Advanced Biosensors for Health Care Applications 0128157437, 9780128157435

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ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

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ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS Edited by

Inamuddin Chemistry Department, King Abdulaziz University, Jeddah, Saudi Arabia

Raju Khan CSIR-North East Institute of Science and Technology, Jorhat, India

Ali Mohammad Department of Applied Chemistry, Aligarh Muslim University, Aligarh, India

Abdullah M. Asiri Chemistry Department, King Abdulaziz University, Jeddah, Saudi Arabia

Elsevier Radarweg 29, PO Box 211, 1000 AE Amsterdam, Netherlands The Boulevard, Langford Lane, Kidlington, Oxford OX5 1GB, United Kingdom 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States Copyright © 2019 Elsevier Inc. All rights reserved. No part of this publication may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, recording, or any information storage and retrieval system, without permission in writing from the publisher. Details on how to seek permission, further information about the Publisher’s permissions policies and our arrangements with organizations such as the Copyright Clearance Center and the Copyright Licensing Agency, can be found at our website: www.elsevier.com/ permissions. This book and the individual contributions contained in it are protected under copyright by the Publisher (other than as may be noted herein). Notices Knowledge and best practice in this field are constantly changing. As new research and experience broaden our understanding, changes in research methods, professional practices, or medical treatment may become necessary. Practitioners and researchers must always rely on their own experience and knowledge in evaluating and using any information, methods, compounds, or experiments described herein. In using such information or methods they should be mindful of their own safety and the safety of others, including parties for whom they have a professional responsibility. To the fullest extent of the law, neither the Publisher nor the authors, contributors, or editors, assume any liability for any injury and/or damage to persons or property as a matter of products liability, negligence or otherwise, or from any use or operation of any methods, products, instructions, or ideas contained in the material herein. British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress ISBN: 978-0-12-815743-5 For Information on all Elsevier publications visit our website at https://www.elsevier.com/books-and-journals

Publisher: Susan Dennis Acquisition Editor: Kostas Marinakis Editorial Project Manager: Theresa Yannetty Production Project Manager: Vijayaraj Purushothaman Cover Designer: Mark Rogers Typeset by MPS Limited, Chennai, India

Contents List of Contributors ix Preface xiii

4. Electrochemical Biosensors for Antioxidants ´ N, DAVID LO ´ PEZ-IGLESIAS, JUAN JOSE´ GARCI´A-GUZMA MARIANA MARIN, CECILIA LETE, STELIAN LUPU, JOSE´ MARI´A PALACIOS-SANTANDER AND LAURA CUBILLANA-AGUILERA

1. Advanced Nanoparticle-Based Biosensors for Diagnosing Foodborne Pathogens

4.1 Introduction 105 4.2 Biosensors for the Determination of Reactive Oxygen Species 108 4.3 Electrochemical Biosensors for the Assessment of Total Antioxidant Capacity of Plants and Foods 115 4.4 Biosensors for the Analysis of Polyphenols in Beverages 120 4.5 Conclusion 135 Acknowledgments 136 References 136

MOHAMMAD LUKMAN YAHAYA, RAHMAH NOORDIN AND KHAIRUNISAK ABDUL RAZAK

1.1 Introduction 1 1.2 Types of Bioreceptors 6 1.3 Types of Transducers 10 1.4 Membrane-Based Biosensors 22 1.5 Multiplex Biosensors 26 1.6 Performance of Nanoparticle-Based Biosensors 1.7 Conclusion and Future Perspectives 33 Acknowledgments 34 References 34

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5. Electrochemical Immunosensors for Rapid Detection of Breast Cancer Biomarkers

2. Aptamer Technology for the Detection of Foodborne Pathogens and Toxins

ANKITA SINHA, DHANJAI, SAMUEL M. MUGO, HUIMIN ZHAO, JIPING CHEN AND RAJEEV JAIN

ALOK KUMAR, MADHU MALINEE, ABHIJEET DHIMAN, AMIT KUMAR AND TARUN KUMAR SHARMA

2.1 Introduction 45 2.2 Detection of Foodborne Pathogens 50 2.3 Biosensors 53 2.4 Classification of Aptamer-Based Biosensors 2.5 Conclusion 64 Competing Financial Interest 65 References 65

5.1 Introduction 147 5.2 Electrochemical Immunosensing of Breast Cancer Protein Biomarkers 151 5.3 Future Prospects and Challenges 165 5.4 Conclusion 166 Acknowledgment 166 References 166

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6. Functionalized Advanced Hybrid Materials for Biosensing Applications

3. Biosensors for Rapid Detection of Breast Cancer Biomarkers

OANA HOSU, ANCA FLOREA, CECILIA CRISTEA AND ROBERT SANDULESCU

AC PEREIRA, MGF SALES AND LR RODRIGUES

3.1 Introduction 71 3.2 Biosensors 79 3.3 Biosensors for Rapid Breast Cancer Detection 3.4 Conclusion 98 References 98

6.1 Introduction 171 6.2 Advanced Inorganic Hybrid Materials 172 6.3 Advanced Organic
Inorganic Hybrid Materials 176 6.4 Advanced Organic Hybrid Materials 188

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CONTENTS

10.3 Drawbacks of Monoclonal Antibodies 265 10.4 Nanobody Characteristics Making Them Suitable for Therapeutic Application 265 10.5 Nanobodies Binding to Transmembrane Proteins 266 10.6 Application of Nanobodies in Medical Imaging 267 10.7 Inflammatory Diseases 268 10.8 Chronic Respiratory Diseases 269 10.9 Application of Nanobodies Against Viruses 269 10.10 Application of Nanobodies Against Bacteria 272 10.11 Conclusion 274 References 274

6.5 Optical Multifunctional Advanced Hybrid Materials 190 6.6 Conclusion 196 References 197

7. Smart, Portable, and Noninvasive Diagnostic Biosensors for Healthcare SRINIVASULU KANAPARTHI, PATTA SUPRAJA AND SHIV GOVIND SINGH

7.1 Introduction 209 7.2 Wearable Sweat Sensors 210 7.3 Gas Sensors for Healthcare 217 7.4 Future Perspectives 224 7.5 Conclusion 225 References 225

11. New Micro- and Nanotechnologies for Electrochemical Biosensor Development

8. Aptamer-Mediated Nanobiosensing for Health Monitoring

NAN HAO, JINWEN LU, RONG HUA, WEI-WEI ZHAO AND KUN WANG

MADHU MALINEE, ALOK KUMAR, ABHIJEET DHIMAN AND TARUN KUMAR SHARMA

8.1 Introduction 227 8.2 Biosensors 228 8.3 Types of Nanobiosensors 232 8.4 Conclusion 244 Competing Financial Interest 245 References 245

9. Biosensing
Drug Delivery Systems for In Vivo Applications

Introduction 279 Microfluidics Chips 280 Quantum Dots 287 Graphene 294 Graphitic Carbon Nitride (g-C3N4) Based Nanomaterials 301 11.6 Conclusion 307 References 308

12. Cholesterol-Based Enzymatic and Nonenzymatic Sensors

M. BIRGUL AKOLPOGLU, UGUR BOZUYUK, PELIN ERKOC AND SEDA KIZILEL

9.1 Introduction 249 9.2 Adaptation of Biosensing
Drug Delivery Systems to In Vivo Applications 250 9.3 Materials and Devices Used as Biosensing
Drug Delivery Systems 254 9.4 Case Studies 257 9.5 Conclusion 259 References 260

10. Nanobodies and Their In Vivo Applications

RAJASEKHAR CHOKKAREDDY, NIRANJAN THONDAVADA, SURENDRA THAKUR AND SUVARDHAN KANCHI

12.1 Introduction 315 12.2 Cholesterol Absorption 317 12.3 Enzymatic Sensors 320 12.4 Nonenzymatic Sensors 331 12.5 Future Prospects 336 12.6 Conclusion 336 References 336

13. Recent Trends in Sensors for Health and Agricultural Applications

PRASHANT SINGH, FANDING GAO AND ANDREA BERNAT

10.1 Introduction 263 10.2 Heavy-Chain-Only Antibodies Generation

11.1 11.2 11.3 11.4 11.5

264

RAJASEKHAR CHOKKAREDDY, NIRANJAN THONDAVADA, SURENDRA THAKUR AND SUVARDHAN KANCHI

13.1 Introduction

341

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CONTENTS

13.2 Smart Health Sensors 343 13.3 Smart Agriculture Sensors 349 13.4 Global Market for Smart Sensors 351 13.5 Future Prospects 352 13.6 Conclusion 353 References 353

14. Hybrid Carbon Nanostructures for Chemical and Biological Sensors PRATIBHA, SUPRIYA SINGH, SUDESH KUMAR, KAKARLA RAGHAVA REDDY, S. NAVEEN AND VEERA SADHU

14.1 Introduction 357 14.2 Synthesis of 3D Structured Graphene Nanosheets 357 14.3 Graphene-Based Electrochemical Sensors and Electrodes for Detecting Biomolecules 362 14.4 Conclusion 368 References 369 Further Reading 374

15. Challenges and Future Prospects of Nanoadvanced Sensing Technology CHRISTINA G. SIONTOROU, GEORGIA-PARASKEVI NIKOLELI, MARIANNA-THALIA NIKOLELIS AND DIMITRIOS P. NIKOLELIS

15.1 Introduction 375 15.2 Nanomaterial-Based Biosensors and Chemical Sensors 376 15.3 Nanostructures, Nanoparticles, Nanowires, Nanofibers, and Nanoprobes 382 15.4 Tubular and Porous Nanostructures 383 15.5 Metal Nanomaterials 383 15.6 Metal Oxide Nanomaterials 385 15.7 Carbon-Based Nanomaterials 386 15.8 Polymer Nanomaterials 388 15.9 Biological Materials 393 15.10 Conclusion 393 References 393

Index 397

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List of Contributors

Chemistry, Dalian Institute of Chemical Physics, Chinese Academy of Sciences, Dalian, P.R. China

Khairunisak Abdul Razak School of Materials and Mineral Resources Engineering, Universiti Sains Malaysia, Penang, Malaysia; Nanobiotechnology Research and Innovation (NanoBRI), Institute for Research in Molecular Medicine, Universiti Sains Malaysia, Penang, Malaysia

Abhijeet Dhiman Department of Biotechnology, All India Institute of Medical Sciences, New Delhi, India; Faculty of Pharmacy, Uttarakhand Technical University (UTU), Dehradun, India

M. Birgul Akolpoglu Chemical and Biological Engineering Department, Koc University, Istanbul, Turkey

Pelin Erkoc College of Engineering and Natural Sciences, Bahcesehir University, Besiktas, Turkey Anca Florea Analytical Chemistry Department, Faculty of Pharmacy, Iuliu Ha¸tieganu University of Medicine and Pharmacy, Cluj-Napoca, Romania

Andrea Bernat Department of Nutrition, Food and Exercise Sciences, Florida State University, Tallahassee, FL, United States Ugur Bozuyuk Chemical Engineering Department, Istanbul, Turkey

and Koc

Fanding Gao Food Science Program, Division of Food Systems and Bioengineering, University of Missouri, Columbia, MO, United States

Biological University,

Juan Jose´ Garcı´a-Guzma´n Institute of Research on Electron Microscopy and Materials (IMEYMAT), Department of Analytical Chemistry, Faculty of Sciences, Campus de Excelencia Internacional del Mar (CEIMAR), University of Cadiz, Campus Universitario de Puerto Real, Polı´gono del Rı´o San Pedro, Puerto Real-Cadiz, Spain

Jiping Chen CAS Key Laboratory of Separation Science for Analytical Chemistry, Dalian Institute of Chemical Physics, Chinese Academy of Sciences, Dalian, P.R. China Rajasekhar Chokkareddy Department of Chemistry, Durban University of Technology, Durban, South Africa

Nan Hao Key Laboratory of Modern Agriculture Equipment and Technology, School of Chemistry and Chemical Engineering, Jiangsu University, Zhenjiang, P.R. China; Key Laboratory of Analytical Chemistry for Life Science and Collaborative Innovation Center of Chemistry for Life Science, School of Chemistry and Chemical Engineering, Nanjing University, Nanjing, P.R. China

Cecilia Cristea Analytical Chemistry Department, Faculty of Pharmacy, Iuliu Ha¸tieganu University of Medicine and Pharmacy, Cluj-Napoca, Romania Laura Cubillana-Aguilera Institute of Research on Electron Microscopy and Materials (IMEYMAT), Department of Analytical Chemistry, Faculty of Sciences, Campus de Excelencia Internacional del Mar (CEIMAR), University of Cadiz, Campus Universitario de Puerto Real, Polı´gono del Rı´o San Pedro, Puerto Real-Cadiz, Spain

Oana Hosu Analytical Chemistry Department, Faculty of Pharmacy, Iuliu Ha¸tieganu University of Medicine and Pharmacy, Cluj-Napoca, Romania Rong Hua Key Laboratory of Modern Agriculture Equipment and Technology, School of Chemistry and Chemical Engineering, Jiangsu University, Zhenjiang, P.R. China

Dhanjai Department of Mathematical and Physical Sciences, Concordia University of Edmonton, Edmonton, AB, Canada; CAS Key Laboratory of Separation Science for Analytical

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LIST OF CONTRIBUTORS

Naveen School of Basic University, Bangalore, India

Sciences,

Jain

Rajeev Jain School of Studies in Chemistry, Jiwaji University, Gwalior, India

S.

Srinivasulu Kanaparthi Department of Electrical Engineering, Indian Institute of Technology Hyderabad, Kandi, India

Georgia-Paraskevi Nikoleli Laboratory of Inorganic and Analytical Chemistry, School of Chemical Engineering, Department 1, Chemical Sciences, National Technical University of Athens, Athens, Greece

Suvardhan Kanchi Department of Chemistry, Durban University of Technology, Durban, South Africa Seda Kizilel Chemical and Biological Engineering Department, Koc University, Istanbul, Turkey

Dimitrios P. Nikolelis Laboratory of Environmental Chemistry, Department of Chemistry, University of Athens, Athens, Greece

Alok Kumar Department of Immunology and Genomic Medicine, Graduate School of Medicine, Kyoto University, Kyoto, Japan

Marianna-Thalia Nikolelis Laboratory of Environmental Chemistry, Department of Chemistry, University of Athens, Athens, Greece

Amit Kumar Discipline of Biosciences and Biomedical Engineering, Indian Institute of Technology Indore, Simrol, Indore, India

Rahmah Noordin Nanobiotechnology Research and Innovation (NanoBRI), Institute for Research in Molecular Medicine, Universiti Sains Malaysia, Penang, Malaysia

Sudesh Kumar Department of Chemistry, Banasthali University, Banasthali Vidyapith, Vanasthali, India Cecilia Lete Institute of Physical Chemistry “Ilie Murgulescu” of the Romanian Academy, Bucharest, Romania David Lo´pez-Iglesias Institute of Research on Electron Microscopy and Materials (IMEYMAT), Department of Analytical Chemistry, Faculty of Sciences, Campus de Excelencia Internacional del Mar (CEIMAR), University of Cadiz, Campus Universitario de Puerto Real, Polı´gono del Rı´o San Pedro, Puerto Real-Cadiz, Spain Jinwen Lu Key Laboratory of Modern Agriculture Equipment and Technology, School of Chemistry and Chemical Engineering, Jiangsu University, Zhenjiang, P.R. China Stelian Lupu Department of Analytical Chemistry and Environmental Engineering, Faculty of Applied Chemistry and Materials Science, University Politehnica of Bucharest, Bucharest, Romania Madhu Malinee Department of Anatomy and Developmental Biology, Graduate School of Medicine, Kyoto University, Kyoto, Japan Mariana Marin Institute of Physical Chemistry “Ilie Murgulescu” of the Romanian Academy, Bucharest, Romania Samuel M. Mugo Department of Physical Sciences, MacEwan University, Edmonton, AB, Canada

Jose´ Marı´a Palacios-Santander Institute of Research on Electron Microscopy and Materials (IMEYMAT), Department of Analytical Chemistry, Faculty of Sciences, Campus de Excelencia Internacional del Mar (CEIMAR), University of Cadiz, Campus Universitario de Puerto Real, Polı´gono del Rı´o San Pedro, Puerto Real-Cadiz, Spain AC Pereira Centre of Biological Engineering, Minho University, Braga, Portugal Pratibha Department of Chemistry, Banasthali University, Banasthali Vidyapith, Vanasthali, India Kakarla Raghava Reddy School of Chemical & Biomolecular Engineering, The University of Sydney, Sydney, NSW, Australia LR Rodrigues Centre of Biological Engineering, Minho University, Braga, Portugal Veera Sadhu School of Physical Sciences, Banasthali University, Banasthali Vidyapith, Vanasthali, India MGF Sales Centre of Biological Engineering, Minho University, Braga, Portugal; BIOMARK/ ISEP, Superior Institute of Engineering of Porto, Porto, Portugal Robert Sandulescu Analytical Chemistry Department, Faculty of Pharmacy, Iuliu Ha¸tieganu University of Medicine and Pharmacy, Cluj-Napoca, Romania

LIST OF CONTRIBUTORS

Tarun Kumar Sharma Center for Bio-design and Diagnostics, Translational Health Science and Technology Institute, NCR Biotech Science Cluster, Faridabad, India Prashant Singh Department of Nutrition, Food and Exercise Sciences, Florida State University, Tallahassee, FL, United States Shiv Govind Singh Department of Electrical Engineering, Indian Institute of Technology Hyderabad, Kandi, India Supriya Singh Department of Chemistry, Banasthali University, Banasthali Vidyapith, Vanasthali, India Ankita Sinha Key Laboratory of Industrial Ecology and Environmental Engineering (Ministry of Education, China), School of Environmental Science and Technology, Dalian University of Technology, Dalian, P.R. China Christina G. Siontorou Laboratory of Simulation of Industrial Processes, Department of Industrial Management and Technology, School of Maritime and Industry, University of Piraeus, Piraeus, Greece Patta Supraja Department of Electrical Engineering, Indian Institute of Technology Hyderabad, Kandi, India

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Surendra Thakur eSkills CoLab, Durban University of Technology, Durban, South Africa Niranjan Thondavada Department of Chemistry, Durban University of Technology, Durban, South Africa Kun Wang Key Laboratory of Modern Agriculture Equipment and Technology, School of Chemistry and Chemical Engineering, Jiangsu University, Zhenjiang, P.R. China Mohammad Lukman Yahaya School of Materials and Mineral Resources Engineering, Universiti Sains Malaysia, Penang, Malaysia Huimin Zhao Key Laboratory of Industrial Ecology and Environmental Engineering (Ministry of Education, China), School of Environmental Science and Technology, Dalian University of Technology, Dalian, P.R. China Wei-Wei Zhao Key Laboratory of Analytical Chemistry for Life Science and Collaborative Innovation Center of Chemistry for Life Science, School of Chemistry and Chemical Engineering, Nanjing University, Nanjing, P.R. China

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Preface

Advanced materials nanotechnology is a promising and developing field that utilizes nanoparticles to encourage the treatment and/ or diagnosis of different diseases, for example, cancer, diabetes, osteoarthritis, cerebrum, retinal and cardiovascular diseases, and bacterial infections. Advanced nanosensor technology improves everyday lifestyles via specially personalized health diagnostic and monitoring systems. Research communities from different fields including material science, physics, chemistry, and biology have met up to develop more modern, sophisticated, practically viable, and realistic diagnostic and monitoring devices. There are several diagnostic methods available to manage these devise’s major requirements such as sensitivity and selectivity. However, biosensors are important sensing tools that have received much interest due to their high sensitivity and selectivity and have inspired new thoughts for creating novel sensing materials, hardware gadgets, scaled-down transduction techniques, and gadget integration to create innovative future diagnostic systems. A biosensor is a small device designed to detect a chemical or biochemical target molecule. It is widely used as a powerful analytical tool in medical research, clinical diagnosis, environmental testing, agriculture, food quality control, bioprocess monitoring, and development of new pharmaceuticals, among others. This book is an endeavor to portray the requirement for novel procedures for building up another class of assay system to retrieve the desired health information of the patient in real

time. This book explores the potential of multidisciplinary science to design and develop smart-sensing technology using micro or nanoelectrodes, novel sensing materials, integration with microelectromechanical systems (MEMS), miniaturized transduction systems, novel sensing strategies such as field-effect transistor (FET), complementary metal–oxide– semiconductor (CMOS), system-on-a-chip (SoC), diagnostic-on-a-chip (DoC), and lab-ona-chip (LOC) for diagnostics and personalized healthcare monitoring. These ongoing advancements on the nanoscale biomedical determination gadgets portrayed in this book will enable us to introduce these innovations into the market in the near future. Therefore we have highlighted different types of prognostic and diagnostic biomarkers associated with cancer, diabetes, Alzheimer’s disease, brain, retinal, and cardiovascular diseases, bacterial infections, and various types of electrochemical biosensor techniques used for early detection of the potential biomarkers of specific diseases. There are 15 chapters written by the world’s foremost authors in this multidisciplinary subject. Some of the contributing authors are from India, Malaysia, Japan, Portugal, Greece, Spain, China, South Africa, Canada, Romania, Turkey, Germany, the United States, South Korea, Australia, and others, and are associated with academia, government, and industry. This book is planned with wide scope of gathering of people having an engineering or science background. This book will be valuable as a reference to researchers and engineers already

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PREFACE

experienced in the field or as a primer to researchers and graduate students just getting started in the art and science of electromechanical miniaturization. The applications of biosensors in the areas such as detection of foodborne pathogens, aptamer biosensor, genobiosensor, metal-based biomarker detection, glucose biosensor and breast cancer biomarkers, are discussed in details. The book is arranged according to these particular themes: Chapter 1 discusses recent advances in biosensors for detecting foodborne pathogens found in food, beverages, and clinical samples. Chapter 2 summarizes the various kinds of aptamer-based biosensors working on different transduction principles (e.g., optical, electrochemical, piezoelectrical, acoustic and cantilever sensors) for the detection of foodborne pathogens and their toxins. Chapter 3 aims to provide insights into the current developments on breast cancer and the role of biosensors in its rapid detection, discussing the disease’s epidemiology, types, biomarkers, and current diagnosis. This chapter provides an overview of the different types of biosensors and their applicability in breast cancer detection. Chapter 4 reviews the design and the use of electrochemical biosensors as diagnostic and decision-making tools in selected healthcare applications like the monitoring and quantification of various reactive oxygen species, assessment of the antioxidant capacity of plants and foods, and the electroanalysis of polyphenols in beverages. A discussion on these topics and the relationships between antioxidants and healthcare is also provided. Chapter 5 discusses recent advances in the development of electrochemical immunosensors for quantification of clinically relevant protein biomarkers of breast cancer tumors. Details regarding multilabeled detection antibodies (Ab2) coupled with specific capture antibodies (Ab1) at sensor surfaces and electrochemical screening of biomarkers as antigens

are discussed. Use of nanostructured surfaces, nanoparticle labels, enzyme labels, and magnetic beads are emphasized as signal amplification strategies. Possible opportunities for further improvement in the area of electrochemical immunoassays toward cancer diagnostics are also emphasized. Chapter 6 discusses multifunctional, advanced, hybrid materials applied in optical biosensors design and development emphasizing the sensing methods based on fluorescence, chemiluminescence, electrochemiluminescence, photoelectrochemical materials, luminescent optical labels, surface plasmon resonance, and surface-enhanced Raman scattering applications. Chapter 7 reports a short review of the current research in noninvasive sweat and gas sensors for healthcare. Recent studies on the detection of various sweat biomarkers such as glucose, lactose, and metal ions using wearable sensors are explored. Respiration sensing using humidity in breath to monitor sleep disorders as well as cardiovascular and pulmonary diseases are discussed. Selective VOC sensors and an array of sensors to detect specific diseases and discrimination of multiple diseases are reviewed. Finally, the use of ingestible sensors, an emerging and convenient technology, is described to diagnose various diseases using gas profiles in the gut. Further, the challenges and future scope of these diagnosing techniques are discussed. Chapter 8 discusses the development of aptamer-based biosensor technologies that can surpass the conventional in vitro diagnostic for disease detection and health monitoring. Chapter 9 discusses biosensing
drug delivery systems for in vivo applications. Different types of materials, devices, sensors, and case studies are discussed. Chapter 10 deals with nanobodies and their in vivo applications. Nanobody characteristics and their binding to transmembrane proteins are highlighted. The applications of nanobodies in areas such as medical imaging,

PREFACE

inflammatory diseases, and chronic respiratory diseases are discussed. The applications of nanobodies against viruses and bacteria are also highlighted. Finally, challenges with nanobodies and future directions are reviewed. Chapter 11 presents recent advances of micro and nanotechnologies for electrochemical biosensor development, including microfluidics, graphene, graphitic carbon nitride, and quantum dots. The applications of quantum dots, graphene, and graphitic carbon nitride in electrochemical biosensors, such as enzyme, DNA, and immunosensing are summarized. The combination of microfluidics and electrochemical sensing in recent years are also introduced. Chapter 12 highlights the cholesterolenabled enzymatic and nonenzymatic sensors to understand the complex interface among cholesterol and bile acids absorptions. The main focus of this chapter is to briefly discuss the advances of cholesterol-based sensors and their applications. Chapter 13 explores sensors for health and agricultural applications. The global market scenario for smart sensors and future prospects of sensors are also highlighted. Chapter 14 reviews the recent progress in synthesis strategies and physicochemical properties of 3D graphene, and its hybrid materials functionalized with different inorganic metallic particles together with its applications for the

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rapid detection of various chemicals and biomolecules at low concentrations. The sensing mechanism and recyclability of sensing performance of graphene-based hybrid electrodes are also discussed. Chapter 15 reviews the principles of nanosensors by describing their classification, principle of operation, and examples of various nanosensors. It also provides directions for research in this field and justifies their practical implementation in sensing by highlighting current challenges and future prospects. This book is the consequence of the commendable cooperation of authors from various interdisciplinary fields of science. It thoroughly examines the most generous, start-to-finish, and forefront research and reviews. We are thankful to all the contributing authors and their coauthors for their commitment. We also thank all copyright holders, authors, and other individuals who agreed for us to use their figures, tables, and schemes. Although every effort has been made to secure copyright permissions from the individual copyright holders to use figures/tables/schemes we should need to offer our sincere proclamations of disappointment to them if, unintentionally, their benefit is being infringed. Editors

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C H A P T E R

1 Advanced Nanoparticle-Based Biosensors for Diagnosing Foodborne Pathogens Mohammad Lukman Yahaya1, Rahmah Noordin2 and Khairunisak Abdul Razak1,2,* 1

School of Materials and Mineral Resources Engineering, Universiti Sains Malaysia, Penang, Malaysia, 2 Nanobiotechnology Research and Innovation (NanoBRI), Institute for Research in Molecular Medicine, Universiti Sains Malaysia, Penang, Malaysia

1.1 INTRODUCTION

minimally processed ready-to-eat meats, dairy products, or fruits and vegetables [1]. The presence of pathogens in ready-to-eat products is a serious concern because these products generally do not receive any further treatment before being eaten. Animals and poultry are the most significant reservoir for many foodborne pathogens [2], although animal by-products, such as feed supplements, may also transmit pathogens to other animals. Seafood is another potential source of pathogens such as Vibrio, Listeria, Yersinia, Salmonella, Shigella, Clostridium, Campylobacter, and the hepatitis A virus [3]. One of the common foodborne pathogens is Escherichia coli. Pathogenic E. coli is composed of six main groups namely, enterohemorrhagic (EHEC), enterotoxigenic (ETEC), enteropathogenic (EPEC), enteroinvasive (EIEC),

The demand for pathogen-free food and beverages is increasing nowadays. Thus concern for disease transmission, which increases the potential for foodborne outbreaks and other associated health issues, has risen across boundaries. To solve this problem, all food and beverages must be screened to ensure that they are free from potential pathogens before entering the market. Some of these pathogens and their screening methods are discussed in the next section.

1.1.1 Typical Foodborne Pathogens Bacterial pathogens have caused human illnesses over the past few decades through the consumption of undercooked or

*

Corresponding authors

Advanced Biosensors for Health Care Applications DOI: https://doi.org/10.1016/B978-0-12-815743-5.00001-9

1

© 2019 Elsevier Inc. All rights reserved.

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1. ADVANCED NANOPARTICLE-BASED BIOSENSORS FOR DIAGNOSING FOODBORNE PATHOGENS

enteroaggregative (EAEC), and diffusely adherent E. coli (DAEC) [4]. Among them, E. coli of serogroup O157:H7 (E. coli O157:H7) is most frequently found to cause foodborne diseases that come from the EHEC group [5]. All members of EHEC, including E. coli O157: H7, can produce Shiga toxins (Stxs) that can cause hemorrhagic colitis, diarrhea associated with abdominal cramps, presentation of mild fever, thrombotic thrombocytopenic purpura, and life-threatening hemolytic uremic syndrome (HUS) in humans [5,6]. The United States Department of Agriculture and Food Safety and Inspection Service (USDA-FSIS) limits the detection of E. coli O157:H7 to one colony-forming unit (CFU) per 65 g of a sample of meat [7]. Salmonella spp. are Gram-negative, rodshaped bacteria, and common foodborne pathogens. Salmonella enterica and Salmonella bongori are two species under the Salmonella genus of the Enterobacteriaceae family. This genus can be further divided into many serotypes, and all of the strains are potentially pathogenic to humans [8]. Salmonella spp. are usually found in dairy and farm products. Salmonella spp. can cause salmonellosis with symptoms of diarrhea, fever, and abdominal cramps. Furthermore, salmonellosis that infects other parts of the body other than the intestines can result in life-threatening and fatal infections [7,8]. However, Salmonella enteritidis and S. typhimurium, as well as S. enterica serovar Typhi, are epidemiologically important as the highest causative agents in human infections worldwide [7,9]. Another genus of the Enterobacteriaceae family that is responsible for foodborne disease is Shigella spp. This genus is composed of four species, namely, Shigella dysenteriae, Shigella flexneri, Shigella sonnei, and Shigella boydii. Shigella spp. cannot be detected in animals because it only infects humans and other primates [10]. Disease transmission mainly occurs via food contaminated with human

feces. When an individual becomes infected, Shigella spp. can invade the intestinal cells, causing inflammation and tissue damage [10,11]. Clinical manifestations vary among species. S. dysenteriae causes dysentery with complications, such as HUS, S. flexneri and S. boydii also cause dysentery but without these complications, whereas S. sonnei causes watery diarrhea [10]. Vibrio spp. are also responsible for foodborne diseases. The three species that commonly infect humans are Vibrio cholerae, Vibrio parahaemolyticus, and Vibrio vulnificus. Vibrio spp. usually contaminate water and seafood and transmit to humans via consumption. Ingestion of contaminated food with Vibrio spp. can cause gastroenteritis and septicemia [12,13]. Meanwhile, Campylobacter spp. are Gram-negative, spiral, and microaerophilic bacteria. A species that is clinically noteworthy as a human enteropathogen is Campylobacter jejuni, which infects poultry, dairy products, and milk. In addition, C. jejuni can attack the peripheral nervous system, causing partial paralysis [7,14]. Other foodborne bacteria that may infect humans are Gram-positive bacteria such as Bacillus cereus, Clostridium botulinum, Clostridium perfringens, and Staphylococcus aureus, and Gramnegative bacteria such as Arcobacter, Helicobacter, and Yersinia enterocolitica. Furthermore, bacterial toxins from Fusarium and fungi such as Aspergillus are also causative agents. Some viruses, such as the hepatitis A virus, and parasites also play a significant role in foodborne diseases [15,16].

1.1.2 Established and Traditional Diagnostic Tools Established and traditional methods for bacterial testing rely on specific media to enumerate and isolate viable bacterial cells in food. These methods are extremely sensitive and

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

1.1 INTRODUCTION

inexpensive and offer both qualitative and quantitative information on the number and nature of microorganisms presence in a food sample [17]. Traditional methods for detecting bacteria involve some basic steps, namely pre-enrichment, selective plating, biochemical screening, and serological confirmation [18]. Hence a complete series of test is often required before any identification can be confirmed [19]. These traditional methods require several days to yield results because they rely on the ability of the organism to multiply into visible colonies [20]. Moreover culture medium preparation, inoculation of plates, and colony counting make these methods labor intensive. Traditional methods that are generally regarded as the golden standard often take days to completely identify viable pathogens. Any modification that reduces the analysis time can technically be called a rapid method. Other established methods include molecular methods such as polymerase chain reaction (PCR) and sequencing analysis, enzyme-linked immunosorbent assay (ELISA), immunological techniques, and fluorescence-based assay using organic dye molecules [21]. These methods offer more rapid analyses compared to traditional methods. PCR is a powerful tool that allows species-specific detection of organisms based on nucleic acid amplification by using a small amount of target DNA. Different types of PCR are commonly used for detecting foodborne pathogens; specifically, real-time PCR uses fluorescent dyes to measure the amount of amplified product as amplification progresses [22,23], reverse transcriptase PCR (RT-PCR) is used to reverse-transcribe and amplify RNA to cDNA [24], and multiplex PCR simultaneously detects several pathogens in one reaction [25,26]. Although amplifying DNA derived from pure cultures is easy, problems arise if the sample is complex (such as food, soil, or

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biological waste) because PCR is easily inhibited by numerous substances, including humic acids, fats, and proteins [27]. Moreover PCR must be followed by agarose gel electrophoresis to view the results. This viewing method is tedious and uses hazardous chemicals, such as ethidium bromide. To overcome these disadvantages, a combination of PCR and other methods, such as biosensors is recommended [28]. Immunological techniques [29,30] and fluorescence-based assays involving organic dye molecules [31,32] are other established methods used for foodborne pathogen detection. These techniques are based on antigen
antibody bindings (immunodetection). Several methods have been developed using these techniques, which are commercially available and dependent on antibody type and format, such as ELISA, enzyme immunoassay (EIA), enzyme-linked fluorescent assay (ELFA), and bioluminescent enzyme immunoassay (BEIA) [21]. ELISA—including direct ELISAs, sandwich ELISAs, and competitive ELISAs—is the most common format used for the immunodetection of pathogens [33]. Many of these ELISA methods are available as commercial kits and approved by regulatory agencies. ELISAbased analyses can be directly applied for detecting foodborne pathogens. ELISAs combine the specificity of antibodies and the sensitivity of simple enzyme assays by using antibodies or antigens coupled to an easily assayed enzyme. However, a major drawback of these assay formats is their lengthy analysis time. Typical ELISAs comprise many steps such as blocking, washing, incubation, and substrate development. These steps can take several hours to complete and are understandably problematic in instances where rapid detection is required [34]. Although immunodetection-based methods are highly accurate, sensitive, and error proof,

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they are time- and labor-intensive and require skilled personnel, sophisticated equipment, and do not always provide the required detectability and specificity toward the target. Moreover these techniques require several enrichment steps and subsequent biochemical and serological identification; hence, the cost for analysis is high. Recent advances in immunodetection combine the established and traditional methods with biosensors as ideal tools for detecting foodborne pathogens [35
37]. Basic information on nanoparticles and biosensors is provided next.

1.1.3 Nanoparticles and Biosensors Nanoparticles (NPs) are particles with diameters in the range of 1
100 nm according to the International Organization for Standardization (ISO) [38]. NPs possess several distinct physical and chemical properties that make them promising synthetic scaffolds to create novel chemical and biological detection systems. Over the past few years, nanostructured materials, such as noble metal NPs, quantum dots, and magnetic NPs, have been employed in a broad spectrum of highly innovative approaches for assays of metal ions, small molecules, proteins, and nucleic acid biomarkers. In addition, NPs can be fashioned with a wide range of small organic ligands and large biomacromolecules by using tools and techniques of surface modification. Each of these capabilities has allowed researchers to design novel diagnostic systems that offer significant advantages in terms of sensitivity, selectivity, reliability, and practicality [38,39]. A typical interaction between NPs and pathogens is via antibody
antigen recognition. For bacteria, many surface antigens are available for specific recognition by using antibodyconjugated NPs. NPs, such as gold (AuNPs), silver (AgNPs), quantum dots (QDs; e.g.,

CdSe
ZnS), magnetic beads (Fe3O4), upconverting phosphor NPs (UCPs), dye-doped NPs, fluorescent-silica NPs (FSNPs), and carbon nanotubes (CNTs) of various shapes, including nanoshells, can be used to construct NP-based assays [39]. Table 1.1 compares the most remarkable NPs used in NP-based lateral-flow biosensors [40]. Biosensors involve any analytical device that incorporates biologically recognized molecules (bioreceptors) such as antibodies, aptamers, phages, nucleic acid (DNA), enzymes, or biomimics with a physical or chemical signaling device (transducer). Bioreceptors are the first element in biosensor systems that recognize targeted analyte in samples. The transducing mechanism converts the signal or recognition event from the bioreceptor into an interpretable result, either visually or as electrical signal to display and record data. Types of commonly used transducers in a biosensor system include electrochemical, optical, magnetic, thermometric, piezoelectric, or micromechanical transducers [7,21,33]. The types of bioreceptors and transducers are discussed in detail in Section 1.2. In brief, the principle of biosensors recognizing pathogens in food is simple. The targeted pathogens in the sample are recognized or captured by a specific bioreceptor. This bioreceptor then responds and triggers the transducer, which is usually in contact with the bioreceptor. The transducer then responds to the input change and converts it into signal output. The signal can be visually recognized by the naked eye or be amplified on an electronic display. Finally, the results can be stored and analyzed [21]. Biosensors possess a number of characteristics, including rapid process, sensitive, specific, accurate, near real-time assay, reproducible, robust, and user friendly. Rapidity of the assay in biosensors is important to differentiate biosensors from conventional methods. Any diagnostic tool must have excellent

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TABLE 1.1 Comparison of the Most Remarkable Methods and Materials Used in Lateral-Flow Biosensors Labels

Advantages

AuNPs

• Easy to synthesize and • No application of enhancement techniques, poor modify sensitivity, and poor limit of detection (LOD) compared with other methods • Highly biocompatible and versatile • Intense color • Usually inexpensive

Carbon-based materials

• Strong contrast against • Unspecific adsorption background (black against • Some are large, resulting in a slow response assay white) • Different shapes and characteristics • Inexpensive and stable with time

Nanoparticles/microparticles loaded/modified with dyes

• Inexpensive and available • Difficult to synthesize and modify as commercial products • Requires high quantity of dyes to provide a favorable • Favorable sensitivity and signal LOD

Fluorescence NPs

• High sensitivity • Low LODs

• Unable to perform naked-eye detection • Requires equipment with both excitation light and fluorescence measurement • Materials normally exhibit photobleaching effect and lose intensity in time

QDs

• Small size, fast assay • Strong intensity

• Difficult to synthesize and conjugate

Other fluorescent materials (e.g., UCPs)

• Requires less energy to be • Rare elements are expensive excited than QDs • Big size, slow assay

Electroactive NPs

• Highly sensitive • Requires equipment to produce and translate the signal • Low LOD • Reproducibility problem related to electrodes • Devices easily miniaturize and are affordable

Nanoparticles/microparticles loaded/modified with magnetite

• Usually inexpensive • Possible colorimetric response (multisignal analysis)

sensitivity, specificity, and accuracy to avoid or minimize false-positive and false-negative results that can cause an expensive recall or loss of credibility. Moreover, desirable biosensors must provide reproducible measurements along an extended shelf life to ensure the accuracy of the test. The bioreceptor

Disadvantages

• Requires magnetic detectors • Sensitivity related to the size of the particles, leading to slow assays

and transducer should be biochemically and mechanically stable, robust, and have a long shelf life. Finally, the device should be user friendly and the given results should be easy to interpret. These factors can avoid or reduce the need for high levels of training and skilled personnel [7,21].

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1.2 TYPES OF BIORECEPTORS As mentioned, a bioreceptor is the most important element in biosensors because it is the first element that is in contact with the sample to be analyzed. Bioreceptors can recognize biological material and produce recognition events. Bioreceptors are later recognized by transducers because a transducer itself cannot recognize biological material. This section discusses the main bioreceptors such as antibodies, nucleic acids, enzymes, and aptamers used in biosensor devices.

1.2.1 Antibodies In general, antibodies can be classified according to their production method, namely, polyclonal, monoclonal, and recombinant antibodies [34]. To produce specific antibodies for bacterial pathogens, living hosts are needed. The host can be immunized with bacterial cells, which may or may not be heat treated. Serum from that host is then collected. Polyclonal antibodies derived from different B cells recognize multiple epitopes on the same antigen. Each of these individual antibodies recognizes a unique epitope that is located on that antigen. Polyclonal antibodies are typically raised in rabbits, goats, or sheep hosts and they are frequently selected in immunosensor-based assays for pathogen detection. However, polyclonal antibodies have a high potential for cross-reactivity due to their ability to recognize multiple epitopes. In cases where high specificity is required, monoclonal or recombinant antibodies are more applicable [34]. Monoclonal antibodies represent antibodies from a single antibodyproducing B-cell, so they only bind with one unique epitope. Each individual antibody in a polyclonal mixture is technically a monoclonal antibody. However, this term generally refers to a process by which the actual B-cell is

isolated and fused to an immortal hybridoma cell line, such that numerous identical antibodies can be generated. Recombinant antibodies (rAbs) are monoclonal antibodies generated in vitro using synthetic genes. The main difference is that monoclonal antibodies are produced using hybridoma-based technologies. However, rAbs do not need hybridomas and animals in the production process. rAbs are produced in vitro using synthetic genes. The synthesis process involves recovering antibody genes from source cells and amplifying and cloning the genes into an appropriate phage vector. The vector is then introduced into a host and the expression of an adequate amount of functional antibody is achieved. Some advantages of using rAbs are animal-free technology, increased antibody production and control, and decreased production time and isotype of antibodies, which can easily be converted into a different species, isotype, or subtype [34,41]. Antibodies are usually immobilized on a substrate, either on a detector surface, carrier, or its vicinity, of the biosensor device [33]. Three ways are available for this immobilization process, namely, adsorption on gold, avidin
biotin system, and self-assembled monolayers (SAMs). In the adsorption process, a clean gold (Au) surface is immersed in an antibody solution and washed. A sample can then be added and the corresponding pathogen is detected through the antigen
antibody reaction. A drawback of this technique is that the orientation of the antibody binding site is uncontrollable because of the random attachment of antibodies on the Au substrate [21,33]. Moreover, the substrate’s surface can be coated with avidin to bind to biotinylated antibodies to create an avidin
biotin system. This method produces strong bonds that can allow for reuse of sensors through multiple washing [42]. Another immobilized method is the use of SAMs. Immobilization is achieved by forming a SAM when the substrate’s

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surface is immersed in an ethanol solution containing thiols or disulfides. The antibody is immobilized through 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide/N-hydroxysuccinimide (EDC/NHS) linker to the thiol end. Unattached antibodies are washed and the remaining site is blocked [33,43]. Antibodies can also be labeled by enzymes, biotins, fluorescent substances, or radioactive isotopes [21]. Antibodies are labeled to antigens in biosensor devices through two methods: direct and indirect. The direct method directly detects labeled antibodies to the targeted antigen. This method is convenient and inexpensive because it uses only one type of antibody. However, the sensitivity of the direct method is limited due to the insufficient labeling signal of the limited antibody. Meanwhile, the indirect method involves two types of antibodies, making it more expensive than the direct method [21,34]. The first antibody is unlabeled and used to capture a specific antigen. The second antibody is labeled and added to bind with the first antibody. Given that the indirect method involves two types of antibodies, signal generation is enhanced and sensitivity is increased.

1.2.2 Enzymes Other bioreceptors that have been used to couple with a transducer are enzymes that are target specific. The catalytic enzyme with a suitable substrate produces electrons and then transfers to a transducer. Most of the enzymes are proteins with the exception of RNA enzymes [21]. For most applications, an enzyme is used as a label rather than as a bioreceptor alone. Enzymes can be conjugated with antibodies or aptamers. Enzymes as a label can increase sensitivity of antibodies or aptamers because the catalytic activity enhances the number of signals that can be detected by a transducer in

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biosensor devices. Examples of enzymes that are commonly used as labels are alkaline phosphatase, horseradish peroxidase (HRP), and beta-galactosidase. These enzymes are mostly applied in ELISA [21]. The advantages of using an enzyme as a label are high sensitivity, possible direct visualization, stability over time, and safety. Direct visualization can be achieved and the sensitivity for detection increases when combined with AuNPs because enzymes provide a colorimetric signal that further enhances visibility of AuNP in a lateral-flow assay [44
46]. Other labels such as fluorescent tags are unstable particularly when exposed to direct light, while the use of a radioisotope label can be a health hazard [21]. Details of the enzymatic reaction used in developing biosensors are discussed in the following section.

1.2.3 Nucleic acids DNA is a genetic component that is unique for each organism. Given its unique sequence, nucleic acid bioreceptors can be designed to complementarily match with the DNA of the interested organism. These matching pairs can be detected by triggering the transducer of a biosensor. Nucleic acid, bioreceptor-based biosensors offer rapid processing and high specificity, and they are simple and inexpensive. They also match only with complementary DNA of the interested organism. However, DNA damage or mutation in an organism results in changes to the DNA structure and DNA replication is disrupted. This change can alter matching complementary sequences between nucleic acid bioreceptors and the targeted DNA. Thus sensitivity and specificity of the biosensor decrease [21]. Genosensor platforms are applied to develop nucleic acid
based biosensors. Biosensors were developed for detecting V. cholera [37] in which the dry reagent
based nucleic acid

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amplification assay was combined with a portable electrochemical genosensor. Fluorescein-tagged amplification reagenttargeted DNA of V. cholera caused corresponding single-stranded DNA (ssDNA) amplicons to be generated. These fluorescein-labeled ssDNA amplicons were then analyzed using capture probe-modified, screen-printed, gold, electrode biosensors produced via the SAM method. Antifluorescein-conjugated alkaline phosphatase produced electroactive α-naphthol through a catalytic reaction to form enzymatic amplification of the hybridization event. The developed biosensor could detect as low as 10 CFU/mL with 100% sensitivity and specificity using 168 spiked stool samples. Other research groups developed excellent and efficient nucleic acid
based biosensors [28,39,47
49]. However, most of their work involved amplification to tag labels with the DNA of the interested organism. This approach requires pre-treatment of samples before they can be applied in the biosensor device. To overcome this limitation, Paniel et al. developed genosensors without amplification [50]. The screen-printed carbon electrode was immobilized with ssDNA, which was hybridized with a complementary DNA sample. Another ssDNA signal probe labeled with horseradish peroxidase enzyme was added for second hybridization. Unfortunately, LOD of this biosensor platform was 102
103 CFU/mL, which was lower than that of the pre-treatment biosensor by DNA amplification. Amplification multiplies the labeled probe compared with a single labeled probe without an amplification process. Peptide nucleic acid (PNA) is an innovative probe in nucleic acid recognition for DNAbased biosensors. PNA is synthesized by substituting the sugar phosphate backbone of DNA with a pseudopeptide. PNA as a probe molecule has superior hybridization characteristics, detects single-based mismatches, and improves chemical and enzymatic stability.

Moreover, developing label-free detection is possible with different molecular structures. Label-free detection contributes to the establishment of rapid, stable, and reliable biosensor devices. PNA-based nucleic acid biosensors can be used for detecting pathogens and offer more stable and increased affinity toward targeted sequences [51
53]. However, some of the disadvantages of PNA as a recognition element include their high synthesis cost, long aggregation-prone PNA oligomers, poor solubility in aqueous media, and difficulty to purify and characterize [21].

1.2.4 Aptamers Aptamers can also be used as a bioreceptor in biosensor devices. Aptamers are a type of nucleic acid in which single-stranded nucleic acids fold together to form a well-defined 3D structure. The structure makes aptamers function like antibodies with high affinity, and they are only specific to their target molecules. Aptamers can be synthesized in the laboratory either chemically, via enzymatic procedures, or a combination of these methods. Aptamers can be both chemical and biological substances [54]. Systematic evolution of ligands by exponential enrichment (SELEX) is a procedure to engineer aptamers in vitro. This procedure involves several repeated cycles, which begins with selecting randomized ssDNA from the library and incubating the target molecule. After incubation, unbound and bound molecules are separated by partitioning. The target-bound sequences are amplified by PCR to produce an enriched pool as input for the next round of selection [55]. The cycle is usually repeated 8
15 times to obtain highly specific aptamers [56,57]. Fig. 1.1 summarizes the SELEX process. Considering the 3D structures of aptamers, these DNA ligands identify their target structurally and not by their sequence. These

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complex targets. However, the technique involves several additional purification steps to ensure that family ligands are selected against the multiple targets, especially in complex mixtures [58]. SELEX was successfully used to identify S. enteritidis [61] and other pathogens [59].

1.2.5 Other Bioreceptors

FIGURE 1.1 SELEX cycle starts with selecting an aptamer library, followed by binding with the target, partitioning of bound and unbound targets, amplifying the bounded target, and producing an enriched pool of selected aptamers for subsequent cycles.

specific aptamers can recognize target proteins specific to some pathogen. Thus aptamers are suitable as bioreceptors to be combined with electrochemical, optical, or colorimetric transducers for biosensor devices [58]. Targeted molecules of a specific pathogen include some carbohydrates, such as lipopolysaccharides and teichoic acid, cell surface proteins, such as intimin of pathogenic E. coli or PilS of S. enterica type IVB pili, and pore-forming toxins, such as listeriolysin of Listeria monocytogenes [59,60]. Live and whole cells can also be used as a target in SELEX. This technique is suitable to design aptamers for microbial pathogens. Aptamers synthesized by this method offer increasing sensitivity of detection because they target several sites of molecules on the cell surface simultaneously. This method does not require laborintensive steps of isolating and purifying

Other bioreceptors used in biosensors are cellular bioreceptors, biomimetic receptors, and bacteriophages. Biomimetic receptors are artificial or synthetic bioreceptors that are designed and fabricated to mimic antibodies, enzymes, or nucleic acids as bioreceptors. Molecularly imprinted polymer (MIP) is one of the techniques in which a polymer is formed around a molecule as an artificial recognition site [62]. A colorimetric biomimetic, specifically polydiacetylene (PDA), produces visible color changes from blue to red when a bacterial target multiplies [63]. S. enterica, E. coli, and B. cereus are among the pathogens studied by this approach. A cellular bioreceptor makes use of a whole cell or specific cellular component that can specifically bind to certain targets. The cellular bioreceptor can be classified into cellular systems, enzymes, and non-enzymatic proteins [21]. A cellular system is based on a whole cell recognition element. Two phases of transducers occur in this system. The whole cell acts as a transducer in the first phase where it produces a cellular response from the detected analyte. Another transducer is required to convert this response into an electrical signal to be processed and analyzed. A cellular bioreceptor has a low detection limit because of signal amplification, is sensitive to biochemical stimuli, and provides a useful assay for biochemical agents [21,64]. In addition, enzyme protein, as discussed in a previous section, is one of the cellular components that can specifically recognize target analytes. Non-enzymatic proteins, such

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as channel or carrier proteins, on the cellular surface are used to recognize molecules through active or potential sensitive sites. For example, lectins can be used as a bioreceptor for electrochemical transducer to bind with lipopolysaccharides on the cell surface for rapid identification of E. coli subspecies [65] and antibiotics to bind with E. coli and methicillinresistant S. aureus (MRSA) for antibiotic susceptibility testing [66].

1.3 TYPES OF TRANSDUCERS Bioreceptors cannot function alone as a biosensor and must be coupled with a transducer. The transducing mechanism converts the signal or recognition event from the bioreceptor into an electrical signal and is then processed or amplified. The interpretable result can be read either visually or as an electronic display. Transducers, such as electrochemical, optical, or mass-based types, are commonly used in biosensor systems and are discussed next.

1.3.1 Electrochemical Biosensors In an electrochemical-based transducer, the interaction of samples with a bioreceptor produces detectable parameters, such as current, potential different, impedance, and conductance. On the basis of these parameters, electrochemical biosensors can be classified into amperometric, potentiometric, impedimetric, and conductometric types. Some advantages of electrochemical biosensors are favorable sensitivity, measurability in complex and turbid samples, potential for compact design, and low cost [7,21,33]. 1.3.1.1 Amperometric An amperometric biosensor measures the current produced by a catalytic enzyme against an analyte through the redox reaction. The current is directly proportional to the analyte

concentration [33]. To measure current, the applied potential and time must be constant to allow electron transfer from the bioreceptor to transducer [21]. Typically, two electrodes are used for this reaction, namely, working and reference electrodes. Materials that are commonly used as working and reference electrodes are gold (Au), platinum (Pt), graphite, silver (Ag), carbon, and conducting polymers [7]. The working principle of amperometric biosensors for detecting pathogens is simple. The antibody against the target pathogen is immobilized on the working electrode surface. When the target pathogen presence in the sample, it binds to the antibody on the working electrode. This reaction produces a current signal. The secondary antibody against the target that is labeled with the enzyme can be used to enhance the first signal. By doing so, the redox reaction from the enzyme catalysis occurs to generate more electrons. Thus the electrical signal is enhanced and improves the sensitivity of the analysis [7]. Fig. 1.2A shows the working principle of amperometric biosensors. NP-based amperometric biosensors have been recently used to diagnose foodborne pathogens. For example, Altintas et al. [67] developed a fully automated microfluidic-based amperometric biosensor for E. coli detection. They compared the LOD for standard and nanomaterial-amplified immunoassays (IAs) with LOD of 1.99 3 104 and 50 CFU/mL, respectively. A specificity study was also conducted by using Salmonella spp., S. typhimurium, Shigella, and S. aureus and the results confirmed the high specificity of the developed amperometric biosensors. Furthermore, the sensor surface could be regenerated multiple times, significantly reducing the cost of the system. In another study, bionanocompositemodified pencil graphite electrode (PGE) was developed using polypyrrole (PPy)/AuNP/ multiwalled carbon nanotubes (MWCNTs)/ chitosan (Chi) [68]. This hybrid bionanocomposite platform was immobilized with

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FIGURE 1.2 Principle of electrochemical-based biosensor. (A) The schematic principle of an amperometric biosensor. The label (enzyme) accelerates the electrochemically active analyte in the solution to transfer electrons to the electrode. (B) The schematic principle of a potentiometric biosensor. The label changes the ion concentration, such as H1 and K1. The ion accumulates on the gate insulator, which produces a potential difference via current flows through the gate. (C) The schematic principle of an impedimetric biosensor. The impedance is increased because electron transfer is blocked upon the attachment of bacterial cells. Source: Reproduced with permission from Ref. M. Xu, R. Wang, Y. Li, Electrochemical biosensors for rapid detection of Escherichia coli O157:H7, Talanta 162 (2017) 511
522, doi:10.1016/j.talanta.2016.10.050.

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E. coli O157:H7 monoclonal antibodies. The study reported selectivity to E. coli O157:H7 with LOD of B30 CFU/mL in PBS buffer. Li et al. [69] also developed amperometric biosensors for detecting E. coli O157:H7. Another amperometric biosensor was developed for detecting L. monocytogenes [70] and Clostridium tetani [71]. 1.3.1.2 Potentiometric In potentiometric-based biosensors, ions in solution from the bioreceptor are converted to a potential signal. The electrical potential proportional to analyte activity is measured from samples by applying the Nernst relationship: E 5 E0 6

RT In a nF

(1.1)

where E is the potential of measurement, E0 is the standard potential at a 5 1 mol/L, R is the gas constant at 8.314 J mol21 K21, T is the absolute temperature in K, n is the number of electrons that is transferred by electrode reaction, F is the Faraday constant, and a is the total number charges of ion based on the concentration at the anode (1) and cathode (
) [7,72]. From the Nernst relationship, potentiometry yields a logarithmic concentration response so it is highly sensitive toward extremely small concentration changes [7,21]. For potential measurement, the current is maintained near zero to allow for electromotive force (EMF) [21]. Similar to amperometric-based biosensors, two electrodes are used to measure ion potential change. One inert reference electrode and one working electrode must be in contact with the sample to be measured. An enzyme is used as a bioactive element surrounding the probe. The catalytic reaction consumes or produces chemical species [7,64]. It can be in the form of an ion concentration fluctuation or pH change, which generates potential that can be measured and directly read from a pH meter display [73]. Fig. 1.2B shows the working principle of potentiometric biosensors. Some

advantages of potentiometric biosensors include their ability to measure pathogens under in situ conditions, portability, low cost, and high sensitivity. However, selectivity or specificity is poor in some samples. Application of this principle toward pathogen detection is lesser compared with other types of electrochemical biosensors. The two major types of potentiometric biosensors are light-addressable potentiometric sensors (LAPSs) and ion-selective, field-effect transistors (ISFETs) [72]. In studies performed on ISFET-based biosensors, poor LODs and device stability have been encountered in detecting pathogens due to the incompatibility between most ISFET fabrication technology and biomolecular immobilization protocols [7,33,72]. The principle of ISFETs is based on the response of specific ions, such as H1, K1, and Cl2, by applying an electric field to create regions of excess charge to a semiconductor substrate. This process helps by using an ion-selective membrane to cover the gate insulator that only allows selective ions to pass [74]. Meanwhile, LAPS is based on inducing a transient photocurrent to an insulated p- or n-doped semiconductor thin film positioned to be in contact with an electrolyte via the transient illumination of an intensity-modulated light source, such as light-emitting diodes (LEDs). The induced photocurrent magnitude depends on the potential applied to the semiconductor plate [72]. LAPS is more feasible for detecting foodborne pathogens than ISFETs [75]. LAPS devices are also available commercially, such as the Threshold Immunoassay System, for the detection of E. coli O157:H7 in food samples [72]. Ercole et al. also reported the use of LAPs for E. coli detection [76]. Bisha and Brehm used the potentiometric alternating biosensing (PAB) system based on LAPs for detecting E. coli in vegetables [77]. Gehring et al. previously developed an immunoligand assay (ILA) combined with LAPs for the rapid detection of E. coli O157:H7 in buffered saline [78].

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A potentiometric aptamer-based bioreceptor, which couples the DNA nanostructuremodified magnetic beads with a solid-contact polycation-sensitive membrane electrode, was developed for detecting Vibrio alginolyticus [79]. In this study, DNA nanostructuremodified magnetic beads were used to amplify the potential response and eliminate the interference effect in a complex sample matrix. The interaction between V. alginolyticus and the aptamer on the DNA nanostructures induced changes in the charge or DNA concentration on the magnetic beads. Chronopotentiometric detection was conducted on a solid-contact polycation-sensitive membrane electrode by using protamine as an indicator. This potentiometric aptasensing method offered LOD of 10 CFU/mL with acceptable specificity for the detection of V. alginolyticus. Ding et al. reported label-free potentiometric aptasensor for rapid, selective, and sensitive detection of L. monocytogenes [80]. Aptamer as the bioreceptor binds specifically to internalin A present on the surface of L. monocytogenes. A polycation-sensitive membrane electrode is used to detect free protamine when targetbinding events prevent the aptamer from electrostatically interacting with protamine. The LOD of this method is 10 CFU/mL with favorable recovery and high accuracy. Another labelfree potentiometric aptasensor uses a network of single-walled carbon nanotubes (SWCNTs) as an ion-to-electron potentiometric transducer and aptamers as the bioreceptor for detecting S. aureus in real time [81] and living E. coli in complex matrices [82]. Herna´ndez et al. [83] also reported a potentiometric aptasensor based on chemically modified graphene for detecting S. aureus in real time. 1.3.1.3 Impedimetric Another type of electrochemical biosensor is the impedimetric biosensor. Impedance refers to the resistance in an electrical circuit to the flow of alternating current. Hence, it is an

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actual electrical resistance to a direct current [64]. Electrochemical impedance spectroscopy (EIS) is a technique that applies the impedance principle for detecting foodborne pathogens. EIS has been accepted by the Association of Official Analytical Chemists (AOAC) for the detection of Salmonella in food [7,73,84]. In EIS, changes in conductance, capacitance, and impedance are analyzed [73]. Changes in electrical impedance and conductance in a medium are due to microbial metabolism of the inert substrate into an electrically charged ionic compound and acidic by-products such as amino acid, acetic acid, and lactic acid [64]. Changes in electrical impedance also occur when immobilized antibodies on the electrode of the electrochemical cell capture the target [7]. A small amplitude of sinusoidal excitation at various frequencies is applied to measure the change in electrical impedance of the medium [7,73]. The imposed signal involves a range of amplitudes and frequencies, thus, it can be interpreted in two ways. The first is by solving the system for partial differential equations governing the system, which is the most rigorous approach. In the second way, data in terms of equivalent circuits are interpreted, which is the preferred approach because it is simple [7,33]. Fig. 1.2C shows the working principle of impedimetric biosensors. Advantages of impedimetric biosensors include being a label-free method, availability of multiplex detection, automated detection, and favorable selectivity [7,21,73]. Label-free impedimetric biosensors can produce inexpensive electrode sensors. However, their LOD is poor compared to other electrochemical-based biosensors [7,21]. Impedimetric-based biosensors with AuNPmodified, free-standing, graphene paper electrode were used to detect E. coli O157:H7 [85]. In this study, one-step electrodeposition was used to grow AuNP on the surface of the graphene paper electrode. E. coli O157:H7

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antibodies were then immobilized on the paper electrode via a biotin
streptavidin system. Captured E. coli O157:H7 on the paper electrode was detected by EIS. Results showed that this method had low LOD of 1.5 3 102 CFU/mL and excellent specificity. Kim et al. also used an impedimetric biosensor to detect 103 CFU/mL of Salmonella in a pork extract within 5 min of incubation [86]. Mutreja et al. first reported the use of OmpD, a specific surface antigen, as a surface biomarker for developing sensitive and selective immunosensors to detect S. typhimurium species [87]. The generated anti-OmpD antibody was used as the detector probe on graphene
graphene oxide (G-GO)
modified screen-printed carbon electrodes. The change in impedimetric response was then measured to specifically detect S. typhimurium in spiked water and juice samples with LOD of 101 CFU/mL, with high selectivity and extremely low cross-reactivity with other strains. Another study was also performed based on biomarkers, such as lectin of Concanavalin A for detecting E. coli [88] and Internalin B (InlB) for detecting L. monocytogenes [89]. Electrode modification has been studied by a few researchers, such as engineered gold electrode with a biocompatible and hydrophilic layer of hyaluronic acid (HA) [90] and nanoporous membrane
based impedimetric immunosensor [91] for the detection of E. coli O157:H7; magnetically assisted impedimetric [92]; electrochemical immunosensor fabricated using double-layer gold nanoparticles and Chi [36]; graphene nanostructure on the sensor surface [93]; reusable capacitive immunosensor [94]; diazonium-grafting layer [95]; antibody-modified polysilicon interdigitated electrodes [96]; and indium tin oxide (ITO)
based impedimetric [97]. Other strategies include pre-treatment using magnetic bead separation [98,99] and the use of synthetic antimicrobial peptides (sAMPs) as novel recognition agents [100].

1.3.1.4 Conductometric The relationship between conductance and bioreceptor techniques is the basis for developing conductometric detection [21,62] performed using a conducting polymer. The conducting polymer acts as an electrochemical transducer to convert biological signals into electrical signals. Examples of conducting polymers are polyaniline, polythiophene, polyacetylene, and PPy [84]. The basic principle of conductometric detection involves a reaction that can change the ionic species concentration. This reaction leads to changes in electrical conductivity or current flow. In this method, two inert metal electrodes are used. The two electrodes are separated at a certain fixed distance before applying AC voltage, which later causes current flow. The change in conductance between the electrodes due to ionic composition changes during biorecognition is measured [21,62,84]. The speed, practicality, specificity, sensitivity, near real-time detection, and need for low sample volume are some advantages of conductometric biosensors [21,84]. Few applications of conductometric biosensors have been reported for the detection of foodborne pathogens. Pal et al. developed a direct-charge transfer conductometric biosensor to detect Bacillus cereus in numerous food samples [101]. In their study an electronic signal was generated by electron charge flow aided through conductive polyaniline combined with the principle of a sandwich IA. The biosensor could decrease LOD to 35
88 CFU/mL B. cereus in food samples and only took 6 min to complete. The biosensor was selective to B. cereus only. Hnaiein et al. reported the development of IAs based on magnetite NPs for detecting E. coli via conductometric measurements [102]. A biotin
streptavidin interaction system was used to immobilize E. coli antibodies on the modified magnetite NPs. Subsequently, a layer of functionalized NPs was directly immobilized on the conductometric electrode

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via glutaraldehyde cross-linking. Results showed favorable specificity toward E. coli cells and not the other non-specific cells of S. epidermidis. The LOD of 1 CFU/mL of E. coli induced a conductivity variation of 35 μS.

1.3.2 Optical Biosensors Optical biosensors have received considerable attention, especially for bacterial pathogen detection, due to their selectivity and sensitivity. This type of sensor signal can be used because they can be directly amplified for detection. Optical biosensors can also be used in conjunction with other techniques. Other reasons that make optical biosensors more versatile are label-free and real-time detection toward chemical and biological substances, including foodborne pathogens [64]. Subclasses of optical detection are based on reflection, refraction, absorption, dispersion, Raman, infrared (IR), chemiluminescence, phosphorescence, and fluorescence [21]. However, the few major techniques adapting optical detection include surface plasmon resonance, optical fibers, and Raman and Fourier transform infrared spectroscopy. 1.3.2.1 Surface Plasmon Resonance In this method, reflectance spectroscopy or optical illumination of the metal surface is used for pathogen detection [21,73]. Surface plasmon resonance (SPR) is the optical phenomenon of metal occurs from interaction of light towards free electron on the metal surface. The p-polarized light beams from LASER causes resonance condition by exciting electrons to produce an electron density wave called surface plasmon wave (SPW) [7]. This method can also measure the binding kinetics of two molecules without using a label such as a fluorescent tag [64]. A bioreceptor used in this transducer method is an antibody which is immobilized on a thin metal surface, usually gold film, near

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the reflecting surface of waveguide to capture the target pathogen. Strong resonance is generated through the interaction of light with the electron cloud in the metal at a certain wavelength and near-IR region [73]. The result is unique to the corresponding metal and is also called the fingerprint region [103]. When an antibody captures the pathogen on that metal surface, it causes a change in the refractive index (RI) and shifts resonance to a longer wavelength. The change in angle of reflected light is measured and the amount of change corresponds to the concentration of bounded pathogens [21,73]. Fig. 1.3A shows the schematic principle of SPR detection in biosensors. SPR can detect minor changes in RI; thus, it can detect molecules down to the femtomolar range [73]. SPR also offers high selectivity with the aid of a label-free antibody bioreceptor, which can exclude additional assays, reagents, and steps, for multiplexing detection and remote sensing for automation [7,64]. However, the high cost of equipment, large size, and complexity of operation and interpretation are some drawbacks of SPR [33]. Vaisocherova´-Lı´salova´ et al. reported the use of an SPR biosensor based on ultralow fouling and functionalizable poly(carboxybetaine acrylamide) (pCBAA) brushes for the rapid, simultaneous, and sensitive detection of Salmonella sp. and E. coli in hamburger and cucumber samples [104]. This approach utilized a three-step detection assay that captured bacteria by antibodies immobilized to the pCBAA coating. These three steps entailed: (1) incubating the sensor with crude food samples; (2) binding of secondary biotinylated antibody (Ab2) to previously captured bacteria; and (3) binding of streptavidin-coated AuNP to the biotinylated Ab2 to enhance the sensor response. Surface resistance to fouling and the functional capabilities of these brushes were also studied, such as the effect of brush thickness on the biorecognition capabilities of gold-grafted functionalized pCBAA coatings.

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FIGURE 1.3 Schematic view of different optical sensors. (A) Geometry and signal transduction mechanism of SPR (surface plasmon resonance) sensors (reproduced with permission from Ref. D. Shankaran, K. Gobi, N. Miura, Recent advancements in surface plasmon resonance immunosensors for detection of small molecules of biomedical, food and environmental interest, Sens. Act. B Chem. 121 (2007) 158
177, doi:10.1016/j.snb.2006.09.014 [113]) and response of a fiber-optic SPR biosensor to various BSA concentrations. The arrows represent the point of change of the flowing solution (reproduced with permission from R. Slavı´k, J. Homola, E. Brynda, A miniature fiber optic surface plasmon resonance sensor for fast detection of staphylococcal enterotoxin B, Biosens. Bioelectron. 17 (2002) 591
595, doi:10.1016/S0956-5663 (02)00013-1 [114]). (B) Cross section of the portable SPR sensors, Spreeta 2000 with their components, and light path inside the sensor (reproduced with permission from Ref. T. Chinowsky, J. Quinn, D. Bartholomew, R. Kaiser, J. Elkind, Performance of the Spreeta 2000 integrated surface plasmon resonance affinity sensor, Sens. Act. B Chem. 91 (2003) 266
274, doi:10.1016/S0925-4005(03)00113-8 [110]). (C) Geometry of a tapered fiber optic biosensor wherein an antibody is immobilized in the waist region with the largest evanescent field (reproduced with permission from Ref. H. Sharma, R. Mutharasan, Review of biosensors for foodborne pathogens and toxins, Sens. Act. B Chem. 183 (2013) 535
549, doi:10.1016/j.snb.2013.03.137 [7]).

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The LOD determined from the hamburger and cucumber samples were 7.4 3 103 and 11.7 3 103 CFU/mL for Salmonella sp., respectively, and 57 and 17 CFU/mL for E. coli, respectively. The same researchers later found that the antibody-functionalized poly (CBMAA 15 mol%-ran-HPMAA) brush exhibits superior biorecognition properties over pCBAA [105]. Rodriguez-Emmenegger et al. used polymer brushes of poly(2-hydroxyethyl methacrylate) [poly(HEMA)]-based SPR for detecting Cronobacter in samples of consumer powder whole-fat milk preparation, fresh whole-fat milk, and powdered infant formula with LOD of 104 cells/mL [106]. Bacteriophages can also be used as specific binders for SPR-based detection. Karoonuthaisiri et al. analyzed filamentous M13 bacteriophages expressing 12-mer peptides by using a Salmonella-specific bacteriophage to detect Salmonella [107]. Several important factors, such as immobilization buffers and methods, and interaction buffers for a successful bacteriophage-based SPR assay, were optimized. Results showed that Salmonella-specific bacteriophage-based SPR assay had an LOD of 8.0 3 107 and 1.3 3 107 CFU/mL for one-time and five-times immobilized sensors, respectively, with extremely low cross-reactivity toward other non-target bacteria. Tawil et al. also demonstrated the use of bacteriophage-based SPR for detecting E. coli O157:H7 and MRSA [108]. A portable and low-cost commercial SPRbased diagnostic tool, such as Spreeta, was evaluated by Waswa et al. to detect E. coli O157: H7 directly in food samples [109]. Assays were conducted at near
real time and results were obtained after 30 min. LOD of E. coli O157:H7 was 102
103 CFU/mL. This biosensor showed specificity to E. coli O157:H7, while no appreciable change was noted in the sensogram for other organisms such as E. coli K-12 and Shigella. Chinowsky et al. studied the

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performance of Spreeta 2000. Fig. 1.3B shows the components of Spreeta 2000 [110]. 1.3.2.2 Optical Fibers Optical fibers can transmit light based on total internal reflection (TIR). In this biosensor, excitation of laser light is sent by tapered fiber to a detection surface that receives emitted light [21]. Specific recognition molecules, such as antibodies, are immobilized in the waist region of the fiber where the evanescent field is the largest due to its small diameter compared with the upstream fiber. Binding of the specific target analyte to the sensing surface alters RI resulting in altered evanescent field and optical throughput [7]. Fig. 1.3C represents the graphical view of optical fibers. Propagation of light through fiber is extremely sensitive to surrounding interactions, which is a major advantage of this method [21]. Furthermore, the impact of PCR inhibitors present in complex matrices in food samples can also be reduced by fiber-optic biosensors. This method can rapidly detect pathogens from complex matrices, while confirmation tests take up to several hours [21]. Propagating light has two components, namely, guided field in the core and exponentially decreased evanescent field in cladding [7]. Tapered geometry fiber optic is mostly used and can be divided into two types, namely the tapered tip type and continuous tapered fiber type. In the tapered type, optical fiber gradually decreases in diameter to a micron-sized tip to function as the sensing element. The tip is conducive to the collection of response light and launch of the source light for measurement [7]. A continuous biconical taper is composed of three areas, namely convergent area of decreasing diameter, constant diameter area called the waist, and increasing diameter of divergent area. Given the small diameter and presence of maximum evanescent field, the waist area functions as a sensing

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element. Changes in resonance, scattering, absorption, and fluorescence also occur in the waist area. The divergent region used for detection collects the transmitted and scattered light as well as the emission resulting from the fluorescence of sensing [7]. DeMarco and Lim developed a fiber optic biosensor based on an intensity-based evanescent sensor to detect E. coli O157:H7 in seeded 10 and 25 g of ground beef samples [111]. A centrifugation method to obtain samples suitable for biosensor analysis (i.e., 10
25 g of ground beef samples) was developed. They reported that the sensitivity of the assay was high and repeatable. The specificity was 100% at 5.2 3 102 CFU/g for 10 g of ground beef samples with polystyrene waveguides and at 9.0 3 103 CFU/g for 25 g of ground beef samples with silica waveguides. The obtained results demonstrated that the reaction was highly specific. Ko and Grant reported fiber optic portable biosensors utilizing the principle of fluorescence resonance energy transfer (FRET) for the rapid detection of S. typhimurium in ground pork samples [112]. FRET involves non-radiative energy transfer through dipole
dipole interactions in which a fluorescent donor molecule transfers to an acceptor molecule in close proximity. FRET was used with an optical fiber tip sensor. The antibody against Salmonella was labeled with a FRET donor fluorophore (Alexa Fluor 546). Meanwhile, the FRET acceptor fluorophore (Alexa Fluor 594) was labeled with Protein G (PG). When S. typhimurium was bound to a labeled antibody, the conformation of the antibody changed; the distance between the donor and acceptor decreased, thereby increasing fluorescence. The LOD was measured by determining the lowest concentration that generated a remarkable change in signal over the baseline, and it was at 103 cells/mL. The fiber probe worked in detecting S. typhimurium in homogenized pork samples with LOD of 105 CFU/g [7,112].

1.3.2.3 Raman and Fourier Transform Infrared Spectroscopy Raman scattering and Fourier Transform Infrared (FT-IR) are vibrational spectroscopy techniques that provide “fingerprints” of an object or organism [21]. They are computation techniques that involve gathering spectral information based on evaluation and calculation of the radiative source coherence with the help of time or space-domain measurements of the electromagnetic radiation or other types of radiation [64]. FT-IR spectrometry is advantageous because it is a non-destructive analytical technique and allows multiplex detection. However, a pure culture of the pathogen isolated from sample constituents is needed to produce sufficient biomass before the whole organism’s fingerprint can be reported and analyzed [21]. Liu et al. [115] developed a new surfaceenhanced Raman scattering (SERS)
based lateral-flow (LF) strip biosensor combined with recombinase polymerase amplification (RPA) for the simultaneous detection of L. monocytogenes and S. enteritidis [115]. In this SERS-LF, AuMBA@Ag core-shell NPs were used. The characteristic peak intensities of SERS tags were measured for highly sensitive quantitative detection. The LOD for this assay was 27 CFU/ mL for S. enteritidis and 19 CFU/mL for L. monocytogenes with remarkably high specificity and applicability in the detection of S. enteritidis and L. monocytogenes in food samples. In another study, pre-treatment of samples via nanoimmunomagnetic separation (NIMS) combined with SERS was developed to detect E. coli O157 from liquid media of apple juice [116]. During pre-treatment, E. coli O157 cells were separated and concentrated from liquid food matrix via capture antibodies (cAbs) that were immobilized on magnetite-gold (Fe3O4/ Au) magnetic nanoparticles (MNPs). SERS labels were prepared by conjugating AuNPs with Raman reporter molecules and the

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detector antibody (dAb) for detecting the target pathogen. The Au-coated MNP
cAb
E. coli O157 complex from pre-treatment interacted with Au-Raman label-dAb, and it was detected via the Raman method (Fig. 1.4). The LOD from the Raman spectra was 102 CFU/mL and the total analysis time was less than 30 min. No cross-reactivity was observed with background non-targeted organisms. Sundaram et al. [117] developed SERS substrate by depositing a biopolymer encapsulated with silver NPs prepared using a silver nitrate, polyvinyl alcohol (PVA) solution and trisodium citrate on a mica sheet. Fresh cultures of S. typhimurium, E. coli, Listeria innocua, and S. aureus were placed on the substrate individually and exposed to 785 nm HeNe laser excitation. SERS spectral data were recorded over the Raman shift between 400 and 1800 cm21. Different species within the

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Gram-negative and Gram-positive group show almost similar SERS spectra because the Gramnegative bacterial share cell walls, while Gram-positive bacterial share cell membranes. The major differences only can be observed between 1200 and 1700 cm21 for nucleic acid and at fingerprint region between 400 and 700 cm21 for amino acid. Mura et al. used FT-IR spectroscopy as an optical transduction method on mesoporous titania thin-film substrates as sensors to detect low E. coli O157:H7 concentration [118]. In this approach, specific antibodies to E. coli O157: H7 were functionalized to titania films that were treated with (3-aminopropyl)triethoxysilane (APTES) and glutaraldehyde (GA) and the absorbance property was monitored. The film-based biosensors provided LOD of 1 3 102 CFU/mL and were a selective method for the effective screening of water samples.

FIGURE 1.4 Stacked surface-enhanced Raman scattering (SERS) spectra showing E. coli O157:H7 concentrationdependent Raman shift signature of malachite green isothiocyanate (MGITC) as Raman labels reported at vibrational frequency zones of 1180, 1370, and 1620 cm21 for the following conditions: absence of E. coli O157:H7 (spectrum a); 101 CFU/mL (spectrum b); 102 CFU/mL (spectrum c); 103 CFU/mL (spectrum d); 104 CFU/mL (spectrum e); 105 CFU/mL (spectrum f); 106 CFU/mL (spectrum g); and 107 CFU/mL (spectrum h). Source: Reproduced with permission from Ref. R. Najafi, S. Mukherjee, J. Hudson, A. Sharma, P. Banerjee, Development of a rapid capture-cum-detection method for Escherichia coli O157 from apple juice comprising nano-immunomagnetic separation in tandem with surface enhanced Raman scattering, Int. J. Food Microbiol. 189 (2014) 89
97, doi:10.1016/j.ijfoodmicro.2014.07.036 [116].

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1.3.3 Mass-Based Biosensors Mass-based biosensors rely on small changes in crystal mass when they vibrate at a specific frequency after the application of electrical signals of a specific frequency. This mass-sensitive technique is suitable for highly sensitive detection [21]. The two major subtypes of mass-based biosensors are piezoelectric and magnetoelastic sensors. 1.3.3.1 Piezoelectric Piezoelectric is a mass-sensitive detector that uses piezoelectric crystal to generate and transmit vibrations at a distinct frequency with the operation of an electrical signal of explicit frequency [64,73]. The applied electrical frequency to the crystal and mass of the crystal depend on the frequency of oscillation. Thus increased mass caused by chemical binding changes the oscillation frequency of the crystal and the change can be measured electrically to determine the additional mass of the crystal [21]. Commonly used piezoelectric materials are quartz (SiO2) and lithium niobite (LiTaO3) [7]. This method is also known as the bulk wave (BW) or quartz crystal microbalance (QCM) sensor [21]. The working principles of piezoelectric biosensors are very simple given that they measure additional masses of crystals. An antibody is coated to the piezoelectric sensor surface, which is quartz crystal. After placing samples containing the pathogen, a specific pathogen binds to the coated antibody. As a result, the mass of the quartz crystal increases. This event causes the oscillation frequency to decrease and the additional mass of quartz crystal can be measured electrically. This method is relatively practical, simple, and cost effective and is characterized by real-time detection, direct label-free analysis, and increased sensitivity and specificity [7,21,73]. Initially, improvement in the procedure for immobilization of oligonucleotide bioreceptor

on piezoelectric quartz crystal coated with gold was achieved by Tombelli et al. [43]. They found that streptavidin linked to a layer of carboxylated dextran when immobilized with biotinylated probe yields high sensitivity for detecting the hybridization reaction and elevated stability with respect to the regeneration step and lacks non-specific adsorption. The oligonucleotide-based piezoelectric biosensors were further improved by real-time detection of E. coli O157:H7 sequences via a circulating flow system of QCM [119] and oligonucleotide-functionalized AuNPs, which function as a “mass enhancer” and “sequence verifier” to amplify the frequency change of the piezoelectric biosensor to detect E. coli O157:H7 [120]. Apart from oligonucleotides, antibodies can also be used as bioreceptors on piezoelectric biosensors. Li et al. immobilized affinitypurified E. coli O157:H7 antibodies onto SAMs of 3-mercaptopropionic acid (MPA) on the surface of a quartz crystal Au electrode [121]. The LOD of this sensor was 102
105 CFU/mL so it can be applied in detecting E. coli O157:H7 in food samples. Another study enhanced sensitivity on piezoelectric biosensors for detecting E. coli O157:H7. Jiang et al. used amplifiers of micro/nanobeads with different materials (magnetic, polymer, and silica) and sizes (diameter range of 30
970 nm), resulting in a change in bacterial resonance frequency rather than an increase in sensor surface mass [122]. Guo et al. enhanced the sensitivity of the system by pre-enrichment of E. coli O157 in the brain and heart infusion (BHI) broth for 18 h, which provided LOD of 0
1 log CFU/mL [35]. Sensitive detection of Campylobacter jejuni using an AuNP-enhanced QCM sensor was also developed which could detect 150 CFU/mL in solution [123]. Such excellent sensitivity was achieved without pre-enrichment steps; hence, it reduced the assay time. Low cross-reactivity was also found for other

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FIGURE 1.5 Principle of fully automated QCMA-1 device. (A) Normal sandwich format assay with polyclonal antibody as the detection antibody to Campylobacter jejuni on the QCMA-1 sensor. (B) Sandwich assay with enhanced polyclonal antibody
conjugated AuNPs as detection antibodies to C. jejuni on the QCMA-1 sensor. Source: Reproduced with permission from Ref. N.A. Masdor, Z. Altintas, I.E. Tothill, Sensitive detection of Campylobacter jejuni using nanoparticles enhanced QCM sensor, Biosens. Bioelectron. 78 (2016) 328
336, doi:10.1016/j.bios.2015.11.033 [123].

foodborne pathogens, such as S. typhimurium (7%), L. monocytogenes (3%), and E. coli (0%), that demonstrated excellent specificity toward Campylobacter detection. Fig. 1.5 shows the principle of AuNP-enhanced QCM sensors. Real-time and sensitive detection of S. typhimurium was also developed using an automated QCM instrument with NP amplification [124]. In another work, QCM-based sensors were used to detect Salmonella contamination in packaged beef via volatile organic compounds [125]. Detection was performed using a conjugate of synthetic peptide of odorant binding protein (LUSH) to piezoelectric sensors for detecting low concentrations of alcohols (3-methyl-1-butanol and 1-hexanol) at room temperature. 1.3.3.2 Magnetoelastic Another method that applies acoustic wave sensors is the magnetoelastic sensor (ME). Amorphous ferromagnetic ribbons or wires are used to construct ME sensors. Magnetic fluxes are generated when ME sensors are excited with a magnetic AC field and they are detected by a sensing coil. The sensing coil can be at distance as it does not require a direct physical connection. Thus in ME analysis can be performed as wireless monitoring.

Moreover, ME has high tensile strength and is cost effective [21,126]. The operation of ME sensors involves the Joule magnetostriction principle. In this principle, the resonator vibrates longitudinally with a characteristic resonance frequency when subjected to a time-varying magnetic field that changes its size cyclically. Mass changes occur when the target analyte binds to a bioreceptor which is immobilized on the surface of the ME sensor. Elevated mass causes a decrease in the resonance frequency. Thus the change in the sensor resonance frequency can be measured rapidly and accurately [21,126]. Wang et al. [127] reported the use of ME for the detection of S. typhimurium in spinach. Filamentous E2 phage was used as a bioreceptor in the study to specifically bind to S. typhimurium. The effects of alternative pre-enrichment broth, incubation time on the detection performance, and negative control as well as different blocking agents to minimize the effect of nonspecific binding were investigated. The researchers observed all blocking agents, which yielded similar performances, and 7 h of incubation at 37 C was necessary to detect an initial spike of 100 CFU/25 g of S. typhimurium in spinach leaves. Beltrami et al. [126] reported the use of ME with hybrid films for detecting E. coli and

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S. aureus in milk. In their study, magnetoelastic alloy (Metglas 2826MB3) was used as sensor contraction material coated with a hybrid film based on silicon alkoxide precursors (TEOS and MAP) via the sol
gel method. The sensor was not functionalized with a specific bioreceptor. This approach allows response only to mass variation due to adhesion or bacterial deposition. Furthermore, the sol
gel coating method provides excellent resistance to corrosion on the metal substrate due to its barrier properties, strong adhesion, chemical inertness, formulation versatility, and ease of application at room temperature. As a result, this method rapidly detects pathogens within 60 min even without using an antibody as a bioreceptor.

1.4 MEMBRANE-BASED BIOSENSORS Membranes used in membrane-based biosensors can be formed from several materials which is selected based on the final application of that membrane in order to form complete biosensors. The membrane materials are mainly classified into inorganic, organic, hybrid, and composite membranes [128]. Inorganic membranes are used mainly as support structures; they increase the surface area of the sensor and form a capillary action. Examples of inorganic membranes include aluminum-anodized oxide or nanoporous alumina, titanium oxide, gold, silver, silicon nitride, and glass. Organic membranes are more frequently used in biosensors than inorganic membranes. Organic membranes function as a support structure and integral component of the sensing process. Examples of organic membranes are nitrocellulose (NC), polyethersulfone, nylon, polypropylene, polydimethylsiloxane (PDMS), polycarbonate, polylactic acid (PLA), cellulose, cellulose acetate, nanofibers, polyacrylamide, polyvinyl chloride, polyvinylidene fluoride (PVDF), and

polyamine/polyurethane. Meanwhile, hybrid membranes are composed of inorganic and organic materials that are fused together. Examples of hybrid membranes are goldcoated polycarbonate track-etched (PCTE) membranes. Composite membranes are formed by multiple membranes sandwiched together either vertically or side-by-side to form a complete sensor. This combination can be a sample pad of cellulose membrane, conjugate pad from glass, signal-generated pad of NC, or an absorption pad made of a cellulose membrane. Membranes are commonly used in various biomedical applications. They can be used as a platform to allow reactions to occur, such as lateral-flow immunoassay (LFIA). They are also employed as a filtration material to concentrate and isolate cells, bacteria, or viruses for the detection of proteins, DNA, or RNA such as in Western, Southern, or Northern blots, respectively, and other tests (e.g., the direct epifluorescence technique; DEFT). Membranes can also be incorporated into various sensors and biosensors for the detection of different compounds including proteins, DNA and RNA, bacterial cells, and virus particles [128]. In this section, membrane-based biosensors for the detection of pathogens are discussed including LFIA, which is widely used as a platform for pathogen detection. Other types of membranebased biosensors are also discussed.

1.4.1 Lateral Flow Immunoassay LFIA is an IA biosensor based on composite membranes that make use of the NC membrane as their main material along with several parts from different materials. Nylon and PVDF membranes are less commonly used due to their limited success because of their high cost and limited utility [129]. Advantages of LFIA biosensors are rapid process, high

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sensitivity, good specificity, low LOD, less sample operation volume requirement, low manufacturing cost, robustness, no complicated equipment involved, and user-friendly format [40]. Moreover, LFIA can be operated for in-field detection because it utilizes a dry form of reagents incorporated within this device and can be read visually in some assays [130,131]. As a result, LFIA is widely used as a point-of-care device even in low-income countries due to the use of low-cost materials in the manufacturing process. LFIA is manufactured in a rectangular shape that is composed of four parts with different materials. It has a width of 4
6 mm and a length of up to 6
7 cm (Fig. 1.6). Each of the parts has its own function and stick together to form a complete functional LFIA [40]. Most of the LFIA is covered with a plastic cage to protect the strip from physical damage or humidity. The first part of LFIA is a sample pad that holds and initiates the sample flow to the next part. It also pretreats samples to produce a continuous sample flow in a smooth and homogenous manner. The sample pad is usually made from cellulose paper or glass-fiber paper. The next part of LFIA is called the conjugate pad. Labeled bioreceptor solution (usually antibodies) in its dry form is impregnated in the conjugate pad. The conjugate pad holds the labeled bioreceptor and releases it when rehydrated by the flowed sample. Here, the reaction occurs between antibody bioreceptor

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with analyte (if presence) to form an analytelabeled-antibody complex. Subsequently, the analyte-labeled-antibody complex flows into the main part of LFIA, the nitrocellulose (NC) membrane, which is available in different grades based on pore size, flow rate, color, or coating materials. Test and control lines are drawn on this NC membrane with a specific antibody against analyte and labeled conjugate, respectively. Here, the analyte-labeled-antibody complex flows from the conjugate pad and binds to a specific test line, while the excess-labeled antibody binds to the control line. The accumulation of the AuNP-labeled antibody at the test and control lines forms a red color due to SPR of AuNPs. The color formation can be detected either by the naked eye (visual detection) or any reader device for quantitative analysis. The subsections 1.4.1.1 and 1.4.1.2 discuss these reading methods. Finally, the sample flows to the adsorbent pad that is placed at the end of LFIA. The adsorbent pad acts as a holding sink, maintaining the flow rate, and preventing any backflow of the sample. The material used is usually cellulose, which is similar to a sample pad [40,132]. 1.4.1.1 Visual Detection Visual colored signals produced by some NP labels can be detected directly by the naked eye. This type of NP is the best choice for developing LFIA and can be categorized FIGURE 1.6 Schematic representation of lateralflow immunoassay (LFIA) strips.

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into colored NPs. Examples of colored NPs are colloidal AuNPs, colloidal AgNPs, colloidal CNPs, and colloidal selenium (SNPs). The resulting interpretation and analysis can be divided into qualitative and quantitative sections. LFIA, which uses colored NPs as labels, usually provides qualitative results either in the presence (positive) or absence (negative) of analytes. However, a major drawback for qualitative interpretation is that it is prone to human error during visual interpretation [133]. Quantitative interpretation is more advanced than qualitative interpretation; the former provides easy interpretation and eliminates human error. However, this approach requires reading devices. One of the colored NPs that is most commonly used as a label in LFIA is colloidal AuNPs. AuNPs have several advantages. It is compatible to biomolecules, stable with time, size-tunable, and easy to synthesize and modify; it provides an intense red color via a visible light source and can be detected by the naked eye [40,132]. Given these properties, AuNPs are widely selected as a label in developing LFIA-based biosensors. AuNP-based LFIA either uses a sandwich IA format in which the results are interpreted as positive with a line color, or a competitive format wherein results are interpreted contrariwise with a line color considered negative [132]. AuNP-based LFIA was developed for detecting foodborne pathogens such as Vibrio spp. [28,134
137], Salmonella spp. [138
140], E. coli O157:H7 [141,142], H. pylori [143], staphylococcal enterotoxin B [144], and mycotoxin [145], and results were interpreted visually and considered for qualitative measurement. By contrast, Song et al. developed multiplex LFIA which was read both qualitatively (naked eye) and semiquantitatively using a strip reader [146]. Another colored label used in developing LFIA biosensors is CNP, which is more sensitive than AuNPs for detecting foodborne pathogens [147
149]. These researchers

combined PCR with LFIA in their studies. The CNPs produced a black color in line(s) against the white background of the NC membrane. 1.4.1.2 Reader Device Detection In contrast to qualitative analysis, quantitative analysis requires a reader device. Some NPs also need a reader device to read the luminescent or fluorescent signal produced by QDs, UCPs, and dye-doped NPs quantitatively. These types of LFIAs are more sensitive than colored NP-based LFIAs. Unfortunately, the need for a reader device, difficulty in NP synthesis, and increased manufacturing cost are major drawbacks of these NPs [40]. These types of NPs are useful in developing multiplex LFIAs as NPs produce a special quantumsize effect, while different particle sizes emit spectra at varying wavelengths when excited and, thus, it can differentiate between analytes [150]. Furthermore, semiquantitative and quantitative results from strip reader devices can be used to detect colored NPs (AuNPs). In this technique, the color intensity, either in test or control lines, is simply measured and converted into optical density (OD) [131,146]. With the advancement in a strip reader device, the technology, sensitivity, and specificity of multiplex LFIA have been improved by reducing human error (visual interpretation) [151]. QDs are among the labels used in multiplex LFIA because they can simultaneously quantify multianalytes. QDs can produce strong luminescence, symmetric photoluminescence spectra, broad absorption, photostability, narrow size, and size-tunable emission [131]. Wang et al. developed aptamer-based LFIAs for detecting mycotoxin and ochratoxin A (OTA) [152]. In their study, the test line was interpreted either qualitatively or quantitatively. Qualitatively, visual results showed an inverse relationship in which the OTA concentration in the solution increased and the fluorescent band in the test zone decreased with

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

1.4 MEMBRANE-BASED BIOSENSORS

LOD at 5 ng/mL and was undetectable at 10 ng/mL. Image Analysis Software was used to quantitatively detect results of the strips and the average values of the peak area in the test zone were plotted as a function of OTA concentration. Under the competitive recognition model, high levels of targeted OTAs in solution resulted in a weak signal in the test zone. The range for quantitative detection was between 0 and 10 ng/mL. Moreover, other research teams also developed a competitive LFIA by using QD nanobeads as amplification probes to quantitatively and ultrasensitively detect zearalenone and aflatoxin B1 with LODs of 62.5 and 0.42 pg/mL, respectively [153,154]. The strip was scanned with an optical reader to provide quantitative measurements. UCPs are also categorized as NPs that need a reader device for detection. UCPs have unique properties; they can emit high energy, either as UV light or visible light, when irradiated with low-energy IR light. Several published works have investigated UCPs for detecting foodborne pathogens [155
158]. Qiu et al. [156] described a portable, light, small footprint, general-purpose analyzer to control flow in IA cassettes and facilitate the detection of test results. The test results were detected with UCP labels that were excited at IR frequencies and emitted in the visible spectrum. Meanwhile, MNPs require a reader device to quantitatively measure the magnetic signal produced. MNPs have advantages over other NPs because they can measure entirely captured analytes that move toward the surface and within the membrane and is usually undetectable [131]. Koets et al. [159] developed such LFIAs for detecting E. coli DNA. Dye-doped polystyrene NPs are another class of nanomaterials used to dope dye molecules. Dye-doped polystyrene NPs can provide higher fluorescence signals than free-dye molecules, thereby increasing the performance of LFIA. Xie et al. [160] reported that the use of fluorescein isothiocyanate (FITC)
doped

25

polystyrene NPs as labels for fluorescent LFIA provided superior results compared with the use of AuNP-based LFIAs in terms of sensitivity, use of antibodies, and coefficients of variation [160]. This approach was applied to detect mycotoxins in food samples [161,162]. Furthermore, Eu chelates substituted FITC to dope polystyrene NPs to fabricate timeresolved fluorescent LFIAs and have been used for mycotoxin detection because of their unique luminescence properties [163,164].

1.4.2 Other Membrane-Based Biosensors Other forms of membrane-based biosensors were reported for the detection of foodborne pathogens. In general, membrane-based biosensors can be classified based on transduction systems, such as electrical, optical, and piezoelectric. In electrical transduction systems, membranes are used as a platform to allow the biorecognition reaction to occur. McGraw et al., [165] used polypropylene microfiber membranes coated with a conductive PPy and antibody. Binding E. coli O157:H7 on the membrane via antibodies increased the resistance at the electrotextile electrode surface to indicate a positive result. Another type of impedance spectroscopy was also used in membranebased biosensors to record the binding of the pathogen on the membrane [166]. Instead of physical support, the properties of porous membranes could be used to measure transduction signals. Target ssDNA bound to a complimentary DNA strand were immobilized at a membrane pore. DNA hybridization could create blockage at membrane pores, which led to changes in the electrochemical parameters, and were then used to detect the DNA of E. coli O157:H7 [91,167]. By contrast, for amperometric detection, a graphite-Teflon-peroxidase-ferrocene electrode was used to detect hydrogen peroxidase [168]. Hydrogen peroxide consumption or generation

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

26

1. ADVANCED NANOPARTICLE-BASED BIOSENSORS FOR DIAGNOSING FOODBORNE PATHOGENS

was a result of catalase activity, either by catalase-positive or catalase-negative bacteria in complex culture media. E. coli and Streptococcus pneumoniae were successfully detected using this approach. Changes in conductivity could also be observed with the use of a conductive membrane as a filter for detecting Salmonella bacteria [169]. In another conductometric transduction, antibodies were used to conjugate with a conductive material to close the electrical circuit. First, a conductive conjugated antibody was bound to the target analyte and concentrated using magnetic separation. The concentrated bacteria
antibody complex then bound to the secondary antibody immobilized on the NC nanofilament membrane. This phenomenon changed resistivity; thus, the concentration of E. coli O157:H7 and the virus could be measured [170,171]. Piezoelectric quartz crystals, as discussed previously in section 1.3.3.1, were also applied in membrane-based biosensors. In mass sensing, the analyte-mediated surface agglutination of ganglioside (GM1)-functionalized lipidsupported membrane was used as a modified gold electrode. The GM1-functionalized membrane was produced via the spontaneous spread of liposomes on a SAM of a long-chain alkanethiol. This gold-modified electrode was used to obtain cholera toxin (CT)-specific agglutination at the surface of GM1. The binding of CT to the surface of GM1 changed the resonance frequencies of the piezoelectric crystal. Furthermore, this reaction increased the density and viscosity at the interface and, thus, the frequency of the piezoelectric crystal decreased. The LOD of detected CT was at 25 ng/mL [172].

1.5 MULTIPLEX BIOSENSORS Multiple types of analytes can be analyzed simultaneously with the single application of the sample, especially when applied to clinical diagnosis. Single analysis can increase the cost

and amount of a limited sample, and it requires different assay conditions and sample preparation; thus, the test is tedious, and more time is required to report the results. To overcome these problems, several analytical methods have been developed to perform multiple detection of analytes simultaneously. Biosensors based on multiplex detection can be grouped according to their transducer mechanisms and principles. These groups include electrochemical, optical, and LFIA as membrane-based biosensors. For electrochemical mechanisms, magnetogenosensing was used for simultaneous detection and three foodborne pathogens were differentiated based on their DNA sequence [173]. Primers were tagged with fluorescein, biotin, and digoxigenin coding for S. enterica, L. monocytogenes, and E. coli, respectively, and used to amplify gene sequences of those pathogens by applying PCR. The single-tagged amplicons were then immobilized on silica MNPs (Fig. 1.7). This detection was based on the nucleic acid-binding properties of silica particles with the help of chaotropic agents, such as guanidinium thiocyanate. Electrochemical magnetogenosensing on silica MNP was detected by amperometric measurement. In another work an NP-functionalized multijunction biosensor used as a sensor chip was developed as a multiplex biosensor [174]. A singlewalled carbon nanotube (SWCNT) and polyethylenimine were used to coat Au tungsten wires. These wires were then aligned to form 2 3 2 junction arrays and functionalized with streptavidin and biotinylated antibodies specific for E. coli K-12 and S. aureus. Changes in current, I, were measured from the antiE. coli
functionalized junctions and anti-S. aureus junctions and then compared with the calibration curves. Another multiplex approach using the electrochemical mechanism was based on a modified screen-printed electrode, such as multiple DNA-based detection [37]. In this approach, a four-channel electrode was

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

1.5 MULTIPLEX BIOSENSORS

27

FIGURE 1.7 Steps involved in electrochemical multiplex magnetogenosensing include: (A) Single-tagging PCR amplification of S. enterica, L. monocytogenes, and E. coli genes; (B) tagged amplicon immobilized on silica MP; (C) enzymatic labeling; and (D) magnetic capture following amperometric measurement. Source: Reproduced with permission from Ref. S. Lie´bana, D. Branda˜o, P. Corte´s, S. Campoy, S. Alegret, M.I. Pividori, Electrochemical genosensing of Salmonella, Listeria and Escherichia coli on silica magnetic particles, Anal. Chim. Acta 904 (2016) 1
9, doi:10.1016/j.aca.2015.09.044.

added, [175] and three antibodies conjugated with specific nanocrystals were measured with square-wave anodic stripping voltammetry (SWASV) [176]. For optical-based multiplex biosensors, unique properties of NPs and fluorescent as colorimetric agents have been applied for foodborne pathogen detection. A fluorescent nanobiosensor coupled with nanobead-based immunomagnetic separation has been developed [177]. This portable and automatic instrument recognizes and differentiates three pathogens by measuring the fluorescence

intensities of each pathogen emitted by different QD-conjugated antibodies to each of the three target pathogens (Fig. 1.8A). Other studies reported on using fluorescent materials [178
180], magnetic beads [181], and AuNPs [182] for multiplex detection. SPR-based detection to simultaneously identify and differentiate a few foodborne pathogens was also reported. Multichannel SPR composed of several channels that corresponded to each pathogen for detection. The sensor chip of each channel was functionalized with the corresponding antibodies specific to the targeted pathogens [183].

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

28

1. ADVANCED NANOPARTICLE-BASED BIOSENSORS FOR DIAGNOSING FOODBORNE PATHOGENS

FIGURE 1.8

Optical multiplex biosensors. (A) Quantum dot (QD) fluorescent-based (reproduced with permission from Ref. L. Xu, Z. Lu, L. Cao, H. Pang, Q. Zhang, Y. Fu, et al., In-field detection of multiple pathogenic bacteria in food products using a portable fluorescent biosensing system, Food Control 75 (2017) 21
28, doi:10.1016/j.foodcont.2016.12.018 [177]). (B) SPR-based detection involving three steps: (1) incubation of the sensor with crude food samples, (2) binding of secondary biotinylated antibody, and (3) binding of streptavidin-coated AuNP to the biotinylated antibody (reproduced ˇ with permission from Ref. H. Vaisocherova´-Lı´salova´, I. Vı´sˇ ova´, M.L. Ermini, T. Springer, X.C. Song, J. Mra´zek, et al., Lowfouling surface plasmon resonance biosensor for multi-step detection of foodborne bacterial pathogens in complex food samples, Biosens. Bioelectron. 80 (2016) 84
90, doi:10.1016/j.bios.2016.01.040 [104]). (C) SERS spectra showing different patterns corresponding to each pathogen (reproduced with permission from Ref. J. Sundaram, B. Park, Y. Kwon, K.C. Lawrence, Surface enhanced Raman scattering (SERS) with biopolymer encapsulated silver nanosubstrates for rapid detection of foodborne pathogens, Int. J. Food Microbiol. 167 (2013) 67
73, doi:10.1016/j.ijfoodmicro.2013.05.013 [117]).

Moreover, ultralow fouling functionalizable poly(carboxybetaine acrylamide) (pCBAA) brushes were introduced as a surface platform with certain advantages such as rapid process,

improvement in fouling resistance, and sensitive detection of bacterial pathogens [104]. This method involved steps for pathogen detection, as illustrated in Fig. 1.8B. Previously, use of

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

1.6 PERFORMANCE OF NANOPARTICLE-BASED BIOSENSORS

29

FIGURE 1.9 Membrane-based multiplex biosensors. (A) The bacteria-sensing mechanism of nanoporous membrane via impedance spectrum. (B) The principle of simultaneous detection using the microfluidic device integrated with nanoporous membranes to detect two types of bacteria. Source: Reproduced with permission from Ref. F. Tian, J. Lyu, J. Shi, F. Tan, M. Yang, A polymeric microfluidic device integrated with nanoporous alumina membranes for simultaneous detection of multiple foodborne pathogens, Sens. Act. B Chem 225 (2016) 312
318, doi:10.1016/j.snb.2015.11.059.

surface-enhanced Raman scattering (SERS) with biopolymer-encapsulated silver nanosubstrate was reported [117]. This method produced different patterns of Raman spectra that corresponded to each targeted pathogen (Fig. 1.8C). Finally, membrane-based biosensors with various multiplex strategies have been reported. In LFIA, strategies for multiplexing [133] include development of multiple test lines corresponding to the amount of target pathogens [184], a single test line to detect multiple pathogens [185], and architecture tuning [186]. For other membranes, non-biofouling polyethylene glycol (PEG)
based microfluidic chip was developed and integrated with a functionalized nanoporous alumina membrane [187]. This PEG microfluidic device was based on impedance measurements for the simultaneous detection of E. coli O157:H7 and S. aureus from mixed samples. The principles are shown in Fig. 1.9.

1.6 PERFORMANCE OF NANOPARTICLE-BASED BIOSENSORS A wide array of different pathogens has been detected mainly through nucleic acid or whole bacteria. The LOD was difficult to compare due to different units under various detection strategies. Comparing considerably

small mass/volume of nucleic acid LOD with whole bacteria is not fair because whole bacteria are larger than their DNA, unless they have been amplified through PCR [128]. As a result, nucleic acid amplification could increase LOD. In contrast, measurement of whole bacteria was logically limited and must be amplified. The amplification for whole bacteria was performed by pre-enrichment and the sample containing bacteria was incubated using a certain growth medium with fixed temperature and time [142]. However, this additional step required more time and cost. To overcome these problems a few enhancement strategies were introduced, especially for LFIA-based biosensors. These methods included enhanced chemiluminescence immunosensors [188]; use of silver enhancement technology [189]; combination of preenrichment treatment of the detection samples by MNPs and fluorescence [179] or AuNPs [190] before using LFIA strips for signal amplification; postenrichment, such as enzyme signal amplification with AuNPs [191] or MNPs [192]; and measurement with surface-enhanced Raman spectroscopy [116]. Other strategies to improve sensitivity include optimization of antibodies conjugated to the surface of AuNPs on the conjugation pad [193], gold nanocomposites such as core-shell structured Ag/Au NPs [194], magnetic

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

TABLE 1.2 Performance of Reported Nanoparticle-based Biosensors by Format and Transducer Mechanism Transducers

Bioreceptors

Nanoparticles

Targets

LOD

Specificity Assay Time

Refs.

Amperometric

Nucleic acid

None

Vibrio cholerae

10 CFU/mL

100%

Not reported

[37]

Impedimetric

Nucleic acid

None

Listeria monocytogenes

267 pM of PCR amplicon

High

Not reported

[49]

Amperometric

Nucleic acid

None

Escherichia coli

2
20 CFU/100 mL

High

3
5 h

[50]

High

Not reported

[61]

Colorimetric

Aptamer

None

3

Salmonella enteritidis

10 CFU/mL

Amperometric

Antibody

AuNPs

E. coli

Standard: 1.99 3 10 CFU/mL; NPs amplified: 50 CFU/mL

High

Not reported

[67]

Amperometric

Antibody

AuNP/MWCNT bionanocomposite

E. coli O157:H7

B30 CFU/mL

High

Not reported

[68]

Amperometric

Antibody

Au-SiO2 nanocomposite

E. coli O157:H7

15 CFU/mL

High

Not reported

[69]

Amperometric

Antibody

AuNPs

L. monocytogenes

102 CFU/mL

High

B1 h

[70]

Amperometric

Antibody

None

Clostridium tetani

0.011 IU/mL

High

,50 min

[71]

Potentiometric

Antibody

None

E. coli

10 cells/mL

High

1.5 h

[76]

Potentiometric

Nucleic acid

MNPs

Vibrio alginolyticus

10 CFU/mL

High

Not reported

[79]

Potentiometric

Aptamer

None

L. monocytogenes

10 CFU/mL

High

Not reported

[80]

2

4

Potentiometric

Aptamer

CNT

Staphylococcus aureus

8 3 10 CFU/mL

High

B8 min

[81]

Potentiometric

Aptamer

SWCNT

E. coli O157:H7

6
26 CFU/mL

High

Few minutes

[82]

Potentiometric

Aptamer

Graphene oxide (RGO)

S. aureus

1 CFU/mL

High

1
2 min

[83]

Impedimetric

Antibody

AuNPs

Bacillus cereus

10 CFU/mL

High

Not reported

[36]

E. coli O157:H7

1.5 3 10 CFU/mL

High

Not reported

[85]

Impedimetric

Antibody

AuNPs

2

3

Impedimetric

Antibody

None

S. enteritidis

10 CFU/mL

High

5 min

[86]

Impedimetric

Antibody

Graphene
graphene oxide nanocomposite

Salmonella typhimurium

B10 CFU/mL

High

Not reported

[87]

Impedimetric

Protein

None

E. coli

75 cells/mL

High

B70 min

[88]

Impedimetric

Antibody

None

Internalin B of L. monocytogenes

4.1 pg/mL

High

Not reported

[89]

Impedimetric

Antibody

None

E. coli O157:H7

B7 CFU/mL

High

Not reported

[90]

Impedimetric

Antibody

Alumina nanoporous

E. coli O157:H7

83.7 CFU/mL

High

Not reported

[91]

High

Not reported

[92]

Impedimetric

Bacteriophage MNPs

E. coli K-12

3

10 CFU/mL

Impedimetric

Antibody

Graphene nanoplatelets (GNPs)

E. coli O157:H7

10 cells/mL

High

Not reported

[93]

Impedimetric

Antibody

AuNPs

Salmonella spp.

102 CFU/mL

High

40 min

[94]

Impedimetric

Aptamer

None

S. typhimurium

6 CFU/mL

High

30 min

[95]

Impedimetric

Antibody

ITO

E. coli O157:H7

1 CFU/mL

High

,5 min

[97]

Impedimetric

Antibody

MNPs

E. coli O157:H7 and S. typhimurium

2.05 3 10 CFU/g and 1.04 3 103 CFU/mL

High

Not reported

[98]

Colorimetric

Antibody

Pt/MNPs

E. coli O157:H7

10 CFU/mL

High

30 min

[99]

3

2

Impedimetric

Protein

None

E. coli, Pseudomonas aeruginosa, S. aureus, and S. epidermidis

10 CFU/mL

High

1.5 h

[100]

Conductometric

Antibody

None

B. cereus

35 CFU/mL

High

6 min

[101]

Conductometric

Antibody

None

E. coli

1 CFU/mL

High

Not reported

[102]

SPR

Antibody

AuNPs

E. coli O157:H7 and Salmonella sp.

57 and 17 CFU/mL; and 7.4 3 103 and 11.7 3 103 CFU/mL

High

,80 min

[104]

SPR

Bacteriophage None

Salmonella

1.3 3 107 CFU/mL

High

Not reported

[107]

SPR SPR Optical fibers

Bacteriophage None Antibody Antibody

None None

E. coli and MRSA

10 CFU/mL

High

,20 min

[108]

E. coli O157:H7

10
10 CFU/mL

High

30 min

[109]

E. coli O157:H7

5.2 3 10 CFU/g

100%

25 min

[111]

10 CFU/g

High

5 min

[112]

27 and 19 CFU/Ml

High

Not reported

[115]

102 CFU/mL

High

,1 h

[116]

S. typhimurium

Optical fibers

Antibody

None

Raman & LFIA

Nucleic acid

AuMBA@Ag core-shell S. enteritidis and NPs L. monocytogenes

Raman

Antibody

Fe3O4/Au MNPs

E. coli O157:H7

3 2

3

2

5

FT-IR

Antibody

TiO2

E. coli O157:H7

10 CFU/mL

High

,30 min

[118]

Piezoelectric

Antibody

AuNPs

E. coli O157:H7

,10 CFU/mL or g

High

24 h

[35]

Piezoelectric

Nucleic acid

AuNPs

E. coli O157:H7

1.2 3 102 CFU/mL

High

B6 h

[120]

Piezoelectric

Antibody

None

E. coli O157:H7

102
105 CFU/mL

High

Not reported

[121]

Piezoelectric

Antibody

Micro/nanobeads

E. coli O157:H7

103 CFU/mL

High

Not reported

[122]

Piezoelectric

Antibody

AuNPs

C. jejuni

150 CFU/mL

High

Not reported

[123]

Piezoelectric

Antibody

AuNPs

S. typhimurium

B10
20 CFU/mL

High

12 min

[124]

Piezoelectric

Protein

None

3-Methyl-1-butanol and 1-hexanol as indicator for Salmonella

,5 ppm

Not very good

Not reported

[125]

2

(Continued)

TABLE 1.2 (Continued) Transducers

Bioreceptors

Nanoparticles

Targets

LOD

Specificity Assay Time

Refs.

Magnetoelastic

Phage

None

S. typhimurium

4 CFU/g

High

7h

[127]

LFIA

Nucleic acid

AuNPs

V. cholerae

10 CFU/mL

100%

Not reported

[28]

Vibrio parahaemolyticus

1.58 3 10 CFU/g

High

B10 min

[135]

LFIA LFIA

Antibody Antibody

MNPs AuNPs

V. cholerae O1 and O139

2

2

98
100% 6 h incubation [136]

4

10 CFU/mL after enrichment

LFIA

Antibody

AuNPs

V. cholerae O139

10 and 10 CFU/mL after enrichment

High

20 min and 12 h incubation

[137]

LFIA

Antibody

AuNPs

S. enteritidis

104 CFU/mL

High

20 min

[139]

Salmonella Typhi

1.14 3 10 CFU/mL

100%

15 min

[140]

LFIA

Antibody

AuNPs

5

6

LFIA

Antibody

AuNPs

E. coli O157:H7

10 CFU/mL

High

15 min

[141]

LFIA

Antibody

AuNPs

E. coli O157:H7

2 CFU/g

95%

Not reported

[142]

LFIA

Antibody

AuNPs

Staphylococcal enterotoxin B

1 ng/mL

High

,5 min

[144]

LFIA

Antibody

AuNPs

AFB1, ZEA, and DON

0.03, 1.6 & 10 μg/kg

High

15 min

[146]

5

LFIA

Nucleic acid

CNPs

E. coli

10
10 CFU/mL

100%

,1 h

[147]

LFIA

Nucleic acid

CNPs

Listeria spp. and L. monocytogenes

10 cells/25 mL

High

28 h

[148]

LFIA

Nucleic acid

CNPs

Cronobacter spp.

1 cell/g

High

16 h

[149]

LFIA

Aptamer

QDs

OTA

1.9 ng/mL

High

10 min

[152]

LFIA

Antibody

CdSe/ZnS QDs

ZEN

3.6 μg/kg

High

Not reported

[153]

LFIA

Antibody

CdSe/ZnS QDs

AFB1

0.42 pg/mL

High

15 min

[154]

LFIA

Antibody

UCPs

Brucella

2 3 103
3.9 3 105 CFU/mg

High

Not reported

[155]

LFIA

Antibody

NaYF4:Yb,Er NPs

Vibrio anguillarum

102 CFU/mL

High

15 min

[158]

LFIA

Nucleic acid

MNPs

E. coli

4
50 pM

High

3 min

[159]

4

4

LFIA

Antibody

FITC-doped polystyrene NPs

E. coli O157:H7

10 CFU/mL

High

Not reported

[160]

LFIA

Antibody

Fluorescent microspheres

AFB1

2.5 mg/L

High

Not reported

[161]

LFIA

Antibody

Eu(III) NPs

OTA

1 μg/kg

High

8 min

[163]

Membrane-based Antibody

None

E. coli O157:H7

log 0
9 CFU/mL

High

15 min

[165]

Membrane-based Antibody

Nanofibrous

E. coli O157:H7

61 CFU/mL

High

8 min

[171]

1.7 CONCLUSION AND FUTURE PERSPECTIVES

nanogold microspheres [195], and shape modification such as gold nanoflowers (AuNFs) [196]. In terms of specificity, almost all developed biosensor devices are highly specific due to the designed bioreceptor in their biosensors that can recognize targeted pathogens, as discussed in previous bioreceptor sections. Antibodies can recognize a pathogen by the unique epitope presented in that pathogen, and enzymes can recognize a specific substrate to produce a reaction. Nucleic acids can recognize a specific sequence of every pathogen. Aptamers are similar to nucleic acid, which has a unique recognition site, and other bioreceptors with their own specific recognition site. Table 1.2 summarizes the performance in terms of sensitivity and specificity of the reported biosensors based on various formats and transducer mechanisms.

1.7 CONCLUSION AND FUTURE PERSPECTIVES Advances in NP-based biosensors for detecting foodborne pathogens aids in controlling and preventing the outbreak of diseases. The improvements in technology can help to identify the pathogens involved, and action can be taken accordingly. Biosensors make use of recognition elements to detect the presence of pathogens and use transducers to process and interpret the results to end users. Neither can function alone as a complete biosensor. The commonly used bioreceptors and transducers provide promising sensitivity and selectivity in detecting foodborne pathogens. Most of the discussed methods in this chapter are commercially available, such as Malthus 2000, Midas Pro, and Bactometer for electrochemical processes; Biosensing Instrument, BIA-core, Nanolane, Lumac Biocounte, and OWLS Sensor for optical electrical processes; and Axela PZ 106 Immuno-biosensor System, Attana sensor technologies, KSC Instruments,

33

and QSENSE Concentris for mass-sensitive mechanisms [73]. Although NP-based biosensor devices offer numerous advantages—such as efficiency, simplicity, sensitivity, selectivity, and robustness—these devices also have several disadvantages such as limited sensitivity and high-cost in some electrochemical-based and optical-based transducer systems. Implementing a reader device and a synthesis method of certain NPs can be complicated, thereby increasing the production cost. The cost to produce and synthesize bioreceptors such as antibodies and aptamers is also high. Therefore better and new equipment and tools are needed to overcome such problems. Among the mentioned transducer mechanisms, electrochemical-based and lateral flow
based bioreceptors are simple and user friendly. These devices are also designed for in-field detection because they are portable, simple, and have a long shelf life. The potential multiplexing capability of the newly developed biosensors is important to reduce cost while simultaneously detecting targeted foodborne pathogens. Recently, the electronic nose (e-nose) has become a potential instrument for food safety assessment due to its capability for rapid early detection of contamination and defects in the food production chain. E-nose was designed as a measurement system based on the detection of complex odors by mimicking the working principle of the mammalian olfactory system [197]. Furthermore, the combination of biosensors with a smartphone device to interpret results is a perfect match in the current era of smartphones and devices. As an example, a novel magnetic biosensor for the visual and quantitative detection of S. enteritidis from milk, cheese, and water has been developed. Signal amplification produces a visual color that can be detected by a smartphone device at an extremely low concentration with high sensitivity [198]. These approaches can be

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

34

1. ADVANCED NANOPARTICLE-BASED BIOSENSORS FOR DIAGNOSING FOODBORNE PATHOGENS

further improved and become future alternative biosensor devices to diagnose foodborne pathogens.

Acknowledgments The authors are grateful for the technical support from the School of Materials and Mineral Resources, Institute for Research in Molecular Medicine (INFORMM), and Universiti Sains Malaysia. This research was funded by TRGS 203/PBahan/6763001.

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C H A P T E R

2 Aptamer Technology for the Detection of Foodborne Pathogens and Toxins Alok Kumar1,*, Madhu Malinee2,*, Abhijeet Dhiman3,4, Amit Kumar5 and Tarun Kumar Sharma6,** 1

Department of Immunology and Genomic Medicine, Graduate School of Medicine, Kyoto University, Kyoto, Japan, 2Department of Anatomy and Developmental Biology, Graduate School of Medicine, Kyoto University, Kyoto, Japan, 3Department of Biotechnology, All India Institute of Medical Sciences, Delhi, India, 4Faculty of Pharmacy, Uttarakhand Technical University (UTU), Dehradun, India, 5Discipline of Biosciences and Biomedical Engineering, Indian Institute of Technology Indore, Simrol, Indore, India, 6Center for Bio-design and Diagnostics, Translational Health Science and Technology Institute, NCR Biotech Science Cluster, Faridabad, India

Moreover, food poising causes B20 million deaths worldwide every year [1]. Food poisoning is a major concern for the food industry. The detection of foodborne pathogens is essential as the failure of such detection could cause severe disease outbreaks. Recently in the United States a multistate outbreak of “Shiga toxin” producing Escherichia coli O157:H7 infections linked to romaine lettuce was reported in April 2018. The outbreak covered 32 states and more than 150 people were infected. In another incident, there was a multistate outbreak of Salmonella typhimurium from infected coconuts [2]. The Center for Disease

2.1 INTRODUCTION Food contaminated with disease-causing agents such as bacteria, viruses, parasites, toxins, and chemicals cause foodborne illnesses (or diseases) which are often referred to as food poisoning. Foodborne diseases (FBDs) are a growing public health concern at the global level and comprise a wide spectrum of illnesses. According to the World Health Organization (WHO) report in 2015, over 250 diseases are caused by consuming foods contaminated with harmful bacteria, parasites, viruses, toxins, and chemical substances. *Represent equal contribution made by authors. **Corresponding author.

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© 2019 Elsevier Inc. All rights reserved.

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2. APTAMER TECHNOLOGY FOR THE DETECTION OF FOODBORNE PATHOGENS AND TOXINS

Control and Prevention (CDC), an organization under the US Department of Health and Human Services, serves as a lead coordinator between public health partners to detect an outbreak, define its size and extent, identify the source, and monitor the effectiveness of prevention and control efforts. According to the report published in December 2015 by the Foodborne Disease Burden Epidemiology Reference Group, a committee established by the WHO to estimate the global burden of FBDs annually, 1 out of 10 people in the world suffers from a FBD. Furthermore, diarrheal diseases which cause more than half of global FBDs are the most common causes of illnesses (550,000 cases) and deaths (230,000 deaths) worldwide [1]. Foodborne illnesses not only harm the health of people, they also affect the economy of the country. According to an estimate, the average cost associated with each case of foodborne illness is US$1068 in the United States. Moreover, food safety incidents account for huge losses (around US$7 billion per year) to the economy of the United States alone [3
7].

2.1.1 Common Foodborne Pathogens The US Food and Drug Administration and US Public Health Service have identified the major pathogens (bacteria and viruses) that cause FBDs pandemically. These pathogens are Salmonella enteritidis, Norovirus (Norwalk Virus), Campylobacter jejuni, E. coli O157:H7, Listeria monocytogens, Clostridium perfringens, Clostridium botulinum, Staphylococcus aureus, Shigella, Toxoplasma gondii, Vibrio vulnificus, Vibrio cholera, Vibrio parahaemolyticus, Yersinia enterocolitica, and others. In October 2016 there was a FBD outbreak associated with poultry and backyard flocks. The outbreak occurred in 48 states in the United States. The causative agent for the outbreak was Salmonella sp. and the associated serotypes were Enteritidis, Muenster, Hada, Indiana, Mbandaka, Braenderup, and Infantis. In March 2017

another multistate FBD outbreak in the United States was reported concerning the spread of L. monocytogens from contaminated cheese. A comprehensive list of foodborne pathogens, the illnesses caused by them, and the symptoms and outbreaks associated with them are listed in Table 2.1.

2.1.2 Norovirus Norovirus (NoV), a highly contagious virus worldwide, is thought to be the most common cause of acute gastroenteritis (diarrhea and vomiting). It accounts for a greater fraction of nonbacterial foodborne illnesses. Norovirus spreads easily through food and drinks. According to a report from the CDC, “Globally, norovirus is estimated to be the most common cause of acute gastroenteritis and is responsible for 685 million cases every year, 200 millions of these cases are among children younger than 5 years old. This leads to an estimated 50,000 child deaths every year that occur mostly in developing countries” [49]. The sign and symptoms usually start to appear within 12
48 hours after infection. The symptoms include nausea, vomiting, abdominal pain, diarrhea, fever, and muscle pain. 2.1.2.1 Salmonella spp. Salmonella, a Gram-negative nonsporeforming bacterium, is the agent causing salmonellosis. Salmonella alone causes over one million illnesses annually worldwide. Salmonella causes two types of illnesses: (1) typhoidal illness, a serious condition with symptoms like high fever, diarrhea, headaches, and lethargy and causes the death of about 1 in 10 patients; and (2) nontyphoidal illness, a comparatively mild condition with symptoms like nausea, vomiting, diarrhea, and cramps and the symptoms last only for a few days. S. enterica and S. bongori are two species that cause human illnesses worldwide. Australia is the world leader in Salmonella outbreaks where it has been responsible for about 15 deaths per

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2.1 INTRODUCTION

TABLE 2.1 List of Pathogenic Microorganisms and the Outbreaks Caused by Them Name of Microorganism Disease

Symptoms

Incubation Period

Infective Dose

Contaminated Food

References

VIRUS Norovirus

Viral gastroenteritis or winter vomiting disease

Abdominal cramps, diarrhea, nausea, vomiting, headache, and fever

1
2 days

Low

Raw oysters, salad, shellfish, coleslaw, baked goods, frosting, and contaminated water and ice

[8
11]

Hepatitis A virus

Hepatitis A

Nausea, anorexia, fever, and jaundice

Unknown

10
100

Sandwiches, fruits, milk, vegetables, salads, and shellfish

[12,13]

Hepatitis E virus

Hepatitis E

Fever, anorexia, vomiting, abdominal pain, and jaundice

2
10 weeks

Unknown Water and raw pig livers

[14
16]

Rotavirus

Gastroenteritis

Vomiting and/or diarrhea 2 days and in some cases fever

10
100

Water

[17,18]

10

Raw or undercooked pork

[19,20]

PARASITIC PROTOZOA AND WORMS Toxoplasma gondii

Toxoplasmosis

Seizures, enlarged liver and spleen, jaundice, severe eye infections, loss of hearing, and mental retardation

Giardia lamblia

Giardiasis

Stomach cramps, diarrhea, 1
2 weeks and nausea

10

Food, drinks water, or coming into contact with surfaces or objects contaminated by the parasite or its cysts

[21,22]

Entamoeba histolytica

Gastrointestinal amebiasis

Diarrhea or bowel movements streaked with blood or mucus, cramps, fever, vomiting, and nausea

Few days to a month

1000

Contaminated food and water

[23]

Campylobacter jejuni

Diarrhea

Diarrhea (sometimes bloody), abdominal cramps, fever, muscle pain, headache, and nausea

2
5 days

400
500

Raw and undercooked, shellfish, poultry, and other meat, raw milk and untreated water

[24,25]

Yersinia enterocolitica

Yersiniosis

Diarrhea, vomiting, fever, and abdominal pain

1
3 days

Unknown Pork, oysters, fish, and unpasteurized milk

[26]

Shigella spp. (Only humans carry this bacterium)

Diarrhea

Diarrhea, fever, stomach cramps, vomiting, and bloody stools

1
2 days

10

[27
29]

Within 1 week

BACTERIAL

Salads, milk and dairy products, raw oysters, ground beef, poultry, and unclean water handled by someone who is infected with the bacterium

(Continued)

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2. APTAMER TECHNOLOGY FOR THE DETECTION OF FOODBORNE PATHOGENS AND TOXINS

TABLE 2.1 (Continued) Name of Microorganism Disease

Symptoms

Incubation Period

Infective Dose

Contaminated Food

References

Vibrio Gastroenteritis parahaemolyticus

Diarrhea, abdominal cramps, nausea, vomiting, and fever

6 h to 4 days

1,000,000

Raw or undercooked seafood like fish, shellfish, and oysters

[30
32]

Vibrio cholerae Serogroups O1 and O139

Cholerae

Sudden onset of profuse, painless, watery diarrhea

Few hours to 5 days

1,000,000

Raw or undercooked seafood

[33
35]

Vibrio cholerae Serogroups non-O1 and non-O139

Gastroenteritis

Diarrhea and abdominal cramps

12
24 h

1,000,000

Raw or undercooked seafood, particularly shellfish

[36,37]

Gastroenteritis, wound infection, and severe bloodstream infections

Within 24 h 100

Raw or undercooked seafood, particularly shellfish, oysters, clams, and crabs

[38]

Vibrio vulnificus Gastroenteritis; primary septicemia

People with liver diseases are especially at high risk Listeria monocytogenes

Listeriosis

Fever, headache, fatigue, muscle aches, nausea, vomiting, diarrhea, meningitis, and miscarriages in pregnant woman

9
48 h

1000

[39] Refrigerated, ready-to-eat foods (meat, poultry, seafood, and unpasteurized milk and dairy products)

Salmonella spp.

Salmonellosis

Diarrhea, fever, vomiting, headache, nausea, and stomach cramps

12
72 h

100,000

Raw and undercooked eggs, raw meat, poultry, seafood, raw milk, dairy products

[40
42]

Bacillus cerus

Gastroenteritis

Diarrhea, abdominal cramps, nausea, and vomiting

1
15 h

1,000,000

Milk and cheese products, vegetables, fish, and rice

[43]

Staphylococcus aureus

Mild skin infection to severe osteomyelitis

Vomiting, nausea, 30 min to stomach cramps, diarrhea, 6 h and pneumonia

100,000

Unpasteurized milk and cheese products, sliced meat, puddings, pastries, and sandwiches

[44]

Clostridium perfringens

Clostridial food poisoning

Abdominal pain, diarrhea, 8
16 h and sometimes nausea and vomiting

Unknown Undercooked meat and meat products

Clostridium botulinum

Botulism

Dry mouth, double vision followed by nausea, vomiting, and diarrhea. Later, constipation, weakness, muscle paralysis, and breathing problems

Low

12 h to 3 days

[45,46]

[47] Improperly prepared home-canned foods; vacuum-packed and tightly wrapped food, meat products, seafood, and herbal cooking oils

(Continued)

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2.1 INTRODUCTION

TABLE 2.1 (Continued) Name of Microorganism Disease Pathogenic E. coli O157:H7

Diarrhea

Symptoms

Incubation Period

Severe stomach cramps, 3
4 days bloody diarrhea, and nauseaCan cause permanent kidney damage which can lead to death in young children

year according to the Commonwealth Scientific and Industrial Research Organization (CSIRO), an independent Australian federal government agency responsible for scientific research. According to the CSIRO, there were 73 cases of salmonella infection per 100,000 people in 2015 in Australia [50]. 2.1.2.2 Escherichia coli O157:H7 E. coli O157:H7, is the most well-known serotype of Enterohemorrhagic Escherichia coli, a Gram-negative bacteria. E. coli O157:H7 was identified in 1982 during an outbreak in the United States. [51]. It resides in the intestinal tracts of livestock and spreads through their meat products. Its infection causes extreme bloody diarrhea and painful abdominal cramps. A severe complication, called hemolytic uremic syndrome, can occur in some cases which causes profuse bleeding and kidney failure. According to the policy of the United States Department of Agriculture Food Safety and Inspection Service (USDA-FSIS), the current limit of detection is 1 cfu/65 g of a meat sample. E. coli O157:H7 produces Shiga toxin which is the most potent toxin known to humanity. Shiga toxin mainly acts on the vascular endothelium. Shiga toxin, a heterodimer molecule (AB5), where B subunit helps in uptake by binding with membrane and subunit A disrupts the 60S subunit of the ribosome thereby completely halts protein synthesis. E. coli O157:H7 produces numerous other putative virulence factors as well in addition to Shiga toxin including some proteins which assist in the attachment and

Infective Dose 10

Contaminated Food

References

Meat (e.g., undercooked or raw hamburgers), uncooked produce, raw milk, unpasteurized juice, and contaminated water

[48]

colonization of the bacteria in the intestinal wall and which can lyse red blood cells and liberate iron to help support E. coli metabolism. According to the CDC report in 2012, this pathogenic strain infects more than 96,000 people in the United States each year [52]. Common sources of E. coli O157:H7 infection are undercooked ground beef, unpasteurized milk and juice, and ciders. During the multistate outbreak in the United States in April 2018, around 150 persons were infected across 32 states and the infection was linked to common lettuce. 2.1.2.3 Campylobacter jejuni C. jejuni, the most common bacterial
foodborne pathogen causing diarrheal diseases, causes around 1.3 million illnesses per year in the United States alone, according to the CDC report 2014 [53]. The common source of infection is raw and undercooked poultry products. People with C. jejuni infection usually have diarrhea (often bloody), fever, abdominal cramps, nausea, and vomiting. People become well generally within a few days without any treatment. But it may be a serious problem for immunecompromised persons. C. jejuni occasionally spreads via the bloodstream to the different organs where it may cause life-threatening infections. Newborn infants and children under the age of one year are at the highest risk of Campylobacter infections. 2.1.2.4 Staphylococcus aureus S. aureus, a Gram-positive bacteria, is an opportunistic pathogen that can cause mild to

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2. APTAMER TECHNOLOGY FOR THE DETECTION OF FOODBORNE PATHOGENS AND TOXINS

life-threatening disease [54]. Staphylococcal food poisoning (SFP) is one of the most common FBDs and is caused by the ingestion of staphylococcal enterotoxins (SEs) preformed in food by enterotoxigenic strains of S. aureus causes mild problems (like skin infections) to severe conditions like pneumonia, endocarditis, and osteomyelitis. The source of contamination can be unpasteurized milk and cheese products, sliced meat, puddings, pastries, and sandwiches. Their toxins act very fast and the patient can develop symptoms within 30 minutes to 6 hours. The symptoms are vomiting, nausea, stomach cramps, and diarrhea. S. aureus infection is not contagious. Although the conditions are not very serious and patients can be cured by their own immunity, sometimes medical advice is advisable. 2.1.2.5 Vibrio spp. Vibrio spp. are Gram-negative bacteria with 12 pathogenic family members. Most of these foodborne infections are caused by V. parahaemolyticus and V. cholerae and, to a lesser, extent by V. vulnificus. Two serogroups of V. cholera (O1 and O139) that produce cholera toxins are associated with the epidemiological features and clinical syndrome of cholera. The other serogroups (non-O1 and non-O131) are nontoxicogenic and do not spread in epidemic form, but do cause mild FBDs [36]. Symptoms include a sudden onset of excessive, painless, watery diarrhea that can rapidly lead to dehydration, acidosis, circulatory collapse, hypoglycemia (especially in children), renal failure, and even death if treatment is not provided quickly.

2.2 DETECTION OF FOODBORNE PATHOGENS There are several steps for food processing before food products enter the consumer market. Contamination with disease-causing pathogens can happen at any stage during food

processing. Although contamination can be reduced through good manufacturing practices, the detection of pathogens is the key to early identification and prevention of foodborne illnesses and so contributes to ensuring that healthy food is provided to consumers. Various approaches ranging from traditional methods to modern biosensors are being used for the detection of pathogens. The routinely used traditional approaches and some advanced tools for pathogen detection are listed in Fig. 2.1.

2.2.1 Conventional Methods A comprehensive list of routinely used or conventional approaches for foodbornepathogen detection is provided in Table 2.2. Culture-based methods are the most reliable and are still regarded as the “gold standard” for the detection of microbial pathogens. Despite these methods being cost-effective and sensitive, they have some serious drawbacks, such as: (1) having a slow-turnaround time, for example, detection of L. monocytogens requires up to 7 days for the result; (2) the difficulty in identification of viable but nonculturable (VBNC) microbial pathogenic strains, which tend to enter the starvation mode of metabolism under stress conditions; and (3) entailing labor-intensive procedures. Immunological detection of foodborne pathogens is based on the antibodies specific for the causing agent’s characteristic antigen or protein. Different types of antibodies are used for immune detection. Based on clonality (clone or clonal population from which an antibody is derived) and variable region of heavy/light chain, the antibodies can be classified as polyclonal, monoclonal, or recombinant antibodies. Antibody-dependent detection procedures have been employed for E. coli, Salmonella spp., L. monocytogens, and others. Although immunological detection can be performed in a relatively shorter time compared to the culture-based method, it also has some

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

FIGURE 2.1 Different methods of foodborne-pathogen detection. TABLE 2.2

Detection of Foodborne Pathogens by Conventional Methods

Method

Pathogen

References

Culture-based

Listeria monocytogenes

[55
57]

Staphylococcus aureus, Salmonella, Coliforms, Escherichia coli, etc.

[58
60]

Campylobacter jejuni

[61]

Yersinia enterocolitica

[62]

E. coli

[63
66]

Salmonella

[67,68]

L. monocytogenes

[69
72]

Staphylococcal enterotoxins

[73
75]

Campylobacter spp.

[76,77]

Fusarium sp.

[78]

S. aureus

[79]

L. monocytogens

[80]

Salmonella

[81]

Bacillus cerus

[82]

Y. enterocolitica

[83]

C. jejuni

[84]

Immunological assay

Polymerase chain reaction (PCR)-based

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2. APTAMER TECHNOLOGY FOR THE DETECTION OF FOODBORNE PATHOGENS AND TOXINS

drawbacks, for example: (1) lacks the ability to detect the pathogen in real time; (2) low affinity of the antibody; (3) interference from other contaminants; (4) low sensitivity of the assays; (5) high batch-to-batch variation among polyclonal antibodies; and (6) the requirement of a cell-culture facility for monoclonal or recombinant antibodies. Another conventional method is the application of polymerase chain reaction (PCR) for the detection of specific DNA sequences from the genome of the food-contaminating bacteria. Owing to its high sensitivity and specificity, PCR-based detection methods are most widely used for the detection of pathogens in the food. For example, a PCR-based method is the most sensitive and reliable approach for the detection of Salmonella in sea food [85]. In multiplex PCR assays, many pathogens can be detected simultaneously by using different primer sets specific for the different genes of each pathogen. Despite high sensitivity, PCRbased methods are prone to contamination, which may lead to false-positive results. Furthermore, performing PCR directly with the food matrix also poses a challenge as the food matrix itself inhibits the PCR that limits

the applicability of PCR-based approaches. Some other demerits of the applicability of PCR are: (1) relatively expensive; (2) requires skilled technicians to operate and analyze the output data; (3) lack of portability; and (4) the analyte must be concentrated enough to have sufficient numbers of template DNA. The specificity and sensitivity of conventional methods can be enhanced by combining approaches. For example, L. monocytogens can be detected with more specificity by combining immunomagnetic separation with flow cytometry [69]. Similarly, staphylococcal enterotoxin B (SEB) produced from S. aureus can be detected by combining immunomagnetic separation with mass spectrometry [86]. A list of the routinely used methods is provided in Table 2.3. All these conventional methods mentioned have some limitations, including: (1) laborious and tedious procedures; and (2) dependency on trained technicians for operating sophisticated instruments. In other words, these methods cannot be performed in resource-limited settings and are, therefore, unsuitable for point-of-care assessments for foodborne pathogen detection. These limitations could be overcome with the use of biosensors, which are

TABLE 2.3 Comparison of Conventional Methods for Specificity, Reliability, Detection Limit, and Related Factors Particulars

Culture Methods

Immunological Methods

PCR-Based Approaches

Distinguishing feature

Pathogen can be grown in differential culture media

Depend on antigen
antibody binding

Based on nucleic acid of pathogen

Specificity comparison

Very specific test

Specific

More specific than antibodybased methods

What can be detected Pathogens only

Pathogen and their toxins

Pathogens only

Optimal analysis time required

24 h to 7 days

B24 h (up to 5 days in C. jejuni detection)

B6 h

Detection limit (cfu/mL)

103
104

103
105

10
103

Cost

Cost effective

Costly

Costly

cfu 5 colony forming units. cfu is a unit used to estimate the number of viable pathogens in the test sample.

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

2.3 BIOSENSORS

easy to handle, available at low cost for in situ real-time monitoring, and do not require highly trained personnel for its operation.

2.3 BIOSENSORS Biosensors are advanced detections systems that resolve most of the limitations of conventional methods. A biosensor can be defined as an analytical device that converts a biological response into a measurable electrical signal. Typically a biosensor consists of three parts, namely a biorecognition element (BRE), transducer component, and the electronic system (which includes a signal amplifier, processor, and electronic display). The different elements of a biosensor are shown in Fig. 2.2. The BRE, often called a bioreceptor, is a biologically derived molecular recognition molecule and interacts with the analyte of interest. The BREs can be antibodies, aptamers or single-stranded DNA, phages, enzymes, cells, tissues, microorganisms, organelle, nucleic acid, biomimics, or a specific protein. The transducer converts the analyte
bioreceptor interaction into some measurable signal which is proportional to the

53

analyte concentration in the sample. Second, the transducing element can be optical, electrochemical, thermometric, piezoelectric, magnetic, or others, depending on the output signal they convert [87,88]. Third, the electronic system amplifies the transduced signal, processes it, and displays the output result. Biosensors have been widely used for analytical problems and diagnoses in several streams, such as medicine, agriculture, security and defense, environmental monitoring, rapid clinical analysis, and the food industry.

2.3.1 Ideal Features of a Biosensor The ideal characteristics of a biosensor have been summarized in Table 2.4. Biosensors should be rapid with the ability to deliver results in real time. Also, it should give a high signal to noise ratio and should be able to discriminate between the analyte and nontarget background molecules with high precision. It should be able to detect an analyte in a trace amount. The calibration curve should have a wider and dynamic linear range that will give the output result with linear or logarithmic

FIGURE 2.2 Schematic diagram of a biosensor. The three parts of a biosensor. (A) Biorecognition element composed of antibody, cell, protein, and aptamers that interact with the analyte in the sample. (B) Transducer, a physicochemical element, which converts the signal from the analyte
bioreceptor interaction into a measurable signal. (C) Electronic system that amplifies the signal from the transducer, processes it, and display the output as a waveform or digital value.

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2. APTAMER TECHNOLOGY FOR THE DETECTION OF FOODBORNE PATHOGENS AND TOXINS

TABLE 2.4 Characteristic Features of a Biosensor Feature

Required Characteristics

Sensitivity

Should be able to detect a pathogen at very low infective dosage (as low as possible)

Linearity

The linear range of calibration curve should be wide, that is, it should give a linear result within a large concentration range of analyte

Specificity

Should be able to discriminate target from nontarget analyte or background; should have a high signal/noise ratio

Reproducibility

The result should be reproducible without failure

Precision and Accuracy

Should have no false-negative and false-positive results

Robustness

Should have robustness against mechanical and chemical injuries

Time for assay

Analysis time should be as minimal to give results in real time

Ease of use

It should be programmed in such a way that it would be easy to handle and not require trained personnel

Compatibility

It should be compatible with other devices to make the analysis more widespread and validated

Cost-effectiveness

Being cost-effective is required, but sensitivity and specificity should not be compromised

values of the analyte concentration. It should be very robust against mechanical and chemical injuries so that they can be used easily in field applications and the setup can be transferred from one place to another for monitoring during field applications. It should be compatible with other devices to make the further analyses more widespread and useable. Additionally, it should be cost-effective.

2.3.2 Merits of Biosensors Over Conventional Methods Biosensors are highly sensitive in foodbornepathogen detection and have a lower detection limit compared with the conventional methods. Biosensors are selective and quick while conventional methods take long time to get the result, for example, the detection of L. monocytogens using conventional methods requires 7 days to get the result. Biosensors are built with easy programs and simple instruments that do not require highly trained professionals to

operate them, while conventional methods require trained technicians to operate the culture-based and immunology-based experiments to detect pathogens. Conventional methods based on culture methods cannot detect the presence of VBNC pathogens while biosensors are able to detect VBNC pathogens as they are based on the physical features of pathogens. The high sensitivity and convenience of biosensors make them ideal testing instruments for foodborne pathogen detection at the point-ofcare or when there are concerns in the field.

2.3.3 Types of Biosensors Biosensors can be divided into different groups depending on the BRE and transducer used. The full classification of current biosensors is provided in Fig. 2.1. Biorecognition elements interact with the analyte in the test sample. It can be categorized into different groups, namely enzyme, antibody/antigen, cellular structures/cell, biomimetic and bacteriophages, and nucleic acid sequences

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2.3 BIOSENSORS

(aptamers). In recent years, much effort has been focused to develop aptamer (singlestranded DNA or RNA sequence)-based biosensors for foodborne-pathogen detection that are more precise, specific, and give output results within a short time, which corresponds to the main concerns in the food processing industry.

2.3.4 Aptamers: Definition and Features Aptamers are a kind of affinity reagent like antibodies. Aptamers are synthetic singlestranded DNA or RNA sequences that can fold into two-dimensional (2D) and threedimensional (3D) structures and can bind with certain targets [89
99]. Aptamers are usually from 25 to 90 nucleotide bases in length and can fold onto themselves to form a variety of secondary structures including, but not limited to, stems and loops, internal loops, purine-rich bulges, hairpins, pseudoknots, and Gquadruplexes to interact with their cognate target [93,100]. Owing to higher surface density and less spatial blocking in 2D or 3D structures, aptamers have high-binding efficiency with their targets [101]. Owing to their nucleic acid nature, aptamers are structurally and functionally stable across a wide range of temperatures and storage conditions. Aptamers can be chemically synthesized, unlike antibodies which require biological systems to be produced. Aptamers are chemically stable in the pH range of 2
12 and have certain thermal refolding capabilities [102]. Another advantage of aptamers is that they can be chemically modified according to the requirement of the detection for the target molecule. Their Kd value is 10212
1026 M, with most being in the range of (1
10) 1029 M, which shows their high-binding affinity with the target [103]. Aptamers can bind to a range of targets, small molecules like nucleotide sequences, dyes, amino acids, and metal ions as well as to large targets like cellular proteins or even whole

55

cells. All these unique characteristics of aptamers make them suitable for the development of fast and efficient point-of-care assay systems for foodborne-pathogen detection.

2.3.5 Selection of Aptamers: Systematic Evolution of Ligands by EXponential Enrichment Method Aptamers can be selected by a selection method called Systematic Evolution of Ligands by EXponential enrichment (SELEX) where a pool of nucleic acid molecules are subjected to bind to the desired target and the bound molecules after separation are amplified to generate new, more-specific molecules that share a common functional property [104,105]. For selection using the conventional SELEX approach, an in vitro synthesized DNA or RNA library (of a size of around 1015) is required. The library must consist of oligonucleotides containing randomized 20
60 nucleotides flanked by constant primer binding sequences on both sides for PCR amplification. There are three steps in the SELEX process: (1) incubation of the large in vitro synthesized DNA or RNA library with the target; (2) separation of aptamer-target complexes from unbound oligonucleotides; and (3) amplification of the target-bound sequence to be used as a template for the next PCR amplification (Fig. 2.3). After around 10
15 rounds of selection, aptamer sequences with high affinity and selectivity for the target can be achieved. After selection, the aptamer sequence will be cloned and sequenced, and the minimum consensus sequence should be determined that can bind the target without compromising its affinity and selectivity. For selecting an RNA aptamer from the library two additional steps are required compared to DNA aptamer process. In the first step DNA library should be transcribed to get the RNA library and in the second step reverse transcription of bound RNA

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2. APTAMER TECHNOLOGY FOR THE DETECTION OF FOODBORNE PATHOGENS AND TOXINS

FIGURE 2.3 SELEX: A method of aptamer production. (A) General scheme of the SELEX method for DNA and (B) RNA libraries. Source: Adapted from Crit. Rev. Microbiol. 42 (6) (2016) 847
865 with permission from Taylor and Francis.

to get cDNA followed by transcription to get RNA library for further selection. RNA aptamers, being more sensitive to cellular nucleases, can be chemically modified to protect it while DNA aptamers are rather more stable [106]. Full details for the chemical modifications to improve the stability of aptamers have been summarized by other research groups [107]. Compared to DNA aptamers, RNA aptamers can form a much larger variety of secondary structures [108]. Other SELEX methods also have been developed like whole-cell SELEX (Fig. 2.4) and genomic SELEX (Fig. 2.5). In whole-cell SELEX, aptamers specific to the target cell can be developed, for example, the aptamer can bind with the target cell
surface protein and can be enriched in further steps. As shown in Fig. 2.4, the bacterial cells were incubated with the chemically synthesized DNA library where some aptamers can attach to the target cells. Then the aptamers were detached and used as a template to get a highly specific aptamer.

Aptamers can be generated at the nucleic acid level as well where a genomic DNA library will be used, unlike conventional approaches which are based on in vitro synthesized DNA/ RNA. In genomic SELEX, the diversity of the initial library is decreased. There are other novel methods of SELEX like “artificially expanded genetic information systems-SELEX (AEGIS-SELEX)” and “Aptamer Selection Express (ASExp) SELEX” [109,110].

2.4 CLASSIFICATION OF APTAMER-BASED BIOSENSORS Transducers in biosensors play a very important role in the detection process. Aptamer-based biosensors (also known as aptasensors) can be classified on the basis of the transducer element that converts the biological signals from the analyte-biorecognition element (aptamer) reaction into physical, chemical, electrical, or optical changes through

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FIGURE 2.4 Whole-cell SELEX. Schematic showing the aptamer selection against live bacterial cells using whole-cell SELEX. Source: Adapted from Anal. Method. 7 (2015) 6339 with permission from Royal Society of Chemistry.

sensor technology to detect the analyte in the test sample [111]. There have been major research developments over the past two decades for improving the quality and sensitivity of transduction mechanisms. The three transduction mechanisms most commonly used in biosensors are optical, electrochemical, and piezoelectric. These different transduction methods can further be classified as labeled or label-free aptasensors. The different types of aptasensors based on their different transduction mechanisms are listed in Table 2.5.

biochemical interaction into a measurable signal. Most of the biosensors used in the food industry for the detection of foodborne pathogens are optical biosensors. Owing to their high sensitivity and selectivity, optical biosensors gained huge interest for the detection of foodborne pathogens. They can be further classified based on the different transducers used, for example, surface plasmon resonance (SPR) aptasensors, surface-enhanced Raman aptasensors (SERS), chemiluminescence aptasensor, and fluorescent aptasensors.

2.4.1 Optical Biosensors

2.4.1.1 Surface Plasmon Resonance Aptasensors

Optical biosensors rely on optical principles for the transduction of the analyte
BRE

When an incident light beam strikes a surface at a specific angle, the SPR phenomenon

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2. APTAMER TECHNOLOGY FOR THE DETECTION OF FOODBORNE PATHOGENS AND TOXINS

Target organism

Isolation of genomic DNA Transcription

RNA library

2

1 High-throughput sequencing and mapping

3

Counter selection unspecific binding of aptamers

Amplification

6 Target immobilized on matrix Reverse transcription

4 Elution Binding

5 Unbound sequences in flow-through

FIGURE 2.5 Illustration of genomic SELEX. (1) Genomic DNA is isolated. (2) The library is transcribed into RNA. (3) A counter selection against immobilization matrix steps. (4) The library of sequences is incubated with the target. (5) The bound RNA is subjected to reverse transcription into cDNA. (6) The target-bound sequences used as the pattern for the next PCR amplification. Source: Adapted from Biomed. Pharmacother. 93 (2017) 737
745 with permission from Elsevier.

occurs where there is graded reduction in intensity of the reflected light according to the thickness of the molecular layer at the metal surface. The binding of aptamer with the analyte causes displacement near the thin filmmetal surface which leads to change in refrance index and SPR-based biosensors measure the changes in refractive index [132]. Lautner et al. immobilized the aptamers of apple stem pitting virus (ASPV) on the surface of a chip by the

self-assembly method based on biotin-avidin linkers to detect the ASPV in the plant extract food samples [133]. In another work, Ahn et al. demonstrated the detection of V. paraheamolyticus using SPR-based aptasensors [117]. 2.4.1.2 Surface-Enhanced Raman Spectroscopy Aptasensors When light passes through a substance, most of the photons are elastically scattered with the

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2.4 CLASSIFICATION OF APTAMER-BASED BIOSENSORS

TABLE 2.5 List of Aptamer-Based Biosensors for Foodborne Pathogen Detection Transduction Mechanism

Pathogen/Toxin (Analyte)

Limit of Detection

References

Immobilized aptamers on the surface of chip by self-assembly method based on biotin-avidin linker




[108]

Milk

Biotin-tagged aptamer immobilized onto a gold surface by affinity capture

nanomolar range (20.5 nM)

[112]

Vibrio parahaemolyticus

Various food sample

50 -Modified biotinylated DNA aptamer immobilized on SA chip




[109]

Lysozyme

Wine

Biotinylated aptamer immobilized by high affinity capture on a gold layer

2.4 nM

[113]

Ochratoxin-A

Wine and peanut Streptavidin immobilized on thesensor 0.005 ng/mL oil chip surface capture biotinylated aptamer

[114]

Salmonella typhimurium

Food

GNP functionalized with aptamer: (1) for 15 cfu/mL the capture and X-rhodamine (ROX)modified aptamer; (2) used as recognition element and Raman reporter

[111,115]

Ricin B toxin

Oil, juice, milk

Silver nanoparticles labeled with 4,40 0.32 fM bipyridyl (Bpy, Raman reporter) and ricin B aptamer

[116]




Electrochemical sandwich immunosensor utilizing lateral flow system

8 cfu/mL

[117]

B. cerus

Various food samples

Direct-charge transfer conductometric biosensor based on lateral capillary flow

35
88 cfu/mL

[118]

Staphylococcus aureus




Potentiometry using SWCNT

800 cfu/mL

[119]

Staphylococcus aureus




Primary aptamer immobilized on MB using biotin-streptavidin and secondary aptamer attached with GNP

1 cfu/mL

[120]

Sample

Methodology

ASPV

Plant extract food sample

Bovine catalase

OPTICAL APTASENSORS SPR

SERS based

ELECTROCHEMICAL APTASENSORS Conductometric E. coli O157:H7 and Salmonella spp.

Potentiometric

E. coli O157:H7




Light-addressable potentiometric

12 cfu/mL

[121]

Amperometric

Vibrio paraheamolyticus




Rolling circle amplification-based assay

10 cfu/mL

[122]

E. coli O157:H7




Bienzyme electrochemical-based assay

15 cfu/mL

[123]

Impedimetric

Staphylococcus aureus

Hospitalacquired infection as well as toxins in foods

Thiol-modified protein A-binding aptamer was co-immobilized with 6-mercapto-1-hexanol onto gold electrodes by self-assembly

10 cfu/mL

[124]

(Continued)

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2. APTAMER TECHNOLOGY FOR THE DETECTION OF FOODBORNE PATHOGENS AND TOXINS

TABLE 2.5 (Continued) Transduction Mechanism

Pathogen/Toxin (Analyte)

Sample

Methodology

Limit of Detection

References

Ochratoxin-A

Milk products

Covalent immobilization of OTA aptamers on SPE

0.15 ng/mL

[125]

Salmonella typhimurium

Meat, egg, milk, beverages

Combination of poly [pyrrole-co-33 cfu/mL carboxyl-pyrrole] copolymer and aptamer used for the detection

[126]

Aflatoxin M1

Milk

A hexaethyleneglycol-modified 21-mer oligonucleotide was immobilized on a carbon screen-printed electrode through carbodiimide immobilization, after diazonium activation of the sensing surface

1.15 ng/L

[127]

Avian influenza virus




Specific AIV H5N1 aptamer was immobilized through biotin and streptavidin conjugation onto the gold surface of QCM sensor to capture the target virus

1HAU

[128]

Streptomycin

Milk and other food products

Based on flow injection analysiselectrochemical quartz crystal nanobalance (FIA-EQCN) technique

0.3 ng/mL

[129]

Gliadin

Food products

Immobilized directly on the AuNPmodified surface by crosslinking

8 ppb

[130,131]

MASS SENSITIVE BIOSENSORS QCM

QCM, Quartz crystal microbalance; SWCNT, sandwich carbon nanotube; ppb, parts per billion; SPR, surface plasmon resonance; ASPV, apple stem hole virus; SPE, screen-printed electrodes. HAU, Haemagglutinin unit (One HA unit in the haemagglutinin titration is the minimum amount of virus that will cause complete agglutination of the red blood cells).

same frequency and energy while only a few photons scatter with a lower frequency. This effect is termed the Raman effect or Raman scattering. Raman scattering can occur with a change in vibrational, rotational, or electronic energy of a molecule. There has been an upsurge in using Raman spectroscopy for the rapid detection of pathogens. Raman scattering is the more frequently reported whole-organism fingerprinting technique. However, there is a need to culture the microorganism to produce sufficient biomass for analysis by Raman spectroscopy. SERS enhances the signal up to 14
15 orders of magnitude. From the past few years SERS is gaining much interest as it is able to detect pathogens with high precision and specificity [118] (Fig. 2.6).

Chang et al. reported an aptamer-based whole-cell biosensor for the detection of S. aureus (a well-known contaminant in food samples causing more than 500,000 infections in the United States alone) [122]. The major sources are milk and cheese products. S. aureus produces around seven different toxins which are often responsible of food poisoning. Although SFP is usually not life-threatening, detection is better to avoid an outbreak. Chang et al. conjugated DNA aptamer with gold nanoparticles (GNPs) for the detection of S. aureus. It can be detected directly where aptamer
GNPs were incubated with S. aureus cells and the bound aptamer
GNPs were eluted and their light-scattering signals were

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FIGURE 2.6 SERS bacterial detection based on aptamer-dependent in situ formations of silver nanoparticles (AgNPs). Source: Adapted from Anal. Chem. 89 (2017) 9836
9842 with permission from American Chemical Society.

analyzed. In the bead-based amplification method biotin-aptamer 1 (SA61-aptamers) conjugated on 60-nm GNPs while aptamer 2 (SA17aptamers) conjugated onto magnetic beads. Both aptamers interacted with S. aureus followed by binding with streptavidin (SA)-coated magnetic beads and reporter-GNPs (conjugated with DNA adapters). The bound reporter’s lightscattering signals were then analyzed. In an effort to detect multiple bacteria, Ravindranath et al. reported an SERS aptasensor with aptamers for S. typhimurium and the antibodies for S. aureus, and E. coli O157:H7, which were immobilized on the surface of Au or Ag followed by Raman dye molecules marked on them [134]. This aptasensor can detect multiple bacteria with a lower detection limit 100 cfu/mL. In another report, Grow et al. demonstrated a label-free SERS-based

aptasensor for the detection of L. monocytogens (Gram-positive bacteria) and Gram-negative bacteria, for example, Legionella bacteria, bacillus spores, and Cryptosporidium oocysts. L. monocytogens cause around 300 deaths each year in the United States alone according to the CDC report in 2011. The common sources of contamination are raw animal products such as meat and cheese. Their SERS fingerprint changed significantly under culture conditions known to affect virulence. These research works show that both Raman spectroscopy and SERS have potential as pathogen monitoring platforms. Moreover, Raman spectroscopy is compatible with other methods for pathogen detection such as optical and fluorescent methods. In support of this study, Kalasinsky et al. reported Raman chemical imaging spectroscopy (a combination of Raman spectroscopy,

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fluorescence spectroscopy, and digital imaging) which can detect minimum levels of pathogens without employing any enhancement techniques [121].

2.4.2 Electrochemical Aptasensors After optical aptasensors, electrochemical aptasensor based on electrochemical detection methods are of much importance where a change in the electrical properties of the sensing platform is measured. Based on observed parameters, they can be classified into amperometric (current), potentiometric (potential), impedimetric (impedance), and conductometric (conductance) aptasensors [111]. The major advantages of electrochemical aptasensors are their low cost and ability to work with turbid samples; however, their selectivity and sensitivity are not as precise as optical aptasensors, which are highly selective. The most important and promising feature of electrochemical aptasensors is that they are relatively small and can be easily adapted into chip-based platforms. These features make them a suitable candidate for onsite use for the detection of foodborne pathogens. 2.4.2.1 Conductometric Aptasensors In conductometric detection, the biorecognition event causes a change in the ionic species concentration, which leads to a change in electrical conductivity or conductance that can be measured. Conductometric biosensors are miniature, two-electrode devices where an AC voltage applied across the electrodes causes a current flow. Muhammad-Tahir and Alocilja reported a conductometric biosensor specific for the detection of E. coli O157:H7 and Salmonella spp. [124]. They claimed that the lower limit of detection was approximately 7.9 3 101 cfu/mL within a detection time period of 10 minutes. In another study, Pal et al. reported a direct-charge transfer conductometric biosensor for the detection of

B. cereus in various food samples [135]. The lower detection limit was in the range of 35
88 cfu/mL in the food samples with a detection time of 6 minutes. Short detection time, sensitivity, and ease-of-use make this a likely device for onsite use for the pathogen detection. 2.4.2.2 Amperometric Method Amperometric biosensors detect only the electrochemically active analyte that can be oxidized or reduced by losing or gaining an electron. Teng et al. demonstrated the detection of V. parahaemolyticus using a rolling circle amplification
based amperometric aptasensor [136]. Electrochemical biosensors based on amperometric detection can be coupled with other biosensing techniques, for example, a bienzyme electrochemical biosensor coupled with the immunomagnetic separation technique has shown promise for the detection of E. coli O157:H7 [137]. 2.4.2.3 Potentiometric Detection In potentiometric detection the analyte
BRE interaction converts into a potential signal. This technique allows the detection of very small analyte concentration owing to the generation of a logarithmic concentration response. Not much work has been done on the potentiometric detection-based aptasensors. Only a few biosensors can be reviewed for the detection of foodborne-pathogen detection [138]. 2.4.2.4 Impedimetric Detection Impedimetric detection is a very sensitive technique where impedance change is recorded following analyte
BRE interaction. Recently, research has been focused on impedimetric detection due to its simplicity and accuracy [128,139
142]. An advantage of the impedimetric transduction system is that it provides label-free direct detection of the analyte in comparison to amperometry or potentiometry, but the disadvantage is that its

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2.4 CLASSIFICATION OF APTAMER-BASED BIOSENSORS

detection limit is lower compared with the traditional methods. Labib et al. have developed the aptamer-based viability impedimetric sensor for bacteria (AptaVISens-B) where highly specific DNA aptamers to live S. typhimurium were selected via SELEX and integrated into an impedimetric sensor by self-assembly onto a gold nanoparticle-modified screen-printed carbon electrode (GNP-SPCE). They found the aptasensor to be highly selective, could detect 600 cfu/mL, and discriminate from other closely related subspecies of Salmonella (Fig. 2.7). In another report, Rivas et al. have developed an impedimetric aptasensor for the detection of Ochratoxin-A (OTA) using iridium oxide nanoparticles with a very low limit of detection (14 pM) [136]. In this study, the researchers used SPCE modified with polythionine and iridium oxide nanoparticles (IrO2 NPs) to detect OTA (Fig. 2.8). The aminated aptamer selective to OTA is exchanged with the citrate ions surrounding IrO2 NPs via electrostatic interactions with the same surface.

63

2.4.3 Mass-Sensitive Biosensors This is a very sensitive detection system where transduction is based on small changes in mass. Piezoelectric crystals are the core of mass-sensitive biosensors. Quartz crystal is a piezoelectric crystal. Piezoelectric crystals vibrate at a specific frequency when an electrical signal is being applied to it. The oscillation frequency of piezoelectric crystal is directly proportional to the applied frequency as well as to the mass of crystal. Further, ligand binding leads to an increase in mass that ultimately changes the oscillating frequency of the crystal. Mass-based aptasensors can be categorized into two broad groups, namely: (1) bulk wave or quartz crystal microbalance (QCM); and (2) surface acoustic wave. Brockman et al. showed the detection of avian influenza virus using quartz crystal microbalance
based aptasensors [143]. In this work, they were able to detect AIV H5N1 using specific DNA aptamers as the biosensing material that was immobilized through biotin and SA conjugation onto the gold surface of QCM.

FIGURE 2.7 Schematic diagram of the aptamer-mediated electrochemical detection of live Salmonella typhimurium bacteria. Source: Adapted from Anal. Chem. 84 (2012) 8966
8969 with permission from American Chemical Society.

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FIGURE 2.8 Schematic illustration of the fabrication steps and working principles of the developed impedimetric aptasensor for ochratoxin-A (OTA) detection. Source: Adapted from Anal. Chem. 87 (2015) 5167
5172 with permission from American Chemical Society.

2.4.3.1 Future Prospects of Biosensors In recent times, aptamer has become a synonym for a biosensor platform for the quantification and detection of various analytes used in the biological system. Aptamers emerged as strong candidates to be used as biorecognition elements in aptamer-based biosensors and could potentially replace its antibody counterpart for point-of-care detection purposes owing to their high stability, specificity, and affinity toward their targets. Moreover, the ease with which the aptamers can be modified is commendable as researchers around the world have exploited this molecule in designing a variety of aptamer-based biosensors such as optical aptasensors, electrical aptasensors, and mass (piezoelectric) aptasensors for the detection of various targets ranging from small molecules to whole cells. Moreover, no one can deny that the progress made in the field of aptamer technology over the past few years has been phenomenal, whether in designing

and introducing new strategies in SELEX like next-generation sequencing or capillary electrophoresis which further provide a great opportunity to scientists in selecting a highly selective and affine aptamer sequence against the target of interest. However, there is still scope to further improve the sensitivity of aptamer-based sensors using modern technologies like paper-based microfluidics. This adaptation of aptamers to simple paper-based assays will certainly make this technology popular among end users.

2.5 CONCLUSION In this chapter, we summarized and showed that a wide variety of aptamer-based biosensors are extensively used for the detection of a broad range of foodborne pathogens and their toxins. As compared to other biosensors, electrochemical aptasensors have been preferably used for

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REFERENCES

the detection and quantification of foodborne contaminants due to their high sensitivity, rapidness, and that they can be used in a label-free approach which further makes things simpler.

COMPETING FINANCIAL INTEREST The authors declare no competing financial interests.

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C H A P T E R

3 Biosensors for Rapid Detection of Breast Cancer Biomarkers AC Pereira1, MGF Sales1,2 and LR Rodrigues1,* 1

Centre of Biological Engineering, Minho University, Braga, Portugal, 2BIOMARK/ISEP, Superior Institute of Engineering of Porto, Porto, Portugal

3.1 INTRODUCTION

expectancy, increased exposure to risk factors, and poor lifestyle as well as genetics all contribute to this number [1
4]. Considering the genotype and phenotype diversity, it is expected that each individual may have different degrees of susceptibility in developing breast cancer. However, considering the population in a global perspective, several risks factors have been identified, such as age, pregnancies, family history or genetics, geographical variation, and lifestyle (Fig. 3.1).

Breast cancer epidemiology has been progressively increasing each year. In 2008, breast cancer represented around 11% of all types of cancer worldwide and by 2012 that number increased to 12% [1]. According to the American Cancer Society [2], in 2015 approximately 40,290 women and 440 men were expected to die from breast cancer and over 290,000 new cases of breast cancer were expected to be diagnosed. With roughly one million new cases worldwide each year, breast cancer is the most common form of malignant cancer, comprising 18% of all female cancers [3]. Statistically, 17 out of a 1000 women before or at the age of 50 will have had some form of breast cancer diagnosed, giving this disease a prevalence of just under 2% [3]. Also, it is expected that female breast cancer incidence will reach 3.2 million new cases per year by 2050 [1]. The growing number of cases cannot be attributed to one single factor. Average lifetime

3.1.1 Breast Cancer Epidemiology 3.1.1.1 Risk Factors for Breast Cancer There are several interesting aspects when considering age as a risk factor. Both menarche and menopause ages seem to correlate with breast cancer
risk [1
5]. An early first menstruation increases the risk of breast cancer and for every year of delay of the menarche age this risk lowers by 5% [2,3]. At

*Corresponding author.

Advanced Biosensors for Health Care Applications DOI: https://doi.org/10.1016/B978-0-12-815743-5.00003-2

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3. BIOSENSORS FOR RAPID DETECTION OF BREAST CANCER BIOMARKERS

FIGURE 3.1 breast cancer susceptibility.

menopause, the opposite happens, with late menopause aggravating risk, which increases by about 3% for each year older at menopause [2,3]. In other words, the longer a woman’s fertile period is, the higher her risk for getting breast cancer [1,2,3,5]. Another important risk factor is with regard to pregnancies, in particular the age of the first full-term pregnancy and the number of pregnancies. Women who have had their first child before the age of 20 have half the risk of developing breast cancer compared to women who got pregnant after the age of 30 [3,5]. For women who have their first-born child after the age of 35, the risk is the highest. In fact, these women have an even higher risk than nulliparous women, whose risk is about 25% higher than for women who have had at least one child [3,5]. Multiple pregnancies seem to further reduce the risk of breast cancer [1,5]. Geographical variation and lifestyle also play an important role as risk factors in breast cancer development, with Western countries women showing breast cancer incidence five times higher than the Far Eastern ones, mainly due to stress, pollution, and poor diet prevail [1
5]. Interestingly, studies regarding migrants from Japan to Hawaii showed that, by

Risk factors for development

embrassing the lifestyle of the host country, the migrants’ risk in breast cancer would become consistent with the risk seen in the host country, within a time period of one or two generations [3,5]. This is a strong indication that environmental factors are strong contributors to the disease’s prevalence. Nevertheless, studies show that breast cancer is dramatically increasing in South America, Africa, and Asia [4]. Regarding lifestyle, alcohol consumption also seems to have an impact, increasing the risk of breast cancer roughly by about 10% per 10 g alcohol (1 unit) consumed per day [2,5]. Obesity seems to have a 1.5
2 times increased effect in developing postmenopausal breast cancer [2,3,5]. This correlation is probably due to the increased levels of estrogen inherent in the excess of fat tissue in overweigh and obese women. Regarding tobacco, diet (in particular animal fat intake), and physical activity, results are, at best, suggestive but not conclusive, that a diet rich in vegetables and low in animal fat and sugar as well as no smoking and moderate exercise seems to help reduce the risk [2,3,5
6]. Family history and genetics entail not only shared genes, but also shared environment and lifestyle, with 10% of breast cancer cases in Western countries being attributed to

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

3.1 INTRODUCTION

genetic predisposition, generally inherited as an autosomal dominant with limited penetrance [3]. Globally, studies regarding these parameters show that the risk doubles for firstdegree relatives of affected patients [2]. There are mutations that have been identified as predisposing factors to breast cancer, including the ones located in genes BRCA1 (Breast Cancer type 1), BRCA2 (Breast Cancer type 2), tumorsuppressor protein p53, PTEN (Phosphatase and Tensin homolog), and ATM (AtaxiaTelangiectasia Mutated), with the first two carrying higher risks [2,3,5]. There are studies that correlate that the probability of developing breast cancer among carriers of these mutations also vary geographically, again suggesting that environment may be important in the expression of those mutated genes [1,5]. Women with previous benign breast cancer usually exhibit a higher risk of developing malignant breast cancer, being this risk about four to five times higher compared to women that never had proliferative changes in breast tissue [2,3]. For this reason, women who have had atypical epithelial hyperplasia should check regularly for abnormal tissue growth to guarantee that in the event of malignant breast cancer development it will be diagnosed in an early stage [2].

3.1.2 Breast Cancer Types For the majority of breast cancers, the disease begins at the milk production glands, called lobules, and ducts that connect the lobules to the nipple [1,7]. The great challenge of achieving an effective cancer diagnosis is mainly due to breast tumor’s heterogeneity [8]. The additional genetic and epigenetic alterations create different clonal populations, further increasing intratumor heterogeneity [8]. Therefore despite common association of breast cancer as a type of cancer, this pathology actually englobes a group of different cancers affecting the breast from where the tumor originates [2].

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There are different types of classification for the different breast cancer subtypes, depending on the focus. Regarding the origin of the tumor and cellular behavior, they can be considered in situ or invasive [2]. If the classification is attributed according to gene expression, then the subtypes to be considered are luminal A, luminal B, luminal B-like, HER2 positive nonluminal, and triple-negative breast cancer (TNBC) [2]. These classifications follow the guidelines and definitions provided by the American Cancer Society, the Canadian Cancer Society, and BreastCancer.org. 3.1.2.1 Noninvasive or In Situ Ductal carcinoma in situ (DCIS) is the most common type of breast cancer in the noninvasive category. It is located only in the lining of the milk ducts and does not invade the walls of the ducts into the tissue of the breast or metastasize to lymph nodes or other parts of the body. It is asymptomatic and cannot be detected by palpation. For this type of breast cancer, mammography is the best exam and treatment is usually the removal of the lump, followed by radiation or hormone therapy in the cases where DCIS is hormone receptorpositive. DCIS is a stage 0 breast cancer, with high successful rates of treatment. However, it can recur, which increases the possibility of future invasive breast cancer development. In these cases, DCIS is graded higher than zero. Lobular carcinoma in situ (LCIS) is represented by the development of abnormal cells in the lobules, without invading the wall of the lobules metastasizing. It rarely develops a lump or a morphological change in the tissues. For this reason, a biopsy is the preferable approach for diagnosis in order to evaluate the presence of other possible breast changes. It is a nonlife-threatening indicator of increased risk of invasive breast cancer development later on and many cases of this type of breast cancer go undiagnosed, without ever causing any health problems.

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Although LCIS rarely develops into invasive breast cancer, patients are usually followed up with regular exams, including mammography and, in some cases, magnetic resonance imaging (MRI). Hormonal therapy is sometimes recommended as a preventive measure when the LCIS is hormone receptor-positive. 3.1.2.2 Invasive Breast Cancer Invasive ductal carcinoma (IDC) is the most common (i.e., 80% of cases) type of invasive breast cancer. It starts in a duct, breaking through its wall and invading the surrounding tissue of the breast. From there it can metastasize to the lymph nodes and other parts of the body. Symptoms of IDC can include a lump in the breast or armpit area, redness, thickening, irritation or dimpling of the skin, persistent breast pain and/or swelling, and nipple changes and discharge. Its detection is usually done by mammography and the confirmation of the diagnosis is obtained by the performance of mammograms and/or further ultrasounds analysis and biopsies that will allow evaluating the hormone receptor HER-2 expression. The treatment of IDC usually involves a lumpectomy or mastectomy, depending on the location and extension. Lumpectomy usually is followed by radiation therapy but, depending on the tumors’ receptor status, other treatments are also frequently used, including chemotherapy, hormone therapy, and HER-2 targeted therapies. Invasive lobular carcinoma (ILC) represents roughly 10% of invasive breast cancers and starts in the lobules of the breast, eventually breaking through the lobules, and invading the surrounding tissue of the breast. From there it may metastasize to the lymph nodes or other parts of the body. This invasive type of breast cancer does not form a distinct lump or a precise location within the breast and since cells grow in the form of a line instead of a lump, its detection can be difficult by mammogram. Therefore a biopsy is usually more suitable for detection, although other techniques like ultrasound or MRI may be used.

Treatments include surgery, radiation, chemotherapy, and HER-2 targeted therapy. Inflammatory breast cancer (IBC) is a rather uncommon type of invasive breast cancer accounting for about 1%
3% of all breast cancers, being more prevalent in younger women and women with African ancestry. In early stages, IBC is often mistaken as infection and symptoms include breast redness, swelling, and pain, skin that feels warm to the touch and/or with an orange peel
like skin texture, and changes in the appearance of the nipple. It has a high spreading rate and is therefore more aggressive. Regarding localization, IBC starts at the milk ducts of the breast and spreads to the lymph vessels. Because symptoms are not usually associated with breast cancer, its diagnosis is rather difficult. Patients with IBC have up to 35% chance of distant metastases. Therefore when IBC is suspected, ultrasound, MRI, and a biopsy, besides a mammogram, are often needed to detect or confirm the diagnosis. Treatments involve a combination of approaches including chemotherapy, surgery, radiation, and hormone and HER-2 targeted therapies, if appropriate. Paget’s disease of the nipple is another form of cancer that accounts for less than 5% of all breast cancers and is more frequent in women over 50 years of age. In this type, cancer cells are located primarily in and around the nipple, but can spread to the areola and other areas of the breast. Symptoms are usually located on one nipple and may include flattening of the nipple against the breast, persistent itchiness, and scaling of the nipple, aggravating into weeping, crusting, and pain. Approximately half of the patients with Paget’s disease develop another form of breast cancer, like DCIS or IDC. Therefore if Paget’s disease is suspected, additional examinations are conducted like mammograms and ultrasounds for clearance of other possible breast cancer types. Surgery is the first procedure of treatment and further therapy will depend highly on how much breast tissue is removed.

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In 2011 St. Gallen International Expert Consensus proposed a new classification for breast cancer subtypes. This classification was then refined at the 2013 conference, scoping the breast cancer molecular subtypes. It is a five subtype classification model made according to the expression of breast cancer receptors (see Table 3.1). This classification is more clinically relevant than the previous one, because it gives some indication on what hormone therapy, if any, might be the best treatment [2]. In luminal A subtype, estrogen-alpha (ERα) and/or progesterone (PR) receptors are expressed but human epidermal growth factor receptor 2 (HER2) is not [9,10]. Because it has low levels of protein Ki-67, which helps control how fast cancer cells grow, it tends to grow slowly and has the best prognosis [2,9,10]. In the luminal B and luminal B-like subtypes, ERα and/or PR are expressed, but HER2 is only expressed in luminal B-like. In both subtypes, there are high levels of Ki-67 [9,10]. For this reason, luminal B cancers tend to have slightly worse prognosis than luminal A cancers because they generally grow slightly faster [2,9,10]. The HER2 positive nonluminal subtype does not express ERα or PR receptors, being positive only for HER2 receptors. This subtype tends to grow faster than luminal cancers and can have

a worse prognosis [9,10]. However, using targeted therapies like trastuzumab, pertuzumab, lapatinib, and ado-trastuzumab emtansine aimed at the HER2 protein significantly increases the treatment’s success rate [2]. In TNBC, the tumor cells do not possess ERα, PR, or HER-2 receptors [9,10]. This means that hormone therapy is ineffective [2]. It is a high-grade and aggressive form of breast cancer which is usually detected in the advanced stages with frequent metastization in the brain and lungs [2,9]. It also can develop between mammogram screening periods and tends to recur after five years [2]. Regarding treatment, the most common approaches are surgery, chemotherapy, and radiation and the success of the treatment greatly depends on how early the disease diagnosis is confirmed [2].

3.1.3 Biomarkers Early detection is the number-one recommendation for successful treatment. In order to achieve that goal, finding specific biomarkers is essential. Indeed, biomarkers are very helpful for predicting a patient’s treatment outcome, enabling the identification of high-risk cases that are good candidates for adjuvant therapy, as well as determining the best therapy approach to be used. Technology has been

TABLE 3.1 Correlation of Molecular Types of Breast Cancer With Receptors’ Expression of ERα (Estrogen-Alpha), PR (Progesterone), and HER2 (Human Epidermal Growth Factor Receptor 2), Expression Levels of Ki-67 (mStands for High expression Levels and k Stands for Low Expression Levels) and Overall Survival Outcome. Receptors Expression/Levels Type of Cancer

ERα

PR

Luminal A

ü

Luminal B Luminal B-like HER2 positive nonluminal Triple negative-breast cancer

HER2

Ki-67

Prognosis

ü

k

Good

ü

ü

m

Intermediate

ü

ü

ü

m

Poor

ü

k or m

Poor

k or m

Poor

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improving in this context and a diverse panel of new biomarkers have been emerging over the past few years [11]. However, the logistics of using a new biomarker in clinical diagnosis is expensive and time consuming. Analytical validation, clinical validation, and demonstration of the biomarker’s clinical value are required steps before considering a biomarker as such. Regulatory approval and evidence of cost-effectiveness are receiving increased weight in this process in a growing number of countries. Another challenge is the heterogeneity of the disease. By being so vast, no molecular biomarker, in isolation or as part of a panel, is sufficient for breast cancer screening or early detection. Hence in current clinical practice, the established breast cancer biomarkers still serve as a complement to conventional clinical methodology like mammograms and X-rays. These biomarkers include ER and PR receptors, HER2, BRCA1 and BRCA2, Ki-67, cancer antigens CA 15-3 and CA 27.29, carcinoembryonic antigen (CEA), urokinase plasminogen activator (uPA), plasminogen activator inhibitor 1 (PAI-1), and multiparameter assays for gene expression [11,12]. The evaluation of ER, PR, and HER2 allows to distinguish the three principal immunophenotypes (luminal A, luminal B, and TNBC) and these biomarkers possess prognostic relevance and a clearly defined guide systemic treatment [13]. However, in clinical practice, subtype boundaries are sometimes dubious and are therefore pragmatic. For example, lowlevel reactivity for PR alone may not indicate substantial likelihood of endocrine sensitivity, and in context with other attributes such as high-grade and/or proliferation, these patients may justifiably be managed as triple-negative [13]. The most informative biomarker in breast cancer remains the ER as over 75% of tumors are ER-positive, but in these breast cancer subtypes the PR has a diagnostic utility, with low or absent expression being associated with

more proliferative and aggressive tumors and probability of recurrence, whereas higher expression levels of PR indicate a more favorable prognosis and endocrine therapy response [12
14]. HER2 overexpression is less common, with approximately half of the cases being ER negative, and it is an indicator of poor prognosis [13]. However, despite being an indication of increased tumor aggressiveness, it also represents a therapeutic opportunity since this receptor is targeted by monoclonal antibodies trastuzumab and pertuzumab [13,14]. Breast cancer 1 (BRCA1) and breast cancer 2 (BRCA2) genes are tumor-suppressor genes and their respective proteins are responsible for DNA repairing [11,15]. Mutations in these genes are very rare, affecting less than 1% of the population, but, when present, represent an increase risk of breast cancer development [2,5]. In patients carrying BRCA1 mutated genes, the chance of developing breast cancer is around 60% and most likely it will be TNBC, whereas patients carrying BRCA2 mutated genes have 50% probability of developing breast cancer with a phenotype likely similar to ones exhibited by patients without those mutated genes [16]. There are commercially available tests to identify some of these mutations which are used often in clinical practice and results are considered useful as predictive markers (especially for BRCA1) of the response to different types of chemotherapy in patients carrying mutated genes [17]. Another complementary biomarker for breast cancer detection is the nuclear nonhistone protein Ki-67 [11,13]. This protein is well correlated with tumor grade and is inversely associated with the ER status [13]. Moreover, meta-analysis of Ki-67 have shown that high scores of this protein’s proliferation strongly associates with reduced overall survival [18,19]. Cancer antigen 15-3 (CA 15-3) and cancer antigen 27.29 (CA 27.29) are carbohydratecontaining protein antigens called mucins [12]. Both antigens belong to the mucin 1 protein

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family, which, although not fully understood how, seem to reduce the cell-to-cell interaction and inhibit cell lysis [12]. The mucin 1 gene is overexpressed in malignant breast cancer cells and both CA 15-3 and CA 27.29 have clinical relevance in patients with breast cancer [12,20,21]. However, as with all the other biomarkers for breast cancer, they are to be considered, at most, as complementary in diagnosis, particularly for breast tumor recurrence detection and metastasis monitoring during active therapy due to low-organ specificity [12,20]. CEAs belong to a family of cell surface glycoproteins and are widely used as tumor markers for several tumors, including breast cancer, in clinical practice [12,21]. Studies regarding this antigen suggest that it might act as an adhesion molecule, gaining relevance since cell adhesion is involved in cancer invasion and metastasis processes [12]. Therefore in patients with continuous rising levels of CEA, this may explain why their cancer treatment is not being successful [12]. In this way, CEA is used as a clinical staging biomarker by helping to detect recurrence after surgery and monitor therapeutic responses in patients undergoing chemotherapy and radiotherapy [12]. The protease system constituted by the serine protease uPA, its receptor uPAR, and inhibitor PAI-1 is a system implicated in cellular angiogenesis, tissue invasion, and metastasis [12,22,23]. High levels are correlated positively with histological grade and negatively with hormone-receptor status [12]. Its value serves more as a predictive patient outcome biomarker and has been validated for individual patients and random trial analysis [12,22
24]. However, because the application of this complex as a predictive test is limited to fresh or freshly frozen tumor tissue it is not widely used in routine practice [12]. Despite all the challenges inherent to establishing new biomarkers for early detection of breast cancer, there have been some candidates in the past few years that have shown promising results. Although insufficient on their own

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for routine use in clinical practice, cyclin D1, cyclin E, p53, and cathepsin D have shown strong correlations with breast cancer, in particular as prognosis complementary data for metastatic breast tumor. Cyclin D1 is a protein encoded in humans by the CCND1 gene and its overexpression occurs in 50% of all breast cancers [11]. Cyclin D1 expression correlates with Erα as it may potentiate transcription of ERα-regulated genes [11,25,26]. Although overexpression is associated with favorable prognosis, the prognostic utility remains controversial [25
27]. Cyclin E is a protein involved in DNA replication and the replication control events of phase S of the cellular cycle [12,28]. The deregulated expression of cyclin E affects the G1/S transition compromising the DNA replication process and cell cycle progression [12,28]. Although the exact mechanism that relates the cyclin E overexpression and tumor instability is still not fully understood, it is likely that high levels of this protein affect DNA synthesis disturbing several checkpoint systems [11,28]. In about 25% of breast tumors, cyclin E is either overexpressed or abnormally stable and it has been consistently related to poor prognosis and risk of recurrence [11,12,28]. Tumor protein 53 (p53) is a nuclear protein with an important role in cell cycle regulation and keeping conserved stability by preventing mutations and working as a tumor suppressor [29]. In normal cells it is expressed in low levels [11,12]. When high scores are observed, it is usually in abnormal cells [29]. Mutations in the p53 gene are one of the most common genetic abnormalities observed in human tumors and is usually associated with more aggressive forms of cancer when overexpressed [30]. For this reason, it is used in prognostic clinical assessments and has been suggested as a biomarker of resistance to chemotherapy and radiotherapy since these treatments damage DNA, triggering apoptosis via p53 [29,31].

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Cathepsin D is a lysosomal aspartyl endopeptidase with no known endogenous inhibitors and is catalytically active at low pH levels [12,32]. Its biological function is to degrade proteins in the lysosome and help newborns’ development through protective action from intestinal necrosis and thymic apoptosis [33]. Also, several studies have shown that this enzyme is highly regulated by estrogen, growth factors, retinoic acid, and tumor necrosis factor alpha [12]. Cathepsin D overexpression seems to be correlated to metastasis development, even facilitating the process, leading to relapse and decrease of overall survival [12,32]. Because research is in constant development, the availability of epigenetic information has allowed better insight regarding the molecular side of breast cancer [34
36]. The databases that have been created collect all the accumulated information allowing an overall knowledge for researchers in order to keep moving forward toward the discovery of new biomarkers that could make a significant difference in the early detection of cancer. Some of the databases, like The Cancer Genome Atlas (TCGA), Gene Expression Omnibus (GEO), Surveillance, and Epidemiology and End Results (SEER) provide good information on biomarkers for breast cancer, thus improving the strategies for diagnosis and treatments [13,37].

3.1.4 Conventional Detection Methodology Conventional breast cancer methodology relies on techniques that allow imaging and although biomarkers complement the diagnosis, traditional diagnostic tools are still fundamental. These procedures include mammography, MRI, ultrasound, and physical examination [2,38,39]. Mammography is the golden standard procedure for breast cancer screening [2,14,40]. It is a low-dosage X-ray exam that allows

visualization of the breast’s internal tissue structure [2,40]. The procedure is mostly conducted digitally by delivering low doses of radiation and it can be used in triple assessment of breast lumps, skin changes, and nipple thickening or discharge [14,39]. Despite being the golden procedure, mammography has some limitations. For instance, it has high rates of false-positive results [14,40]. According to an American study, in 10 screening examinations, 50% of the women will experience a falsepositive result and 19% of these women will need to undergo biopsies [2]. Another limitation is the radiation, which still concerns many patients despite being a very low-risk radiation level [2,41]. Other limitations include breast tissue density and postmenopausal hormone replacement therapy [2]. Still, mammography is the single-most efficient procedure of early detection, allowing cancer detection several years before physical symptoms develop [2,40]. MRI is another standard procedure but, unlike mammography, uses magnetic fields instead of X-rays [2,39]. This technique allows very detailed, cross-sectional images of the breast using a contrast material, usually gadolinium diethylenetriamine pentaacetic acid (Gd-DTPA), that is injected into the bloodstream before or during the examination, thereby improving the images details [2,42]. Because there is the need of specific and sophisticated equipment to use MRI for breast tissue, not all hospital facilities can perform this exam. Moreover, the cost for an MRI scan is higher than mammography, therefore MRI is usually recommended for women that are suspected of being at high-risk [2]. A breast ultrasound is usually a follow-up examination when abnormal cellular tissue from a screening or diagnostic mammogram or physical exam is detected [2,43]. Studies have shown that ultrasound can be a more suitable procedure than mammography for the detection of breast cancer in dense breast tissue [2,39,44,45]. However, it also comes with

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an increase of false-positive results. Therefore it is not a procedure recommended as a substitute for mammograms, but as a complementary one, with the upside of being noninvasive and safe [2,44,45]. Physical examination, particularly selfexamination, has been recommended for decades to all population groups. For women over 40 years of age periodic selfexaminations for lump screening is highly recommended [2]. It does not replace the exams performed in check-ups, mostly because there are several cases of asymptomatic tumors and lumps are very difficult to detect in dense breast tissue. Nevertheless, self-awareness seems to make women more attentive, improving chances of regular and periodic clinical check-ups [2]. Advances in diagnosis involve the development of tests that enable the screening of multiple genes at once. These tests make use of tumor tissues, either fresh or paraffinembedded and formalin-fixed [34]. This multiple screening greatly contributes to reduce the administration of adjuvant chemotherapy without compromising the outcome [13]. Oncotype DX is a test that analyzes the expression of 21 genes, of which 16 are tumor-related and the remaining 5 are for housekeeping, using quantitative polymerase chain reaction (PCR) for mRNA levels measurement [9,35,46]. MammaPrint is another multianalyses test for breast cancer patients’ outcome prediction [9,13,35,36,46]. This test measures the mRNA levels of 70 genes implicated in the 6 classic cancer hallmarks. Other multianalyte tests include EndoPredict, Prosigna, Genomic Grade Index, and the Breast Cancer Index, among others, with Prosigna and EndoPredict being the most studied [34,36,46]. These two tests can be helpful in prognosis and adjuvant therapy choices for patients with ER-positive, HER2-negative, and metastatic lymph nodes [34,36].

3.2 BIOSENSORS 3.2.1 Overview Biosensors are analytical devices that allow the detection of biological molecules of interest through a physicochemical transducer signal [47,48]. These devices also allow quantification where the signal generated is proportional to the concentration of the analyte [49]. Despite the boom in biosensor devices over the past two decades, the first device of this kind was an electrode for oxygen detection developed in 1956 by Leland C. Clarck, Jr. and known as the Clarck Electrode [48
50]. In 1962 an amperometric enzyme electrode for glucose detection was developed followed by a potentiometric biosensor to detect urea in 1969 [49,51]. The first biosensor to be commercialized was a biosensor for glucose detection in 1975, the same year that the first microbebased immunosensor was developed [48,49,52,53]. The portability of these systems enabled a remarkable progress in the biosensors research field of different areas of science and engineering in research and development that we see nowadays [48,49]. Although there is a wide range of sophistication in biosensors, there are transversal components common to all of them, including a biorecognition element (BRE) that identifies the molecule of interest in the sample and a transducer that converts a positive detection event into a measurable signal. The selectivity of the recognition element and the sensitivity of the transducer is what defines the overall accuracy of the biosensor device. However, the surface in which the recognition element is immobilized and the design of the biosensor as a whole are key features to take into account when developing a functional and reliable device for analytical purposes. Apart from the components, the analytical properties of a given biosensor are key elements for its application. These include

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different parameters, such as selectivity, reproducibility or stability, and sensitivity [49,54]. Selectivity is probably the most important aspect in any biosensor since this is the ability to detect, without bias, a specific analyte of interest [49,54]. Thus it can be expected that the higher the selectivity, the more reliable the results, and the lower the probability of falsepositive or false-negative results [49]. Reproducibility relates to precision and, indirectly, to accuracy. It represents the ability of a biosensor to generate identical results in repeated assays or in different units of the same biosensor [49]. It is the property that provides high reliability [49]. Stability is related to reproducibility and is considered as the susceptibility of the device to the surrounding disturbances [49]. The higher the susceptibility, the bigger the drift in the output response [49]. In other words, stability is related to the precision of a biosensor and when a biosensor is more susceptible to a certain parameter of the surrounding environment, like temperature for example, the higher the probability of errors [49]. Sensitivity is another important property of a biosensor and it is related with the ability to discriminate accurately among close concentration levels of the target molecule [49]. Thus it also determines the accuracy of the result, being it is also interfaced with the limit of detection value [49,54]. In applications such as food safety or healthcare, in which it is necessary, for example, to detect contaminants or antibiotics in food and water or detect biomarkers for specific diseases, the ability to detect accurately amounts of analyte as low as ng/mL or even fg/mL in a sample is crucial to support adequate and reliable decisions [49]. The applicability of these devices is endless, and whether the purpose is health diagnosis or food-quality monitoring, biosensors allow rapid and inexpensive analyses with high reliability [55]. Moreover, because of the portability and user-friendly features, point-of-care

analysis without the need of further equipment becomes possible, and increases the potential of these devices. Biosensors can monitor all kinds of molecules, from toxic metals, organophosphates, and pesticides to glucose, cholesterol urea, and drugs as well as enzymes, antibodies, bacteria, biomass, and fermentation processes with remarkable selectivity and costeffectiveness[49,54,55].

3.2.2 Classification Biosensors can have different classifications depending on the nature of the BRE or the technology used for the transducer. 3.2.2.1 Biorecognition Element The BRE, or bioreceptor, is one of the key components in a well-structured biosensor. The BRE, or bioreceptor, is one of the key components in a well-structured biosensor and its nature varies widely depending on the target molecule to be recognised. Overall, there is no perfect bioreceptor, each kind has advantages and disadvantages and, therefore, its selection should always take into account the analytical method being developed and the application intended. 3.2.2.1.1 ANTIBODIES

Antibodies, due to their specific antibodyantigen binding properties, are perhaps the most popular class of BREs used in biosensors (Fig. 3.2), due to their antibody
antigen binding properties [47]. One of the major advantages of making use of this property to develop immunosensors is that samples containing the antigen do not need purification before analysis [47]. Antibodies are naturally occurring proteins and are either monoclonal or polyclonal [7,56]. During a humoral immune response, an antibody that is derived from a single B-cell clone is termed a monoclonal antibody [57].

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FIGURE 3.2 Schematic illustration of a biosensor with an antibody as a biorecognition element (on the left side of the figure). On the right side of the figure, it is shown the crystal structure of the complex formed by trastuzumab and the human epidermal growth factor receptor 2.

If, on the other hand, it derives from numerous B-cell clones, a polyclonal antibody will be produced [57]. There are advantages and disadvantages in using both types. Monoclonal antibodies can only recognize one specific epitope of an antigen while a polyclonal antibody can detect multiple epitopes, therefore are able to recognize an antigen from different orientations [57]. In terms of biosensing response, this is particularly reflected in sensitivity. Monoclonal antibodies cannot detect the antigen from its different regions, meaning that many antigen molecules arriving at the biosensor surface may not be detected, thereby reducing sensitivity. In contrast, polyclonal antibodies are more likely to detect the different sides of the antigen surface, increasing the number of antigen molecules detected, and yielding increased signals. Considering the reproducibility of batches, producing polyclonal antibodies means dealing with variations from batch to batch, including reactivity and titer, which does not happen in

monoclonal antibody batches [57]. In fact, cultures of B-cell hybridomas offer continuous remarkable specificity, making monoclonal antibodies powerful tools for macromolecules, cell investigations, and clinical diagnosis tests [57]. Additionally, there are recombinant antibodies that consist in genetically manipulated and fused antigenbinding domains of common antibodies [56,58]. These type of antibodies are considered the third generation of antibody treatment and tackle the shortcomings of monoclonal and polyclonal antibodies by triggering a range of effector functions like antigen phagocytosis and toxin neutralization, agglutination, or precipitation [58]. Compared to polyclonal or monoclonal antibodies, recombinant antibodies are less expensive and time consuming and do not suffer from lack of supply, plasma-suitability problems, or risk of infectious agent transmission, thereby improving the safety profile [47,58,59].

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3.2.2.1.2 NUCLEIC ACIDS

Nucleic acids (NA) are BREs that can detect the presence of a complementary target sequence in complex mixtures through complementary hybridization [60]. DNA and RNA follow simple complementarity rules due to their powerful molecular recognition systems, thereby allowing the detection of biomarkers of interest and represent an easy approach in designing specific interactions between different probes and affinity reagents [61,62]. In human samples, the circulating NAs in plasma and serum have implications for minimally invasive diagnostic and predictive applications in benign and malignant conditions [63]. Nucleic acid aptamers (NAAs) are a particular class of nucleic acid that can form secondary and tertiary structures capable of specifically binding proteins or other cellular targets, and are playing a similar role as antibodies [64]. Biosensors that use NAAs as BREs are classified as genosensors [56,65]. In this context, NAAs are short, single-stranded oligonucleotides with high affinity and specificity to a broad range of target molecules [47]. Although predominantly unstructured in

solution, they have incredible folding capabilities, even higher than their protein counterparts, enabling the association with their ligands to the point where the ligand becomes part of the aptamer architecture (Fig. 3.3) [47]. Because of their folding abilities and wide application, as well as the fact that they are easy to synthesize and store, NAAs represent a more effective approach compared to antibodies. Nevertheless, NAAs also present some disadvantages that should be considered when developing a new biosensor, these include aptamer degradation, cross reactivity, labelling costs and time required for sample preparation and/or modification [47,66]. Besides aptamers constituted by NAs alone, two particular types of such aptamers namely, aptazymes and peptide nucleic acids (PNAs), can also be considered [47,67]. Aptazymes are aptamers with catalytic properties much like enzymes, but with the ability of enduring repeated denaturation without losing their catalytic and binding capacities. [47] Also, because of their particular high affinity and high signal-to-noise ratio, they can be very useful to monitor very low concentrations of

FIGURE 3.3 Schematic illustration of a biosensor using an aptamer as a biorecognition element.

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key metabolites of diagnostic, environmental, and military value [47]. PNAs are synthetic DNA analogues with a polyamide backbone instead of a sugar-phosphate backbone [47,62,67,68]. As a main advantage over natural NAs, PNAs exhibit higher hybridization characteristics and better chemical and enzymatic stability due to their uncharged nature [67,68]. PNAs can also be associated with NAs in double- and triple-stranded complexes and perform better in nanoparticle aggregates [47]. 3.2.2.1.3 ENZYMES AND PROTEINS

Proteins are also employed as BREs if they interact selectively with a specific target compound. Enzymes are a very important subgroup as they display catalytic activity. Typically enzymes are very effective and diverse, being extensively used in biosensing due to the variety of reactions that can be measured and of products that result from the catalytic process, including protons, electrons, light, and heat. There are several mechanisms by which enzymes allow an analyte recognition [69]. For example, the enzyme may react directly with the analyte producing a product that is detectable, but the enzyme and analyte can also interact in a way that inhibits, activates, or even alters the enzyme properties [47]. Another property of interest is the wide variety of transducers that can be used to monitor the reaction [47]. Whether the purpose is an electrochemical, fluorescent, or colorimetric signal, enzymes offer an appealing range of choices with relatively simple assemblies [70]. Despite this attractiveness, allosteric proteins are multimeric in nature and present considerable instability, expression difficulties, and shorten biosensors’ lifetime [47]. Another type of protein used as BREs is lectins. Lectins constitute a wide family of proteins with high affinity for saccharide moieties on cell surfaces and protein aggregates via multivalent interactions arising from the spatial organization of oligosaccharide ligands,

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which is a property that has been extensively exploited as a basis for biosensor design [47]. Of all the lecithins, concanavalin A is one of the most widely used for saccharide detection by coupling this lecithin to fluorescent moieties for specific ligand detection [47]. 3.2.2.1.4 CELLS AND TISSUES

In cell- and tissue-based biosensors (usually bacteria, fungi, yeast, algae, or tissue-culture cells) the regulatory system of the cell is used to induce expression of a specific reported gene [48,56]. In other words, the cell is engineered to react to certain chemical signals producing, ideally, an easily quantifiable marker protein [56]. This approach can be done in either ex vivo or in vivo cells and is a great way to study hormones, drugs, or toxins in a continuous and noninvasive fashion by using biophotonics or other physical principles [48,56]. The use of cells as BREs has a number of advantages, including the batch amount since bacteria and yeast can be produced in large amounts in a very short period of time with a wide spectrum of enzymes at low cost. They are also easy to handle and maintain as there is no need for extraction and purification, as well as having considerable stability in pH and thermal variations [56]. However, biosensors developed with this type of BRE tend to be slow in response and less selective [48,56]. 3.2.2.1.5 MOLECULAR IMPRINTS

Molecular imprints represent a technique where a template of interest is used to create a cavity, or imprint, in a polymeric matrix (Fig. 3.4) [71
73]. The result is a polymer with affinity to the original molecule that served as the template [71,73]. Molecular imprinted polymers (MIP) offer the potential of very stable “solid-state
like” artificial biosensing elements [73]. In recent years, the technology of molecular imprinting has proliferated because it is an inexpensive, accessible, durable, and effective strategy for

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FIGURE 3.4

Schematic illustration of a molecular imprinting process: (A) target molecule; (B) the target molecule is inserted in a matrix to be polymerized; and (C) after polymerization the target molecule is removed, leaving a cavity in the matrix.

attaining high-specificity features [72,74]. The imprinting possibilities against a countless number of analytes makes this approach very appealing, allowing the construction of a robust artificial receptor that can effectively analyze multiple clinical samples without pretreatment or reagents [71,75]. It is generally accepted that it is the template that acts as the critical molecule and that the compounds responsible for the polymerization, cross-linkers, functional monomers, and solvents should be selected based on the physicochemical properties of the template [72,75]. The type of interactions that occur during the imprinting within the matrix-forming material (like covalent and noncovalent binding and metal-ion mediated imprinting) can, therefore, be customized according to the desired MIP and target analyte [71,73]. Equally important is the proportion of each compound in the mixture prior to polymerization [72,76]. Typically, this includes monomers that create the polymeric network and cross-linkers that ensure a high degree of reticulation in the final structure and the subsequent formation of a three-dimensional (3D) network, and initiators that are responsible to imitate the polymeric reaction. Overall, the monomer mediates specific chemical interactions with the target compound and the target is held in place by the cross-linking agent [72,76]. The selective artificial recognition cavities are then formed

by removal of the imprinted target, which is established in appropriate solvents [71,72]. Therefore different ratios between the monomer and cross-linker affect the type of dominant chemical interactions and inherent stability of the binding strength with the analyte of interest [72]. In this way, the correct compounds as well as their ratios can be optimized to improve selectivity and sensitivity [71]. Compared to natural compounds, MIPs have a higher shelf life and few storage requirements allowing highly selective and sensitive sensing systems with low cost [76]. For these reasons, this approach has attracted researchers’ interest and has had a profound impact on the development of biosensors for various applications [71,72,75]. 3.2.2.2 Transducer Technology 3.2.2.2.1 OPTICAL BIOSENSORS

Optical biosensors use a BRE integrated with an optical transducer system [77]. This allows a visible response when the analyte of interest is present in a sample and can also increase the signal intensity with an increasing concentration of the analyte [77]. Detection in this type of biosensor is based on the quantification of luminescence, fluorescence, or any color change, appearance, or disappearance by measuring absorbance, reflectance, phosphorescence, or

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

3.2 BIOSENSORS

fluorescence emissions occurring in the ultraviolet, visible, or near-infrared spectrum region [77,78]. Therefore the principle is to exploit the optical properties of the interaction of the BRE with the analyte of interest. An optical biosensor can use a wide range of biological materials, from enzymes, antibodies, antigens, receptors, and NA to whole cells and tissues as BREs [77
79]. Among the optical properties, fluorescence is by far the most exploited [78]. This is because most organic fluorophores are sensitive to environmental changes, which is the key to sensing applications [78]. Another attractive feature is the fact that they are easy to build and provide the detection of multiple compounds in a single device [78]. When the analyte is detected, the fluorescent signal is transduced and measured. These probes can detect biomarkers of organic or inorganic natures, even in complex samples, without compromising sensitivity [48,77]. Optical biosensors have the advantage of allowing a safe nonelectrical remote sensing of materials and usually do not require reference sensors since the comparative signal can be generated using the same source of light as the sampling sensor [79]. Also, when developed to give a response in the range of visible light, it discards the need for equipment to read results, making them particularly appealing for portability [78]. Several optical devices have been developed and successfully applied in cancer detection and metastasis as well as arthritis, inflammatory, cardiovascular, and neurodegenerative diseases, viral infections, and drug screening [48,78]. These can be effective devices in early detection in molecular and clinical diagnosis as well as for monitoring disease progression and therapy response [48]. 3.2.2.2.2 ELECTROCHEMICAL BIOSENSORS

Electrochemical biosensors usually require a reference electrode, counter electrode, and working electrode, the latter working as the

85

BRE [80]. These electrodes need to be stable with regard to conductivity and chemical properties in order to be reliable, which is why platinum, gold, graphite, and silicon compounds are the most-commonly biorecognition electrodes used, depending on the analyte of interest [80]. Electrochemical biosensor devices comprise a group of devices that can be of amperometric, potentiometric, conductometric, or voltammetric nature depending on the detection principle employed in the biosensor [78]. Amperometric biosensors operate at a given applied potential between the working and the reference electrodes [81]. Then a current signal is generated due to the oxidation or the reduction process that is as extensive as the concentration of the analyte [81]. These types of biosensors have similar response times, dynamic ranges, and sensitivities as the potentiometric biosensors [78]. A conductometric biosensor works under the principle of production or consumption of ionic species involved in the metabolic process [80]. Its attractiveness is due to its enhanced sensitivity, speed, and suitability for miniaturization because of the unnecessary reference electrode in the system [80]. Its limitation is in the charge carriers that affect the conductivity process, which directly affects the device’s poor selectivity [78,80]. A potentiometric biosensor works on the principle of potential difference between working and reference electrodes [81]. The measured species are not consumed like in the amperometric biosensor. Its response is proportional to the analyte concentration by comparison of its activity to the reference electrode [81]. The great advantage of potentiometric biosensors is their sensitivity and selectivity when a highly stable and accurate reference electrode is used [78]. Voltammetry is an interesting and versatile technique that can be used in biosensors. It combines electric current and potential

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difference enabling a reasonable system response with major applications as multicomponent detectors [78]. 3.2.2.2.3 PIEZOELECTRIC BIOSENSORS

Piezoelectric biosensors are based on acoustics and the measurement of changes in resonance frequency of piezoelectric crystals as a result of mass changes on the crystal structure [48,82]. The adsorption of the analyte causes characteristic vibrations in the charged crystals, altering their frequency [82]. This alteration can then be detected by electronic devices [82]. Several biosensors with piezoelectric transducers have been developed, for instance, for the detection of molecules like organophosphorus insecticide, formaldehyde, and cocaine, demonstrating the flexible applicability of this type of technology [83]. A major advantage is that synthetic quartz crystals are mass produced, allowing low costs and besides they can perform in several modes (i.e., direct, indirect, or label-free interactions) with the analyte [83]. The biggest limitation, especially for clinical or medical applications, is the fact that piezoelectric biosensors are not appropriate for detection of analytes in solution. In solution, the crystals’ vibration is compromised and, therefore, interferes with the principle of the technology [83].

commonly used for pesticides and pathogenic bacteria estimation [86]. Thermal biosensors do not need frequent recalibration. However, in general, they have limited applications because the pool of analytes that react exothermically is relatively small [84,85]. 3.2.2.2.5 MAGNETIC BIOSENSORS

Magnetic biosensors use magnetic nanoparticles and microparticles of 5
300 nm and 300
500 nm, respectively, in microfluidic channels using the magnetoresistance effect [48,87]. These particle surfaces are modified and functionalized to recognize specific molecules with appealing sensitivity [87]. Magnetic biosensors have attracted researchers’ attention because they offer great advantages compared to fluorescent-based methods. Magnetic probes are more stable over time in culture and can be used for long-term labeling assays without leading to background noise effects [87]. Magnetic fields on external surfaces provide remote measurement and regulation of the biological environment [87]. Moreover, their potential high sensitivity allows detection at significantly lower-protein concentrations compared with fluorescent-based techniques [87].

3.2.2.2.4 THERMOMETRIC BIOSENSORS

3.2.3 Biosensors versus Conventional Techniques in Health Care

Thermometric or calorimetric biosensors use a physical transducer that can detect heat differences when the analyte of interest is recognized [48]. The temperature changes between the substrate and the product can be measured and even a small change in the temperature can be detected by thermistors [84]. Thermometric biosensors are suitable for enzyme-based reactions and combine enzymes with temperature sensors. [84,85]. When the analyte reacts with the enzyme, the heat of the reaction is measured and calibrated against the analyte concentration. These biosensors are

In breast cancer, the concept of traditional approaches for diagnostics is associated with common techniques like mammographies, MRI, X-rays, and others. This approach is, however, time-consuming and expensive for the patients and healthcare systems. Mammographies, for example, have a sensitivity of 67.8% and specificity of 75.0%, but are not suitable for subjects with dense breast tissue [88]. Moreover, it seems to only reduce breast cancer death associated with a late or no diagnosis by 0.0004%, which may not be as useful as previously thought [88]. Ultrasound is performed as a

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3.2 BIOSENSORS

supplement to mammography, improving sensitivity of imaging to 83% [88]. However, specificity is reduced to 34% because ultrasonography fails to detect many tumors as a consequence of the similar acoustic properties of healthy and cancerous tissues [88]. In the case of MRI, because of the high false-positive rates, along with logistic issues such as costs, time consumption, and need of experienced radiologists, this approach is usually recommended only in high breast cancer
risk cases [88]. Moreover, despite being the technique with higher sensitivity (i.e., 94.4%) than mammography and ultrasound, it has the lowest specificity (26.4%). Hence, the high number of false-positive results [88]. Along with physical examination, these constitute the main conventional techniques for breast cancer diagnosis and it is clear that the process from a screening routine examination to a complete diagnosis is cumbersome. Taking into account how critical early detection is for a positive prognosis, these techniques, although reliable, are not the best possible solution. As technology progresses, the type of equipment available in medicine has become more sophisticated. The goal is to work toward a more dynamic healthcare system, where patients not only spend less time waiting for the results of tests and examinations, but also guarantees that those exams get less and less invasive and uncomfortable. This will ultimately lead to a decrease in healthcare costs and to a more dynamic healthcare system. New innovations in biosensor development address these concerns in healthcare which are particularly relevant in under-developed countries. In fact, according to several studies cancer mortality seems to be higher in countries with lower-income populations [2,4,89]. For instance, a study showed that women in Uganda and Kenya only seek treatment when they are already at advanced stages of the disease [89]. This means that any treatment will have a very low probability of success, besides

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the lack of proper facilities and diagnosis tools [89]. By contrast, in countries like England and Australia the percentage of women diagnosed in late stages of the disease is very low [89]. Overall, biosensors potentially offer high precision, accuracy, specificity, fast response, portability, and cost-effective properties to monitor virtually any biomarker of any disease. These devices are already in use in a number of healthcare systems. The most wellknown biosensors are glucose [48,52] or urea [90] monitoring devices, where, instead of having to go to a hospital and take blood samples, people can go to a health center or pharmacy and easily get their sugar levels checked within few minutes with a single drop of blood. Several other devices include quantitative measurements of cardiac markers in undiluted serum, immunosensor array for clinical immunophenotyping of acute leukemias, and the effect of oxazaborolidines on immobilized fructosyltransferase in dental diseases. In breast cancer, regardless of the transducer approach or biomarker choice, biosensor technology is a growing market. With an estimation of the global point-of-care diagnostics market being roughly US$40.5 billion by 2022, it is not surprising that there are so many studies invested in these devices [91]. From electrochemical transducers applied to the detection of proteins like osteopontin with an aptamer as a BRE [92] to a colorimetric technology to detect HER2 through an approach with HER2 antibodies anchored gold nanocluster
loaded liposomes [93], the list of studies regarding biosensing applied to the breast cancer detection field is wide, and is discussed in Section 3.3. Whether straightforward with simple assembly or rather sophisticated by means of the use of quantum dots (QDs) labeled with primary antibodies against MCF-7 cell surface proteins [88], the goal remains the same—to construct a scalable, portable device that can provide reliable, safe, and user-friendly diagnosis in early breast cancer detection.

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3.3 BIOSENSORS FOR RAPID BREAST CANCER DETECTION An important aspect regarding biosensors lies in the vast possible combinations between BREs’ nature, support materials, and signal transducer technology. This alluring aspect is what allows to experiment with similar principals and still create so many different and innovative devices. Elegant and/or unconventional principles are being used to develop devices that try to answer the need of affordable, portable, and time-effective tests without compromising sensitivity and specificity. Although most of them are not yet at a development stage to be considered suitable for clinical and/or market applications, some approaches show promising and cost-effective outcomes regarding sensitivity and specificity towards known breast cancer biomarkers. These studies could have a positive impact

regarding early breast cancer detection in a near future and are discussed briefly next as well as summarized in Tables 3.2
3.6.

3.3.1 BRCA1 Regarding detection of BRCA1 genes (Table 3.2), there are several approaches with incredibly low limits of detection. Tiwari et al. developed an electrochemical biosensor using chitosan-co-polyaniline, a moderately inexpensive and stable electroactive material, as a sustainable support matrix applied on a support of indium
tin-oxide (ITO) [94]. A probe with cDNA sequences associated with BRCA1 was immobilized onto the surface giving an electrochemical response in the presence of the ssDNA [94]. The detection limit of this biosensor was 0.05 fmol, also showing excellent sensitivity and reproducibility, with promising application for the efficient and

TABLE 3.2 Studies on Biosensor Construction for the Detection of Breast Cancer Biomarker BRCA1 (Breast Cancer 1). LOD; Linear Range

Response Time

ShelfLife

Immobilization of complementary DNA probe

2.104 μA/fmol; 0.05
25 fmol

16 s

6 months [94]

Electrochemical (EIS)

BRCA1 complementary sequence immobilized on AuNPs, within a highly cross-linked aminemodified PEG film

1.72 fM; 50.0 fM 2 1.0 nM

ND

ND

[95]

Electrochemical

Attachment of DNA probes to PANI/PEG nanofibers

0.0038 pM; 0.01 pM 2 1 nM

30 min

.10 days

[96]

Electrochemical

Zwitterionic peptides anchored to a conducting polymer of citrate doped PEDOT

0.03 fM;


ND

ND

[97]

Fluorescence

Fluorescent dual-channel based on carbon dots and AuNPs for detecting nucleotide BRCA1 sequences


; 4 2 120 nM

ND

ND

[98]

Colorimetric

Three spots labeled with digoxin per detected 10 fM; target, amplifying the typical enzymatic reading 10 fM 2 10 nM

ND

ND

[99]

Transducer

Principle

Electrochemical (CV and EIS)

References

ND, Not disclosed; CV, cyclic voltammetry; EIS, electrochemical impedance spectroscopy; AuNPs, gold nanoparticles; PEG, polyethylene glycol; PEDOT, poly(3,4-ethylenedioxythiophene); PANI, polyaniline; PEG, polyethylene glycol.

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TABLE 3.3 Studies on Biosensor Construction for the Detection of Breast Cancer Biomarker Estrogen-Alpha (ERα), Progesterone (PR), or Carcinoembryonic Antigen (CEA). LOD; Linear Range

Response Time

Shelf-Life References

Hollow-core of photonic crystal fiber with anti-ER labeled primary and secondary antibodies

0.4 μg/mL

ND

ND

[100]

Electrochemical (EIS)

Aptamer on Au electrode and iron redox probe readings

0.90 ng/mL; 10 2 60 ng/mL

40 min

ND

[101]

CEA

Colorimetric

AuNPs/few-layer black phosphorus hybrid

0.20 pg/mL; 1
104 pg/mL

ND

ND

[102]

CEA

Chemiluminescence CEA aptamer linked to hemin aptamer, with 1,10 oxalyldiimidazole

0.58 ng/mL; 0
200 ng/mL

30 min

ND

[103]

CEA

Colorimetric

AuNPs as carriers of anti
CEA antibody, labeled with biotin

48 pg/mL; 15 min 0.05 2 50 ng/mL

ND

[104]

CEA

Electrochemical

Paper-based microfluidic immunodevice

0.01 ng/mL;


ND

ND

[105]

CEA

Fluorescence

Fluorescence resonance energy transfer between up-converting nanoparticles and PdNPs

1.7 pg/mL; 4 2 100 pg/mL

ND

ND

[106]

Biomarker Transducer

Principle

ERα

Optical

PR

ND, Not disclosed; EIS, electrochemical impedance spectroscopy; AuNPs, gold nanoparticles; PdNPs, palladium nanoparticles.

precise detection of breast carcinoma at its early stage [94]. Wang et al. also developed an electrochemical biosensor to detect BRCA1 from serum samples in levels as low as 1.72 fM [95]. This label-free DNA sensor was constructed by the modification of a glassy carbon electrode (GCE) with highly crosslinked polyethylene glycol film containing amine groups. This sensor was then modified with gold nanoparticles (AuNPs), yielding an outstanding sensitivity and effective readings [95]. Hui et al. developed an ultrasensitive electrochemical biosensor to detect BRCA1 based on polyaniline/polyethylene glycol nanofibers [96]. The biosensor allows the detection of BRCA1 in human serum without being affected by nonspecific adsorption in

complex biological media, and the nanofibers show antifouling properties through high immobilization ability to capture the probes [96]. Wang et al. [97] used zwitterionic peptides modified with a polymer of citrate doped poly(3,4-ethylenedioxythiophene) (PEDOT) creating an electrochemical biosensor with excellent antifouling ability and good conductivity for subsequent binding of a suitable DNA probe as the BRE [97]. This biosensor displayed a lower detection limit than Wang et al. [95], 0.03 fM versus 1.72 fM. Zhong et al. developed a fluorescent dualchannel biosensor model based on carbon dots and AuNPs assisted by a hairpin structure [98]. BRCA1 RNA/DNA targets bound specifically to its complementary sequence on

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TABLE 3.4 Studies on Biosensor Construction for the Detection of Breast Cancer Biomarker Human Epidermal Growth Factor Receptor 2 (HER2). Transducer

Principle

LOD; Linear Range

Response Time

ShelfLife References

Electrochemical (amperometric)

Sandwich-type immunoassay based on nanobodies


; 1 2 200 μg/mL

2 or 20 min

.3 [107] weeks

Electrochemical

DNA generated electric current with DNA 0.047 pg/mL; self-assembly for signal amplification 1 2 100 pg/mL

ND

ND

[108]

Colorimetric

HER2 antibodies anchored AuNPs-loaded liposomes

5 Sk-Br-3 cells; —

2h

ND

[93]

Electrochemical (amperometric)

Modified AuNPs and graphene oxide loaded on GCEs

0.16 nM; 0.37 2 10 nM

ND

ND

[109]

Electrochemical

Sandwich-type aptasensor that uses molybdate to generate an electrochemical current


; 0.01 2 5 ng/mL

ND

ND

[110]

Electrochemical

Organic-electrochemical-transistor-based biosensor


; 10214 2 1027 g/mL

ND

ND

[111]

Electrochemical (voltammetry)

Reduced graphene oxide-chitosan film as electrode material using MB redox probe

0.21 ng/mL; 0.5 2 2 ng/mL

ND

ND

[112]

Electrochemical

Immunosensor with hydrazine and aptamer-conjugated AuNPs

37 pg/L; 1 ng/L 2 10.0 μg/L

ND

ND

[113]

Electrochemical (voltammetry)

Anti
HER2 antibodies conjugated with iron oxide nanoparticles on Au electrode

0.995 pg/L; 10 ng/L 2 10 μg/L

ND

ND

[114]

Photoelectrochemical Zinc oxide/graphene composite and S6 aptamer on a portable indium
tin-oxide microdevice

58 cells/mL; 102
106 cells/mL

20 min

[115] .2 weeks

Electrochemical (EIS) The charge-transfer resistance of an iron redox probe changes with the amount of protein bound to the antibody

7.4 ng/mL; 10 2 110 ng/mL

35 min

ND

[116]

Electrochemical

Immobilized polycytosine DNA sequence in an AuNP matrix

0.5 pg/mL; 1 2 1000 pg/mL

ND

ND

[117]

Electrochemical

Interdigitated Au electrodes modified with aptamer

1 pM; 1 pM 2 100 nM

ND

ND

[118]

Electrochemical

Inkjet-printed 8-electrodes array, requiring 12 pg/mL; biotinylated antibody, and polymerized
horseradish peroxide labels

15 min

ND

[119]

ND, Not disclosed; EIS, electrochemical impedance spectroscopy; AuNPs, gold nanoparticles; GCE, glassy carbon electrodes; MB, methylene blue.

the AuNPs and released carbon dots, inducing a positive fluorescent signal with a linear range of 4
120 nM [98]. Yang et al. designed a “sandwich-like” biosensor on a magnetic

bead platform. In this, a tetrahedron-shape reporter probe was designed having three vertices labeled with digoxin and the fourth vertice labeled with a detection probe. The

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TABLE 3.5 Studies on Biosensor Construction for the Detection of Breast Cancer Biomarker Mucin 1 or Cancer Antigen 15-3 (CA 15-3). Biomarker Transducer

Principle

Response LOD; Linear Range Time

ShelfLife References

Mucin 1

Electrochemical

Aptamer/cell/aptamer sandwich architecture on an electrode surface

100 cells/mL; 102 2 107 cells/mL

ND

ND

[120]

Mucin 1

Electromagnetic Aptamer-functionalized Au nanorods

100 cells; 102 2 105 cells/mL

30 min

ND

[121]

Mucin 1

Electrochemical (voltammetry)

Polyadenine-aptamer functionalized AuNPs/graphene oxide hybrid

8 cells/mL; 10 2 105 cells/mL

40 min

ND

[122]

Mucin 1

Electrochemical

Biotinylated aptamer immobilized on a composite of AuNPs-graphene oxide-PEDOT

0.031 fM; 3.13
31.25 nM

15 min

14 days

[123]

CA 15-3

Optical

Antibodies immobilized by surface standard amine coupling on an optofluidic ring resonator

1 unit/mL;


20 min

ND

[124]

CA 15-3

Optical

Cadmium sulfide QDs modified by cysteamine capping

0.002 kU/L;


15 min

ND

[74]

CA 15-3

Electrochemical (voltammetry)

Detection of 7 tumor marker using alkaline phosphatase-based competitive immunoassay for hydroquinone readings

0.7 U/mL 1.2
3.7 U/mL

ND

ND

[125]

CA 15-3

Electrochemical

Nanoporous Au/graphene hybrid platform combined with horseradish peroxidase

5 3 1026 U/mL; 2 3 1025
40 U/mL

ND

ND

[126]

CA 15-3

Electrochemical

Label-free highly conductive graphene N-doped graphene sheets modified electrode

0.012 U/mL; 0.1
20 U/mL

ND

ND

[127]

CA 15-3

Electrochemical

Functionalized graphene with 1-pyrenecarboxylic acid sensor probe and MWCNTs with ferritin labels

0.009 U/mL; 0.05
100 U/mL

ND

ND

[128]

CA 15-3

Electrochemical

Electrically-conducting poly (toluidine blue) employed as synthetic receptor film

0.10 U/mL; 0.10
100 U/mL

ND

ND

[129]

ND, Not disclosed; PEDOT, poly(3,4-ethylenedioxythiophene); AuNPs, gold nanoparticles; MWCNTs, multiwall-carbon nanotubes; QDs, quantum dots.

antidigoxin antibody with labeled with suitable enzymes had three places to bind for each detection probe, thereby contributing to signal amplification [99]. This approach

also allowed distinguishing DNA sequences with only 1 base mismatch, and the performance proved to be comparable to PCR products [99].

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TABLE 3.6 Studies on Biosensor Construction for the Detection of miRNA 21 or miRNA 155. Biomarker Transducer

Principle

LOD;Linear Range

Response Time

ShelfLife References

miRNA 21 Electrochemical

Probes attached to a pencil graphite electrode

1.0 mg/mL;


ND

ND

[130]

miRNA 21 Electrochemical

Two auxiliary probes that self-assemble to form 1D DNA concatemers

100 aM; 100
105 aM

ND

ND

[131]

miRNA 21 Electrochemical

MB as a redox indicator

84.3 fM; 0.1 2 500.0 pM

60 min

ND

[132]

miRNA 21 Electrochemical (amperometric)

Hybridization to a specific biotinylated DNA probe immobilized on magnetic beads

0.04 pM; 1.0
100.0 pM

30 min

ND

[133]

miRNA 21 Fluorescence

2-Aminopurine probe in conjunction with a G-quadruplex structure

1.48 pM;


ND

ND

[134]

miRNA 21 Electrochemical

Probe modified with a pyrrolidinyl peptide nucleic acid/PPy/silver nanofoam

0.20 fM; 0.20
106 fM

ND

ND

[135]

miRNA 21 Electrochemical

Target-induced glucose release from propylamine-functionalized mesoporous silica nanoparticle

19 pM; 50
5 3 103 pM

ND

ND

[136]

miRNA 155

Electrochemical

Graphene oxide sheet on the surface of the glassy carbon electrode with thiolated probe-functionalized Au nanorods

0.6 fM; ND 2.0 2 8 3 103 fM

ND

[137]

miRNA 155

Electrochemical

Immobilization of the anti-miRNA-155 on Au-SPE

5.7 aM; 10
109 aM

ND

ND

[38]

miRNA 155

Colorimetric

DNA probe covalently bound to negatively charged AuNPs

100 aM; 100
105 aM

ND

ND

[138]

ND, Not disclosed; MB, methylene blue; PPy, polypy.

3.3.2 ERα The development of biosensors for ERα biomarker detection are summarized in Table 3.3. Padmanabhan et al. developed an approach with an immunobiosensor to detect ERα using an optical transducer that can detect the protein in volume samples as low as 50 nL [100]. The overall biosensing system, using a hollow core
photonic crystal fiber in a total internal reflection configuration, allowed a fluorescence

green and red response from the recognition of ERα by the secondary antibody [100].

3.3.3 PR Studies on the development of PR biosensors are described in Table 3.3. Jime´nez et al. selected a progesterone aptamer by systematic evolution of ligands by exponential enrichment (SELE) to develop a label-free aptasensor with

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3.3 BIOSENSORS FOR RAPID BREAST CANCER DETECTION

enhanced signal gain monitored by electrochemical impedance spectroscopy [101]. The conformational change of the aptamer immobilized on the gold electrode upon binding to progesterone resulted in increased electron transfer resistance upon an iron standardredox probe, and allowed a linear range detection of progesterone from 10 to 60 ng/mL, with a detection limit of 0.90 ng/mL [101].

3.3.4 CEA The latest developments on CEA biosensors are also provided in Table 3.3. Peng et al. took advantage of the catalytic features of few-layer black phosphorus modified by onsite production of AuNPs against 4-nitrophenol that was detected by colorimetric assays [102]. This catalytic activity was reversibly reduced in the presence of the antibody, but reactivated when CEA was added. The detection limit and linear detection range proved adequate for sample analysis, with values of 0.20 pg/mL and 1 pg/mL
10 μg/mL, respectively [102]. Khang et al. developed an all-in-one chemiluminescence aptasensor [103]. A dual DNA aptamer for competitive binding of CEA and hemin along 30 minutes at room temperature was used. Amplex Red and H2O2 were then added into the system to form resorufin, which depended on the concentration of horseradish peroxidase (HRP)-mimicking G-quardruplex DNAzyme formed upon the binding interaction between hemin and the dual DNA [103]. Bright red light was observed after the addition of 1,10 oxalyldiimidazole to detection cell, and decreased with the increasing CEA concentrations [103]. Liu et al. developed a colorimetric enzyme immunoassay with AuNPs as carriers of HRP-labeled anti
CEA detection antibody, and magnetic microparticles were used as supporting substrates [104]. The complex generated an optical signal, exhibiting improved sensitivity compared with a CEA ELISA

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kit [104]. Wu et al. used a sandwich immunoassay in which the secondary antibody allowed the growth of the long chain polymeric material providing numerous sites for subsequent HRP binding [105]. This approach turned out to be a way to amplify the signal, as the more secondary antibodies bound to the support (more CEA was on the platform), the higher the electrochemical signal generated by HRP-O-phenylenediamine-H2O2 system. The support was a carbon electrode printed on a paper-based microfluidic electrochemical immunodevice [105]. Li et al. constructed a biosensor based on fluorescence resonance energy transfer (FRET) between up-converting nanoparticles (UCPs) and palladium nanoparticles (PdNPs) [106]. Having the aptamer bound to the UCPs, the close proximity of the PdNPs to the aptamer resulted in the fluorescence quenching of the UCPs. When CEA was present, the aptamer preferentially combined with CEA yielding conformational changes that weakened its interaction to the PdNPs, thereby recovering fluorescence signals. This system allowed an ultrasensitive detection of CEA performed in diluted human serum with a linear range of 4
100 pg/mL and a detection limit of 1.7 pg/mL [106].

3.3.5 HER2 The HER2 breast cancer biomarker is among the most targeted compounds for biosensor applications (Table 3.4). Patris et al. developed a sandwich-type immunoassay based on nanobodies developed to detect another epitope of HER2 on screen-printed electrodes (SPEs) [107]. The capture nanobody was immobilized on the carbon-working electrode and the detection antibody was labeled with HRP. The signal corresponded to the electroreduction of p-quinone, generated at the SPE, by the HRP in the presence of hydroquinone and hydrogen peroxide [107]. Shen et al. developed a DNA self-assembly amplification biosensor capable

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of generating electric current to power electrochemical biosensing. An HER2 aptamer was used both as a ligand for recognition and as a signal-generating reporter on a sandwich format [108]. The detection limit was 0.047 pg/mL with a detection range of 1
100 pg/mL [108]. Tao et al. developed a colorimetric biosensor that uses a probe with HER2 antibodies anchored on liposomes that were loaded with BSA or gold nanoclusters. These gold nanoclusters have an intrinsic peroxidase property and react with 3,30 ,5,50 -tetramethylbenzidine (TMB) in the presence of hydrogen peroxide, thereby changing the color of the solution [93]. This platform allowed the detection of HER2-positive breast cancer cells in human serum samples and breast cancer tissue with a detection limit as low as five cells [93]. Saeed et al. used AuNP with a short complementary sequence for HER2 and covalently bond to a GCE modified by graphene oxide [109]. After binding HER2 to the AuNPs, an additional DNA short sequence modified with HRP was able to hybridized with the free sequence of HER2 producing an electrochemical signal in the presence of TMB and hydrogen peroxide [109]. Hu et al. used an HER2specific aptamer as a ligand to capture HER2 and another to generate a redox current signal. This current was obtained by the reaction between the phosphate moieties in the aptamer and molybdate [110]. Overall, the electrochemical current generated by the aptasensor was proportional to the HER2 concentration in the range of 0.01
5 ng/mL [110]. Fu et al. developed an organic-electrochemicaltransistor-based biosensor that detects electrochemical activity on gate electrodes to the concentration of 10214 g/mL [111]. The gold gate electrode was modified with a capture specific polyclonal anti
HER2 antibody and the detection was enabled by a secondary antibody bound to HRP. The current was produced in the presence of HER2 and hydrogen peroxide [111]. Tabasi et al. developed an

ultrasensitive electrochemical aptasensor that uses a graphene and chitosan film as a suitable electrode material for aptamer binding [112]. After HER2 interaction with the aptamer, the conformational changes dictated that the electrochemical probe MB would produce a higher signal, which was concentrationdependent [112]. Zhu et al. developed a sensing system based on a similar-to-sandwich approach. In this, the probe was prepared by immobilizing the antibody on a nanocomposite of AuNPs capped with 2,5-bis(2-thienyl)-1Hpyrrole-1-(p-benzoic acid) directly on the bare electrode surface. The detection was achieved by a hydrazine
AuNP
aptamer bioconjugate, having the hydrazine reductant bound to the AuNPs and containing silver that should be reduced for signal amplification [113]. Another interesting aspect of this approach was that the silver-stained target cells exhibit a black color which is easily observed through a microscope, providing a simple and convenient approach for clinical analysis of cancer cells [113]. A label-free immunosensor was designed by Emami et al. for HER2 detection in real samples [114]. Anti
HER2 antibodies were attached to iron oxide nanoparticles forming stable bioconjugates laid over the gold electrode surface [114]. The immunosensor was responsive to HER2 concentrations as low as 0.995 pg/mL and a sensitivity of 5.921 μA 3 mL/ng [114], against an iron redox probe. Liu et al. developed a photoelectrochemical biosensor for the detection of SK-Br3, which is a HER2-positive cell line, using an oxide zinc and graphene composite and a S6 aptamer [115]. The high photoelectric signal of the zinc oxide, the graphene’s superior charge transportation and separation, and the S6 aptamer’s specificity to target Sk-Br-3 cells seemed to be an improvement in sensitivity and selectivity, making this approach a promising candidate for accurate detection of cancer cells [115]. Arkan et al. developed an electrochemical immunosensor for the analysis of HER2 by

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preparing carbon paste electrodes based on graphite powder, multiwall-carbon nanotubes (MWCNTs), an ionic liquid, and paraffin and further decorated with AuNPs by electrodeposition [116]. The charge-transfer resistance increased linearly with increasing concentrations of HER2 antigens for an optimum incubation time of 35 minutes, with linear dependency between 10 and 110 ng/mL [116]. Li et al. also used the electrochemical immunosensor approach, but with an immobilized polycytosine DNA sequence in an AuNP matrix [117]. The HER2 captured by the immunosensor was detected due to a reaction between polycytosine DNA phosphate backbone an molybdate, in a similar approach to that in [112] generating an electrochemical current at the surface of the electrode [117]. The biosensor showed linear behavior from 1 pg/mL to 1 ng/mL, with a limit detection of 0.5 pg/mL and no cross reactivity with human IgG, human IgA, p53, CEA, or protein kinase [117]. Arya et al. combined interdigitated microelectrodes modified with a thiol terminated
DNA aptamer for HER2 to develop a simple and sensitive biosensor for HER2 [118]. The use of interdigitated gold electrodes is the biggest difference compared with previous approaches [118]. The biosensor proved excellent selectivity when challenged with other serum proteins and exhibited a dynamic linear range from 1 pM to 100 nM [118]. Carvajal et al. found an inexpensive approach (under US$0.25) by developing a fully inkjet-printed
electrochemical sensor [119]. The device platform featured an inkjetprinted gold working 8-electrode array, a counter electrode, and an inkjet-printed silver electrode that was chlorinated with bleach to produce a Ag/AgCl quasireference electrode [119]. A full sandwich immunoassay was constructed into the microfluidic device in which the labeling was achieved through a streptavidin/HRP composite. The assay time

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was 15 minutes and the limit of detection was 12 pg/mL [119].

3.3.6 Mucin 1 Table 3.5 summarizes the biosensors developed to detect mucin 1 protein (MUC1). Zhu et al. used an aptamer-cell
aptamer sandwich architecture approach to detect MUC1 in Michigan cancer foundation-7 (MCF-7) human breast cancer cells [120]. The biosensor presents a sandwich architecture that can only be formed in the presence of the targeted cells. The electrochemical response comes from the enzyme HRP-labeled on the MUC1 aptamer and the subsequent reading of the electron mediator thionine [120]. The specificity is further increased with the aptamer doubled recognition ability [120]. Li et al. selected an electromagnetic approach with surface plasmon resonance as the detecting method of MUC1 on MCF-7 cells [121]. MUC1 aptamer functionalized
gold nanorods allowed an excellent dynamic range from 100 to 105 cells/ mL with a detection limit of 100 cells/mL in only 30 minutes [121]. Wang et al. developed a sandwich electrochemical biosensor based on a polyadenine-aptamer
modified gold electrode and a polyadenine-aptamer functionalized AuNPs/graphene oxide hybrid for the labelfree and selective detection of MUC1 in breast cancer cells MCF-7 [122]. Under optimized experimental conditions the biosensor detected down to 8 cells/mL, along with a linear range of 10
105 cells/mL [122]. Gupta et al. developed an electrochemical aptasensor based on the conducting properties of a polymer nanocomposite [123]. The nanocomposite film of AuNPs and graphene oxide
doped PEDOT was deposited onto a surface of fluorine tin oxide glass [123]. This approach allowed the detection of MUC1 in concentrations as low as 0.31 fM with a reusability of the aptaelectrodes of 8 times [123].

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3.3.7 CA 15-3 CA 15-3 biosensors are also described in Table 3.5. Using an optical approach, Zhu et al. developed a label-free optofluidic ring resonator sensor that could rapidly detect CA 15-3 [124]. The sensor was capable of detecting about 1 U/mL CA 15-3 in diluted human serum samples within approximately 30 minutes [124]. Elakkiya et al. also applied optics as the transducer technology to develop a biosensor for CA 15-3 using a cadmium sulfide QD surface that was cysteamine capped [74]. The device was tested in saline and serum samples spiked with antigens and was able to detect a very low concentration of 0.002 KU/L with a constant response time of 15 minutes [74]. Marques et al. developed the first multiplexed electrochemical immunosensor for the simultaneous detection of CA 15-3 and HER2 [125]. The immunosensor was constructed on a personalized dual screen-printed carbon electrode with surfaces modified with in situ electrodeposited gold nanoparticles [125]. These electrodes were then individually coated with a monoclonal antihuman
CA 15-3 or a monoclonal antihuman
HER2 antibody [125]. The antigen
antibody interactions were detected by voltammetric analysis with a limit of detection of 5.0 U/mL [125]. Ge et al., Li et al., and Akter et al. all used graphene to develop electrochemical immunosensors. Ge et al. used a nanoporous/graphene hybrid as a platform, using liposomes with enzyme HRP encapsulated as labels [126]. The presence of CA 15-3 released the enzyme from the liposome, thereby reducing hydrogen peroxide with thionine as an electron mediator. The encapsulation proved to be a good amplification strategy, allowing a limit detection as low as 5 μU/mL [126]. Li et al. used graphene applied to an electrochemical immunosensor, but in a N-doped graphene sheets manner [127]. This approach incorporated high conductivity to the graphene-modified electrode, exhibiting

significant electron transfer and high sensitivity without the need for labeling [128]. The immunosensor exhibited a detection limit down to 0.012 U/mL with a linear performance in the range of 0.1
20 U/mL [127]. In another amplification approach, Akter et al. used noncovalent functionalized graphene oxides as sensor probes and multiwalled carbon nanotube
supported numerous ferritin as labels, and both bound to a suitable CA 15-3 antibodies [128]. The amide bond between amine groups of secondary antibody and ferritin and carboxylic acid groups of MWCNTs allowed the detection of CA 15-3 through an enhanced bioelectrocatalytic reduction of hydrogen peroxide mediated by hydroquinone at the functionalized graphene probe [128]. Ribeiro et al. developed an electrochemical biosensor with a synthetic receptor film using molecular imprinting (MIP) strategies [129]. In this approach, CA 15-3 was imprinted on a poly(toluidine blue) film and assays were performed in buffer and artificial sera, showing selective adsorption of CA 15-3 onto MIP film after 30 minutes incubation [129]. Calibration plots showed a linear dependency of target protein concentration from 0.10 to 100 U/mL, with a 0.10 U/mL detection limit [129].

3.3.8 miRNA 21 Table 3.6 provides examples of miRNA 21 and miRNA 155 biosensors. Kilic et al. designed an electrochemical biosensor based on enzyme amplified biosensing of mir21 from cell lysate of total RNA [130]. The detection of mir21 was achieved by capture probes and/or cell lysates covalently attached onto the pencil graphite electrode by coupling agents of N-(dimethylamino)propyl-N0 -ethylcarbodiimide hydrochloride and N-hydroxysulfosuccinimide [130]. The proposed enzymatic detection method was compared with the conventional guanine oxidation based assay in terms of

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detection limit and specificity [131]. The biosensor approach proved to be sensitive with a limit detection of 1 μg/mL [131]. Hong et al. used electrochemical technology to develop an ultrasensitive biosensor for the detection of cancer-associated circulating miRNA-21 [132]. Using a self-assembled DNA concatamer, a long DNA chain of repeated copies of the same DNA sequences linked end-to-end allowed the detection of miRNA-21 in complex biological samples enzymes or labels with a detection limit as low as 100 aM [132]. Vargas et al. developed a sensitive amperometric magnetobiosensor for quick detection of microRNAs [133]. The strategy involved direct hybridization of the target with a specific biotinylated DNA probe immobilized on magnetic beads modified with streptavidin. The label was provided by a specific DNA
RNA antibody and the bacterial protein A conjugated with a HRP homopolymer for signal amplification [133]. This single-step device achieved a linear concentration range between 1.0 and 100.0 pM and a limit detection of 10 attomoles in a 25 μL sample, without any target miRNA amplification, in just 30 minutes [133]. Another approach for miRNA-21 detection with an electrochemical transducer was conducted by Raffiee-Pour et al. In this study, methylene blue was used as a redox indicator, therefore discarding the use of labels. Kinetic assays showed that methylene blue had stronger and more stability with miRNA/DNA than with ss-DNA, achieving detection limit of 84.3 fM. Li et al. combined a 2-aminopurine probe with a G-quadruplex structure to develop a simple sensor that can detect overexpressed miRNA-21 from human breast cancer cell lysate without quenchers and enzymes [134]. The biosensor contained two DNA hairpins that significantly enhanced the probe’s fluorescence providing a limit detection of 1.48 pM [134]. Kangkamano et al. used a modified electrode to develop a label-free electrochemical biosensor for miRNA-21 detection [135]. The

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probe was modified with a pyrrolidinyl peptide nucleic acid/polypyrrole/silver nanofoam and the electrochemical signal was proportional to miRNA-21 concentrations in the detection range 0.20
106 fM, with a detection limit of 0.20 fM [135]. Deng et al. developed an electrochemical biosensor based on targetinduced glucose release from propylaminefunctionalized mesoporous silica nanoparticles [136]. Glucose was employed as the signalgeneration tag for glucometer readout and the overall strategy allowed avoiding labeling and the time-consuming, repeated washing steps and had a limit detection of 19 pM [136].

3.3.9 miRNA 155 Azimzadeh et al. developed an electrochemical nanobiosensor applied to miRNA 155 detection on plasma samples and attributed the great selectivity and sensitivity to the combination of a graphene oxide sheet on the surface of the GCE with thiolated probefunctionalized gold nanorods [137]. The electrochemical signal showed a linear detection range from 2.0 fM to 8.0 pM and a detection limit of 0.6 fM [137]. Cardoso et al. developed a simple electrochemical biosensor that can monitor attomolar levels of miRNA 155 breast cancer [38]. This biosensor can detect concentrations of miRNA 155 as low as 1
10 aM in a serum background, thus allowing a high degree of sample dilution. In addition, it surpasses interferences and therefore, it can be reused along consecutive readings with new solutions while exhibiting high selectivity towards other proteins in biological fluids and cell extracts from other cancers [38]. Hakimian et al. took an optical approach for the detection of miRNA 155, being able to specify 3-basepair mismatches and genomic DNA from target miRNA 155 [138]. The strategy included using a DNA probe that can covalently bind to the negatively charged AuNPs, allowing

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electrostatic adsorption of the target miR-155 onto the positively charged AuNPs surface [138]. After hybridization, an optical signal with a detection limit of 100 aM and a wide linear range from 100 aM to 100 fM could be detected [138].

3.4 CONCLUSION The future of the biosensors field regarding their application in breast cancer detection is, primarily, to consider the panel of stablished biomarkers and develop a device that allows multianalyses of those markers at the same time and with a single biological sample. It is highly unlikely that such devices would replace completely conventional methods. The established biomarkers, although with good predictive performances, are not specific to a particular breast cancer type nor breast cancer in general. Therefore it is the abnormal levels of the biomarkers, as a panel, that can provide important information as complementary examinations providing, in a single analysis, valuable information for prognosis and therapy approaches. Along with the advancements in the development of multiplex biosensors that can provide a more accurate analysis of cancer biomarkers, there are continued improvements in biosensor technology toward miniaturization to make the devices easier to carry. With this comes the idea of wearable devices, which a direction that can increase the ability to test for a panel of analytes at or near the patient [139]. The generally increased cost per test can then be reconciled with the potential to decrease the overall cost of care by the improved turnaround time [139]. People who need long-term care and people living far from health and medical services would benefit from wearable biosensors, by remote monitoring, allowing patients to spend less time in

hospitals, improving their comfort, and reducing the burden on manual hospital checks. Implantable biosensors are just an additional small step from wearable devices and it will probably be the future in healthcare regarding monitoring. Not only medicine can become personalized, this technology can provide treatment in real-time comprising the advancement of the field. Implanted biosensors can provide biochemical analysis and give feedback regarding treatment status and disease stage. In breast cancer, these devices could be implanted at the time of a biopsy, for example, to track chemotherapy agents, helping doctors understand whether cancer drugs are reaching the target tumors [140]. Also, information like pH or oxygen levels could be monitored to help understand the tumor metabolism and response to treatments [140]. These features, once achieved in a sensitive, safe, and precise devices, will improve global healthcare because it will allow real-time, personalized diagnosis and drug delivery, decreasing the time from presenting symptoms, diagnosis, and treatment. In addition, it will bring medicine closer to one of the main goals in breast cancer—early diagnosis.

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C H A P T E R

4 Electrochemical Biosensors for Antioxidants Juan Jose´ Garcı´a-Guzma´n1, David Lo´pez-Iglesias1, Mariana Marin2, Cecilia Lete2, Stelian Lupu3,*, Jose´ Marı´a Palacios-Santander1,* and Laura Cubillana-Aguilera1,* 1

Institute of Research on Electron Microscopy and Materials (IMEYMAT), Department of Analytical Chemistry, Faculty of Sciences, Campus de Excelencia Internacional del Mar (CEIMAR), University of Cadiz, Campus Universitario de Puerto Real, Polı´gono del Rı´o San Pedro, Puerto Real-Cadiz, Spain, 2 Institute of Physical Chemistry “Ilie Murgulescu” of the Romanian Academy, Bucharest, Romania, 3 Department of Analytical Chemistry and Environmental Engineering, Faculty of Applied Chemistry and Materials Science, University Politehnica of Bucharest, Bucharest, Romania

4.1 INTRODUCTION An antioxidant can be defined from a chemical point of view as “any substance that, when present in low concentrations compared to that of an oxidizable substrate, significantly delays or inhibits the oxidation of that substrate” [1,2]. In other words, antioxidants reduce the damage caused by reactive oxygen species (ROS; superoxide anion: O2  2, hydroxyl:  OH, peroxyl: ROO  , alkoxyl: RO  radicals, hydrogen peroxide: H2O2, hypochlorous acid: HClO, ozone: O3, and singlet oxygen: 1 O2), either being free radicals by themselves or causing the generation of other free-radical

species. In fact, antioxidants act by preventing the formation of these kinds of substances, scavenging them, or by promoting their decomposition [3]. From a physiological point of view, the role of antioxidants seems to be directly related to damage prevention of cellular components (DNA mutations, malignant transformations, etc.) as a consequence of chemical reactions directly involving oxygen and/or ROS. Typically, some peroxidase and dismutase enzymes (catalase and tyrosinase, and superoxide dismutase, respectively), some vitamins (C or ascorbic acid, E or α-tocoferol), metal ions (selenium) carotenoids (β-carotene), and polyphenols (flavonoids, anthocyanins, etc.)

*Corresponding authors.

Advanced Biosensors for Health Care Applications DOI: https://doi.org/10.1016/B978-0-12-815743-5.00004-4

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© 2019 Elsevier Inc. All rights reserved.

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[4
6], among others, are well-known antioxidants, being capable of counteracting the damaging effects of oxidation. A possible classification of the major antioxidants according to their origin, that is, exogenous (daily intake) or endogenous (synthesized by the organism itself, or “first line of defense”), can be found in Fig. 4.1 [7]. The mechanism of action of antioxidants, which can act at different steps of the oxidative radical process, can be summarized as [7]: In 2 In-In  1 In  Initiation In  1 L 2 H
! In 2 H 1 L  kiLH

L  1 O2
! L 2 OO  kperox

Propagation

L 2 OO  1 L 2 H
! L 2 OOH 1 L  kp

2L 2 OO 
! ½L 2 OO 2 OO 2 L Termination kt

½L 2 OO 2 OO 2 L-Nonradical species 1 O2 As can be seen, the mechanism implies the successive steps of initiation, propagation, and chain termination and can be described by taking into consideration the lipid peroxidation in cell membranes or foodstuffs [8]. Nevertheless, endogenous antioxidant defenses are insufficient to prevent damage completely. This first line of defense is typically composed of superoxide dismutase and H2O2-removing enzymes, metal-binding proteins and lowmolecular weight scavengers, such as glutathione, uric acid, bilirubin, coenzyme Q, α-lipoic acid, and melatonin, among others (see Fig. 4.1) [9]. Thus diet-derived antioxidants (also called exogenous antioxidants) like vitamins and polyphenols are critical in maintaining health [10]. These antioxidants can usually be found in vegetables [11,12] and fruits [13] and products derived from them, such as juice, wine, beer, coffee, tea, and others [14], and also in cosmetics [15
17]. Recently, much

antioxidant activity has been found in seaweed-associated bacteria as well, which could be of great interest for promoting blue economy and marine sustainability, since this kind of sea product might be used as food, for making drugs, and others as it is an issue still open for exploration with enormous potential [18]. The World Health Organization (WHO) advises that diets rich in fruit and vegetables promote good health. In fact, according to the WHO and supported by several research studies [19
21], a daily intake of at least 400 g of fruit and vegetables is recommended to reduce the risk of cardiovascular diseases [22]. Another possibility might consist of the intake of dietary supplements or pharmaceuticals. However, rigorous research has not demonstrated antioxidant supplements or drugs to be beneficial in preventing diseases, mainly due to the synergistic effects with other substances present in fruit and vegetables, differences in chemical composition of the antioxidants in the supplements versus those in foods, specificity of some antioxidants for some diseases, the amount of the antioxidant intake, and others [23
25]. In fact, no specific antioxidants have been recommended or offered by healthcare systems, neither have any been approved as therapy by regulatory agencies that base their decisions on evidencebased medicine. This is simply because, so far, despite many preclinical and clinical studies indicating the beneficial effects of antioxidants in many disease conditions, randomized clinical trials have failed to provide the evidence of efficacy required for drug approval [26]. Over the past decade, the key role of ROS species has been substantially evidenced in many fundamental cellular reactions. ROS may attack biological macromolecules, giving rise to protein, lipid, and DNA damage, cell aging and oxidative stress-originated diseases [27,28]. In fact, oxidative stress is suggested to be implied in the etiology of various disease states of an organism,

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

Endogenous antioxidants

H N N H

N H

– O

N

O

HOOC

H3N +

HN

N H

NH2

CH3

H3C CH3

HS O

O NH

O

H

Nonenzymatic

O

Exogenous antioxidants (diet)

H N

CH3

HO OH

COOH

H

HO

O

H

O

CH3

O

HO

OH HO

Uric acid

Histidine

Glutathione

H N

HO

O

H3C

NH

HN

NH

HN

CH3

H3C

OH

O

Vitamin D

HO

OH

O

Vitamin C

O

O

H2N

Vitamin A

O

N

H2C

N H

CH3

OO H3C

Carnosine

CH3

HN

HN

H2C

Vitamin E

B Carotene

O

Bilirubin

Melatonin

O

OH HO

OH

HO HO

Enzymatic

OH O

Resveratrol

O O

CH3

H3C OH S

H3C

O

H

S

O

Alpha Lipoic acid Catalase

CH3

O

6–10

Coenzyme Q10 Peroxidase

Caffeic acid

Superoxide dismutase

Terpene

O

Flavone (Zn)

DNA-repair enzymes DNA-glycosylase Apurin/apyrimidinic endonuclease

FIGURE 4.1 Schematic classification of major antioxidants according to their exogenous or endogenous origin.

(Se)

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being considered as crucial in the pathophysiology of common diseases including atherosclerosis, neurodegenerative diseases like Alzheimer’s, chronic renal failure, diabetes mellitus, and cancer [29]. Epidemiological studies have proven antioxidants ability to constrain the effects of ROS activity, and diminish the incidence of cancer and other degenerative diseases [7]. It is noteworthy to mention as well that scientists are also paying attention to reactive nitrogen species (RNS; nitroxyl anion, nitrosonium cation, higher oxides of nitrogen, S-nitrosothiols, and dinitrosyl iron complexes) as well as their implication in cell damage and death by inducing nitrosative stress at high concentration levels. RNS have been recognized as playing a crucial role in the physiologic regulation of many living cells, such as smooth muscle cells, cardiomyocytes, platelets, and nervous and juxtaglomerular cells. They possess pleiotropic properties on cellular targets after both posttranslational modifications and interactions with ROS [30,31]. In fact, both ROS and RNS, usually put together and known as reactive oxygen and nitrogen species (RONS), are considered to cause oxidative damage to biomolecules, contributing to the development of a variety of diseases [32,33]. Many different studies regarding the establishment of direct relationships between RONS and different illnesses have been conducted over the past few years affecting the liver [34], systemic autoimmune diseases [35], arthritis [36], a wide range of neurodegenerative diseases [37,38], immune system illnesses [9], and cardiovascular pathologies [39], among others. This chapter provides an overview of the current developments and achievements related to the use of electrochemical biosensors in the assessment of antioxidants in plants, food, and beverages, thereby critically showing their relationships with healthcare in most cases. The main reason for selecting electrochemical biosensors is due to their advantages versus other commonly used

analytical techniques like chromatography or mass spectrometry. Advantages include simple instrumentation, no sample treatment, high specificity, low-cost, rapid response, sensitivity, relatively compact size, and ease of implementation to detect biomolecules [40,41]. The first part of the chapter (Section 4.2) is devoted to the methods and procedures for ROS determination, with an emphasis on the applications of enzymatic biosensors. The second part (Section 4.3) focuses on the determination of antioxidants, mainly polyphenols, in plants and foods, by using electrochemical biosensors based on oxidase enzymes. The functioning principle of electrochemical biosensors is described together with selected applications of various tyrosinase-based and laccase-based biosensors applied in the assessment of polyphenols in real samples. The analytical performances and advantages of electrochemical biosensors compared to classical analytical methodologies are discussed. Finally, the last part (Section 4.4) pays special attention to the analysis of polyphenols in the most internationally popular beverages such as wine, beer, coffee, tea, and juice by using electrochemical biosensors.

4.2 BIOSENSORS FOR THE DETERMINATION OF REACTIVE OXYGEN SPECIES Oxidative stress is a consequence of the imbalance between oxidative processes induced by ROS and the antioxidant capacity of aerobic organisms. As introduced earlier, ROS is a general term referring mainly to different radicals (HO  ), superoxide radical anions (O22  ), and molecules containing oxygen (e.g., H2O2) which are highly reactive. Along with RNS, they form the defense system that protects the human body against different viruses and/or bacteria. In a healthy body, ROS are usually trapped by different lowmolecular weight (vitamin C, glutathione,

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4.2 BIOSENSORS FOR THE DETERMINATION OF REACTIVE OXYGEN SPECIES

vitamin E, lipoic acid, etc.) and high-molecular weight (enzymes) antioxidants. However, antioxidants cannot scavenge excessive ROS and oxidative stress is installed concomitantly with its harmful effects on the human health [42,43]. The incomplete reduction of oxygen is the main factor generating ROS, whereas their overproduction and the pathological consequences can be determined by different conditions such as environmental pollution (water, air, and soil), radiation, stress, smocking, and others. Generally, ROS play a dual role in the living organisms being involved in both “good” biochemical processes (redox regulation, defense against viruses and bacteria, modulation of vasodilatation) and in the “bad” ones (inflammatory processes, arterial hypertension, cancers, neurological disorders, aging, diabetes mellitus, etc.) [1]. Nowadays, modern medicine achieves superior results concerning the quality of human health by bringing together prophylaxis and efficient treatments. Permanently connected to the fast scientific progresses, medicine has applied different approaches, technologies and materials in order to assess the effects of oxidative stress on human health by monitoring ROS in various biological systems. Fluorescence and chemiluminescence are the most employed methods for determination of ROS, but specificity and selectivity are their main disadvantages. Moreover, they are not well suited for in vivo determinations owing to the bulky apparatus and, therefore cannot provide accurate information for medical purposes. Electrochemical methods have attracted considerable attention due to their outstanding features like rapidity, sensitivity, selectivity, miniaturization, low costs, simplicity, and portability. The in situ monitoring of ROS by means of electrochemical tools is probably the most important achievement of medicine, knowing that ROS are present in low concentrations and for an extremely short period of time in biological environments.

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4.2.1 Biosensors for Hydroxyl Radical Hydroxyl radical (HO  ) results from Fenton reactions of different transition metal ions such as Fe(II), Cu(II), and Cr(III) with H2O2 and/or from different biochemical reactions normally occurring in the human body. Owing to its short half-time, HO  is one of the most aggressive ROS, being involved in lipid peroxidation, inflammatory processes, neurological disorders induced by damaging of dopaminergic neurons, hyperthermia, arterial hypertension, and others. The monitoring of HO  is very important especially for clinical diagnostic purposes, but is challenging due to its high reactivity. One approach to track HO  in both in vivo and in vitro biological environments is the electron spin resonance (ESR). The employment of 4-hydroxy-2,2,6,6-tetramethylpiperidine-N-oxyl (hydroxy-TEMPO) as a spin probe [44] allowed to monitor HO  in the brain of mice and to assess their damaging effect toward dopaminergic neurons. 5,5-Dimethyl N-oxide pyrroline (DMPO), another spin probe, was used to achieve the ESR signal in the presence of HO  [45]. An optical sensor obtained by immobilization of nitrophenol on fiber-optic material enabled sensitive determination of HO  by reflectance spectroscopy for clinical analyses purposes and showed good stability over time (more than 6 months in conditions of dry storage) [46]. On the other hand, different fluorescence methods have been developed for HO  sensing in biological environments. The hybrid phenothiazine platform-based cyanine dye (MPT-Cy2) used as fluorescent probe for HO  determination in living cells and bacteria has proved good selectivity over other ROS and was employed for in vivo monitoring of HO  released in zebra fish in the presence of TiO2NPs under sunlight-like illumination conditions [47]. Another fluorescence sensor for HO  was based on coumarin-modified

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cyclodextrin derivatives and allowed sensitive and specific determination of HO  in water solution and yeast cells [48]. The rapid progress in nanotechnology and the outstanding features of nanoparticles have encouraged their application in the development of fluorescence biosensors systems for HO  . Thus, gold nanoclusters protected by bovine serum albumin (AuNCs-BSA) and conjugated with 2-[6-(40 -hydroxy)phenoxy-3Hxanthen-3-on-9-yl]benzoic (HPF) were successfully tested for HO  determination in living cells [49]. The fluorescence biosensor showed a linear response in the range 1
150 μM (R2 5 0.998) and a detection limit of 0.68 μM, with no interferences from other ROS or biologically active compounds usually present in biological samples. Another fluorescence nanobiosensor was obtained by immobilization of single-stranded DNA onto gold nanoparticles (AuNPs) surface via strong amide bonds [50]. The fluorescence intensity increased linearly with HO  concentration from 0.1 to 50 μM, the detection limit being 42 nM. The dyes employed in fluorescence measurements can promote potential damage toward human health or can generate false-positive results, thus becoming less attractive for determinations in biological environments. Electroanalytical approaches have received increasing attention thanks to their attractive features such as low cost, fast and simple analyses, possibility of miniaturization, and others. Meanwhile, the sensitivity and selectivity of the determinations can be optimized by choosing the appropriate electroanalytical method and sensing materials. A major problem encountered when electrochemical methods are employed for measurements in biological systems is the fouling of the electrode surface caused by the adsorption of either analyte or other compounds present in the sample. In order to avoid this situation different redox mediators can be immobilized onto electrodes

surface. The resulting modified electrodes are less prone to lose their electroactivity and generally show superior analytical performance over unmodified electrodes. An amperometric sensor obtained by the deposition of doped polyaniline (PANI) onto a gold electrode showed a linear response toward HO  . The sensor response, that is, electrical conductivity of PANI, changed with HO  concentration and was linear in the range from 0.2 to 0.8 μM [51]. Different biologically active compounds such as enzymes, nucleic acids, proteins, and others have been employed for the construction of biosensors. Over the past few decades, biosensors have received increasing attention for biomedical applications as they enable to perform the clinical analysis and to monitor in small sample volumes with good sensitivity and selectivity. Different electrochemical enzymeless biosensors have been obtained by immobilization of flavonoid compounds on the surface of an optically transparent electrode. Thus, quercitin (quer), primuletin (prim), and morin (mor) deposited on the surface of (3-aminopropyl)triethoxysilane(APTES) fluorine-doped tin oxide (FTO) electrode leads to voltammetric biosensors that allowed the rapid determination of HO  with low detection limits, that were 2, 3, and 5 nM, respectively, (see Table 4.1) and good selectivity with respect to other potential interfering ROS [52]. DNA-based biosensors for HO  have been developed starting from the harmful potential of ROS toward nucleic acids. The construction of these biosensors involves the immobilization of nucleic acids either on the surface of an unmodified electrode or onto an electrochemical transducer previously modified with a redox mediator. When nucleic acids have been adsorbed onto carbon-based electrodes, the resulting DNA-biosensors showed good sensitivity toward HO  and was also successfully applied in order to assess the

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TABLE 4.1 Analytical Performances of Different Sensors and Biosensors for HO  Determination Sensor

Technique

Linear Range (μM)

Detection Limit (μM)

References

MPT-Cy2

Fluorescence

1
10

1.16

[47]

AuNCs-HPF

Flourescence

1
150

0.68

[49]

AuNPs-DNA

Fluorescence

0.1
50

0.042

[50]

Au/PANI

Amperometry

0.2
0.8

0.2

[51]

quer/APTES-FTO

SWV

0
0.05

0.002

[52]

mor/APTES-FTO

0.003

prim/APTES-FTO

0.005

AuNPs/MCH/DNA1/Au

SWV

5
10000

3

[56]

MBs-DNA-Ag

ASV

0.05
4

0.01

[57]

APTES-FTO, ((3-aminopropyl)triethoxysilane-fluorine doped tin oxide) electrodes; ASV, anodic stripping voltammetry; AuNCs, gold nanoclusters; AuNPs, gold nanoparticles; HPF, 2-[6-(4-hydroxy)phenoxy-3H-xanthen-3-on-9-yl]benzoic acid; MBs, magnetic beads; MCH, 6-mercaptohexanol; mor, morin; MPT-Cy2, cyanine dye based on a hybrid phenothiazine platform; PANI, polyaniline; prim, primuletin; quer, quercetin; SWV, square wave voltammetry.

antioxidant capacity of different antioxidants [53
55]. The outstanding features of nanomaterials have encouraged their application in DNA-biosensor development in order to improve their analytical performances. The voltammetric response of the DNA-Au biosensor was enhanced after its subsequent modification by DNA-functionalized AuNPs. An increment of the linear response in the range from 5 μM to 10 mM and a lower limit of detection (eight times) have been obtained when DNA1/Au biosensor (DNA1: 5_-SH(CH)6-GGT CCG CTT GCT CTC GC-30 ) was replaced by DNA2-AuNPs/MCH/DNA1/Au (DNA2: 5_-SH-(CH)6-CGG GCG AGA GCA AGC GGA-30 ) for the assessment of HO  , where MCH is 6-mercaptohexanol [56]. The enhancement of electroanalytical nanoparticlebased DNA-biosensor performances is due to a more effective surface area that AuNPs show against their bulk homologue. Another selective nanoparticle-based biosensor for HO  was obtained by functionalization of DNA-immobilized magnetic beads (MBs) with

silver nanoparticles (AgNPs) [57]. The working principle of the MBs-DNA-AgNPs consisted in the oxidation of the nucleic acid and the breakage of the strands from DNA-nanobiosensor followed by AgNPs detaching from the nanobiocomposite surface. The as-“released” AgNPs were subsequently detected by the anodic stripping voltammetry (ASV) technique, proving the peak current to be proportional to the concentration of HO  . Thus, the linear response range (0.05
4 μM) and the limit of detection (10 nM) were significantly improved by choosing appropriate electroanalytical method and biosensor components. These features together with good reproducibility and selectivity of MBsDNA-AgNPs electrochemical biosensor against the competing ROS species strongly recommend its application for HO  quantification and for the assessment of antioxidant capacity for different antioxidants. Table 4.1 summarizes the main results concerning the analytical performance of different sensors and biosensors for HO  determination.

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4.2.2 Biosensors for Superoxide Anion Radical Superoxide anion (O22  ) is an intermediate of molecular oxygen reduction. O22  and together with other ROS has an important contribution to immune response of living systems. On the other hand, O22  is very unstable and can interact rapidly with different biologically active compounds and/or other radicals containing oxygen, leading to different pathological effects. Hence, the reaction of O22  with NO  is responsible for vasoconstriction [58] whereas that with NO produces peroxynitrite, which is involved in neurodegenerative diseases [59]. Under normal physiological conditions O22  appears at low concentrations owing to its efficient scavenging by endogenous antioxidants. The overproduction and/or accumulation of O22  , when the concentration is expected to be in the range of 1026
1027 M [60], leads to oxidative stress and harmful effects on human health including cancer, neurodegenerative diseases, arterial hypertension, and others. In other words, the border between the physiological and pathological conditions is a function of O22  concentration, location, and duration. An efficient assessment of health status and/or disease progress assumes permanently monitoring of O22  by clinical analysis and various imaging methods that involve the use of sensitive, selective, and biocompatible analytical tools for the detection of O22  . One approach to detect O22  in the vascular wall is chemiluminescence, achieving a low detection limit (20 nM) with negligible cellular toxicity. However, this approach shows some limitations regarding the concentration of probe and existence of antioxidants in the sample. The fluorescence generated in the presence of dyes such as dihydroethidium (DHE), its derivatives, [61,62], and 20 ,70 -dichlorodihydrofluorescein [63] has been successfully applied for detection of O22  in vascular cells. The fluorescence

nanosensor obtained by microemulsion polymerization [64] and the nanobiosensor developed by immobilization of horseradish peroxidase (HRP) and superoxide dismutase (SOD) [65] have showed good sensitivity and specificity toward O22  quantification. However, the electrochemical methods are very attractive for clinical monitoring due to their specific features that allow the rapid and selective determination of O22  both in small volumes of samples or onsite during physiological and/or pathological processes in living systems. Biosensors are powerful analytical devices for the determination of very reactive species in biological samples, both in vivo and in vitro, owing to their specificity, fast response, and small size affording spatial and time resolution. Selective electrochemical biosensors have been designed by employing enzymes, proteins, or permeable membranes deposited on the surface of different conducting materials and applied to O22  monitoring either for clinical analysis or imaging purposes. Cytochrome c (Cyt c) is a small redox protein that significantly increases the electrochemical response when immobilized onto electrode surfaces. A nonenzymatic multilayer structure very sensitive to O22  was obtained by alternating deposition of Cyt c and poly(aniline(sulfonic acid)) (PASA) onto Au wire electrodes [66]. Under optimum conditions (six layers), the nonenzymatic Au/PASA/Cyt c electrochemical biosensor showed a linear response in the range of 0.4
1.5 μM O22  and a sensitivity of 0.398 A/mol cm2. The amperometric biosensor obtained by covalent immobilization of Cyt c onto an array of Au electrodes allowed the determination of O22  produced by stimulation of A172 human glioblastoma cells with a sensitivity of 10.3 nA/μM cm2 and no interferences from ascorbic acid and H2O2 [67]. Superoxide dismutase (SOD) is an enzyme that catalyzes the dismutation of O22  to O2 or H2O2 and plays an important role in the defense systems of living organisms by ROS



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4.2 BIOSENSORS FOR THE DETERMINATION OF REACTIVE OXYGEN SPECIES

scavenging. Because of its specificity and high reactivity toward O22  , SOD has become very attractive in the development of sensitive and selective biosensors. A third-generation biosensor obtained by immobilization of SOD onto a Au electrode modified by cysteine (Au/Cys) was employed for the amperometric determination of O22  with good sensitivity and no interferences from H2O2, uric acid, ascorbic acid, and DOPA C [59]. The entrapment of SOD within polypyrrole (PPY) matrix during electropolymerization of the monomer onto a platinum (Pt) electrode was used to obtain an amperometric microbiosensor with good analytical performances toward O22  [60]. Thus, the wide linear response range and the detection limit of 15 nM together with the small size have determined the use of Pt/PPY/SOD microbiosensors in the quantification of O22  in abdominal aorta of dogs after stimulation. New enzymatic biosensors with good electrocatalytic properties starting from the outstanding properties of the nanomaterials have been also reported. The covalent bonding of SOD to Fe3O4 nanoparticles followed by their immobilization onto the Au surface allowed to achieve a detection limit of 11.5 nM for O22  [68]. Furthermore, three different nanostructured Au electrodes (pyramidal, spherical and rod-like) were modified with SOD. They were successfully applied for the amperometric determination of O22  over a wide range of concentrations and with low-detection limits, depending on the applied potential [69]. Another selective approach enabling the assessment of O22  in ovarian cancer cells was based on the immobilization of the biomimetic enzyme-like Mn3(PO4)2 on TiO2 nanoneedles by using Nafion membranes. The synergistic effects generated by employing biomimetic enzymes and highly conductive TiO2 have provided superior performances to the sensor such as a wide linear range (5 3 1027
1.5 3 1023 M) and a low detection limit of 170 nM [70].

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4.2.3 Biosensors for H2O2 H2O2 is another ROS with mild oxidant power. It can be involved in physiological processes such as redox signaling or immune defense system, and can be responsible for different pathologies like cancer, myocardial infarction, or Alzheimer’s disease. Its dual role is a consequence of its concentration, location, and/or accumulation in target organs. In the human body, H2O2 can result as a response of different cells to various bacteria and viruses or by O22  scavenging in the presence of SOD. Catalase and peroxiredoxin are the enzymes responsible for maintaining H2O2 under physiological levels, thus avoiding the occurrence of pathological effects. The quantification of H2O2 in different biological samples (tissues, cells, plasma, serum, urine) is an attractive approach to assess the human health and/or the related diseases progression. Optical methods such as fluorescence [71
74] and chemiluminescence [75,76] are sensitive to H2O2 and have been the first ones employed for its determination and for the assessment of the antioxidant activity for different compounds. Despite their sensitivity, these methods have become less attractive as are not selective and usually employ dyes with potential harmful effects. H2O2 is a redox active molecule, and therefore electrochemical methods are well-suited for its monitoring. The continuous efforts and scientific achievements in the field of materials science have enabled the development of new sensing electrochemical tools with superior activity toward H2O2, that were employed not only for in vitro analysis, but also for in vivo monitoring of H2O2 released from living cells. Various electrochemical sensors for H2O2 determination have been reported in the literature and an increased prevalence of electrochemical methods within the methodology dedicated to H2O2 monitoring was observed

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over the past few decades. The main concerns regarding the electrochemical detection of H2O2 are the potential interference of other biocompounds existing in biological samples and/or the fouling of the electrode surface that can affect the accuracy of measurements. There are different ways to overcome these drawbacks and to obtain reliable results such as the selection of appropriate approach (e.g., determination by electrochemical reduction instead of oxidation) and the employment of redox mediators and/or different permselective membranes. The electrochemical reduction of H2O2 takes place at negative potentials, where the others compounds present in the sample are not electroactive. Moreover, the use of small-sized sensing tools (providing spatial and temporal resolution) and redox mediators showing electrocatalytic activity have enabled new achievements in medicine. Thus a platinized-carbon microelectrode has allowed the monitoring of H2O2 released by cells under physical polarization and has also served to highlight when no H2O2 is developing in the case of singleimmunostimulated macrophage cells [77]. Nanosized carbon-based materials deposited on different conductive surfaces have been successfully employed in the development of sensitive electroanalytical tools for H2O2 determination in biological systems. The indium tin oxide electrode (ITO) modified in a first step by single-walled carbon nanotubes (SWCNTs) and by osmium bipyridine in a second step, was used for the monitoring of H2O2 released in macrophage cell (RAW 264.7) under bacterial stimulation [78], without interferences from other ROS, and offered new insights regarding the mechanism of the immune response inside the cells. The amperometric sensor obtained by immobilization of multiwalled carbon nanotubes (MWCNTs) onto platinum black electrodes was employed to acquire information concerning the level of H2O2 in

aging of nervous central systems [79] thanks to the possibility to cross the cell membrane. Since its discovery in 2004, graphene oxide (GO) has attracted much attention due to the outstanding properties such as electrical conductivity and low weight. The electrical conductivity of GO increases by chemical or electrochemical reduction, and the additional modifications of the resulting reduced graphene (RGO) lead to different sensors and biosensors for H2O2 monitoring. The functionalization of RGO by CuS allowed a linear response of the resulted sensor (CuS/RGO/ GCE) on a wide range of concentration (5
1500 μM) and a detection limit of 0.27 μM, without interferences from other biologically active compounds present in the biological fluids [80]. The sensor was tested with good results in human serum, urine and HeLa cells. Following the same line of action, the biosensor obtained by immobilization of Cyt c onto the surface of RGO/GCE was used for the selective amperometric detection of H2O2 [81]. Other Cyt c
based biosensors for determination of H2O2 in biological systems have been obtained by deposition of Cyt onto different nanosized TiO2 structures [82
84]. Myoglobin, hemoglobin, and ferredoxin can be also utilized for the construction of H2O2 biosensors [85]. Nanomaterial-based biosensors are very attractive sensing tools due to the synergistic effects obtained by combining the outstanding features of nanomaterials (small size and increased active surface) with substrate specificity that enzymes and proteins offer. The nanobiosensor designed by means of HRP immobilization onto AuNPs [86] and ZnONPs [87] has demonstrated electrocatalytic activity toward H2O2 reduction. Also a 3D biomimetic electrochemical biosensor was obtained by immobilization of MnO2 within the pores of PPY and successfully employed to detect H2O2 released from cells [88].

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4.3 ELECTROCHEMICAL BIOSENSORS FOR THE ASSESSMENT OF TOTAL ANTIOXIDANT CAPACITY

ROS are normally present in living aerobic organisms and are involved in various biochemical reactions corresponding either to physiological or pathological states. The in vivo and/or in vitro monitoring of ROS are valuable tools for the assessment of human health and allow more accurate diagnostics and effective therapies of diseases that improve the quality of people’s lives. Electrochemical biosensors demonstrate to be remarkable analytical tools for selective and sensitive determination of different ROS, being successfully applied to track the oxidative stress and related pathologies even at the level of a single cell.

4.3 ELECTROCHEMICAL BIOSENSORS FOR THE ASSESSMENT OF TOTAL ANTIOXIDANT CAPACITY OF PLANTS AND FOODS Plants are a rich and natural source of antioxidant compounds. Especially medicinal plants containing a tremendous variety of natural antioxidants such as flavonoids, phenolic acids, and tannins, which possess more powerful antioxidant activity than common dietary plants. Owing to their antioxidant and freeradical scavenging activity they are used in adjuvant therapy for the treatment of antiinflammatory, antitumor, antiallergic, antiviral, and antibacterial processes. The presence of polyphenolic and phenolic compounds in plants assures protection against pathogen attacks or ultraviolet radiation. Several studies have demonstrated that the consumption of some cereal products is of importance for the prevention of diabetes, cancer, and cardiovascular diseases. The phenolic compounds prevent oxidative damage to cellular organelles, proteins, lipids, DNA, and RNA [89]. For example, oat contains different kinds of phytochemicals with antioxidant activities as flavonoids, phenolic acids, and tocols [90].

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Also, this kind of cereal includes antioxidants with low-molecular weight, and substituted N-cinnamoylanthranilic acids that are named avenanthramides (AVs) [91]. Emmons and Peterson [92] have shown that these acids have a significant antioxidant capacity, with 10
30times higher radical scavenging activities than caffeic acid, ferulic acid, and vanillin. Nie and Wise [93] discovered that AVs have an antiproliferative influence on vascular smooth muscle cells, limiting the development of restenosis and atherosclerosis after angioplasty. These phytochemicals in plants exhibit health protection against several biomolecules damages. Daily, at the cell level around 2 3 104 DNA damaging events occurs leading to diseases such as arteriosclerosis, hypertension, diabetes, and neurodegenerative disorders [94,95]. The damage could be assigned to ROS  as HO  , H2O2, and superoxide radical (O2 2 ) [96]. Thus the monitoring of ROS has attracted a great deal of interest over the past few years. For instance, it was demonstrated that a DNAbased biosensor against ROS is a useful approach in total antioxidants capacity assessment because the principle of measurement is closer to the antioxidant activity in biological systems [97]. The radical attack on DNAmodified electrodes is similar to the process that occurs within the cell, which may generate replication errors and subsequent misleading protein synthesis that produce cell ageing. Consequently, the development of analytical methodologies for the quantification of the antioxidants or evaluation of the antioxidant capacity in plants, food, and beverages is of paramount importance. The analytical methods could be classified as hydrogen atom transfer (HAT) and single electron transfer (SET) methods based on the reaction mechanisms involved. Both mechanisms almost always occur together according to their antioxidant structure and pH [98]. HAT-based methods used in the literature are total radical-trapping antioxidant parameter (TRAP), total oxidant scavenging capacity (TOSC), oxygen radical

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absorbance capacity (ORAC), low-density lipoprotein (LDL) oxidation, and β-carotene bleaching by ROO  . SET-based techniques used in antioxidant capacity determination are total phenolic assays by Folin-Ciocalteu (FC), ferric reducing antioxidant power (FRAP), trolox equivalent antioxidant capacity assay (TEAC), and DPPH-based assay (2,2-diphenyl-1-picrylhydrazyl) [99]. In the Introduction, Fig. 4.1 illustrates most studied antioxidants classified according to their exogenous or endogenous characters. Due to the utilization of herbal extracts in medicine and food technology, it is important to evaluate their antioxidant capacity and this issue has been performed by a range of analytical methodologies including liquid or gas chromatography, capillary electrophoresis, spectrophotometry, fluorescence, and mass spectrometry [100]. In spectrophotometric assay, antioxidant activity of plant extracts is often determined against different freeradical species, such as the stable free radical 1,1-diphenyl-2-picrylhydrazyl (DPPH), and the bleaching rate is monitored at a characteristic wavelength in the presence of the plant extract. These methods provide precise quantification of several antioxidants in various samples, but they are time consuming and sometimes sample pretreatment is required. The use of electrochemical methods like cyclic voltammetry and differential pulse voltammetry in the quantification of various antioxidants has several benefits like simplicity, high sensitivity, rapidity of analytical measurements, and analysis of colored samples without pretreatment, compared to traditional analytical methodologies [101]. The electrochemical parameters of importance in antioxidants determination via electrochemical methods are: (1) the anodic (oxidation) peak potential, Epa; (2) the anodic (oxidation) peak current, ipa; and (3) the electric anodic charge, which is related to the area under the anodic oxidation wave, Qa. The Epa value is related to

the electron donor capability of the measured antioxidant, while the ipa parameter refers to the concentration of the antioxidant. The electric anodic charge has been successfully used to quantify the antioxidant capacity usually expressed as total polyphenolic content. The electrochemical methods are typically applied in connection with metals (Pt, Au) or glassy carbon electrodes and semiconductors, as well as electrodes modified with various nanomaterials and organic polymers to improve the selectivity and the sensitivity of the analytical measurements. The judicious modification of conventional electrode substrates with a range of inorganic, organic, and composite materials represents the most important achievement and development in the electrochemical science. The modification of the electrode surface revealed the possibility to achieve new properties like selectivity, sensitivity, and polarity that greatly expand the final applications of obtained modified electrodes. Chemically modified electrodes (CMEs) are usually obtained by deposition of chemical modifiers, that is, inorganic, organic, and polymeric compounds, in forms of monomolecular, multimolecular, and polymeric layers. The CMEs are actually functioning as electrochemical sensors and have found several applications in the electroanalysis of antioxidants. Another approach consists in the use of electrochemical biosensors based on oxidase enzymes such as tyrosinase, laccase, and HRP for the quantification of various antioxidants. Electrochemical biosensors represent a subclass of electrochemical sensors and have the advantages of both the sensitivity of the electrochemical transducers (electrode substrates) and the selectivity of the biological recognition element, namely, the enzymes. The immobilization of the enzymes onto electrode surfaces is of paramount importance for the proper functioning of the electrochemical biosensors. The main enzyme immobilization procedures are adsorption, entrapment into a matrix,

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

4.3 ELECTROCHEMICAL BIOSENSORS FOR THE ASSESSMENT OF TOTAL ANTIOXIDANT CAPACITY

microencapsulation, cross-linking, and covalent bonding. The incorporation of enzymes within conducting polymer-based matrices has attracted great interest because of the simplicity, versatility, and reproducibility of this procedure. PPY, PANI, PEDOT, and their derivatives have been successfully and extensively used in the fabrication of enzymatic electrochemical biosensors [102
104]. The use of tyrosinase [105
107] and laccase [108,109] for the detection of antioxidants like polyphenols is very convenient since the enzymatically generated quinone derivative may be reduced at low potentials, thus the electrochemical interferences are considerably minimized. The reduction current of the generated quinone derivative is measured and is related to the concentration of the investigated polyphenols. The functioning principle of the electrochemical biosensors based on oxidase enzymes for polyphenols detection is schematically depicted in Fig. 4.2. Tyrosinase (polyphenol oxidase, E.C. 1.14.18.1), in the presence of oxygen, catalyzes the hydroxylation of monophenols to o-diphenols (Reaction 1), and the oxidation of o-diphenols to o-quinones (Reaction 2) [110]. The resulted quinone derivatives can be easily detected by electrochemical reduction at low potential values (Reaction 3). This is the basic principle of tyrosinase-based amperometric biosensors and can be described by these reaction schemes:

117

mono-phenol1TyrosinaseðO2 Þ-o-diphenols (4.1) o-diphenols1TyrosinaseðO2 Þ-o-quinones1H2 O 2

1

o-quinones12e 12H -o-diphenols

(4.2) (4.3)

Laccase (EC 1.10.3.2) catalyzes the oxidation of phenol, diphenols, and various polyphenols to quinone derivatives and does not require the H2O2 as a cosubstrate [111,112]. Similarly to tyrosinase-based amperometric biosensors, the reduction of the enzymatically generated quinone derivatives provides the electrochemical signal in laccase-based biosensors. Actually, the antioxidant capacity is measured using a standard compound like caffeic acid, catechin, chlorogenic acid, or catechol, and this compound displays good electrochemical behavior at the electrode surface. The analytical performance of tyrosinase and laccase-based amperometric biosensors depends mainly on the enzyme immobilization method, the enzyme loading and activity, and pH of the sample solution. The immobilization method is the most important parameter in the development of enzyme-based electrochemical biosensors. The main achievements in this topic will be discussed taking into account the enzyme immobilization procedure. For instance, the adsorption of enzymes is a simple and versatile approach in biosensors construction [105]. In this study, the beneficial role of FIGURE 4.2 The functioning mechanism of electrochemical biosensors based on oxidase enzymes.

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the immobilization matrix alongside to the enzyme adsorption procedure by covering the adsorbed tyrosinase enzyme with polyethylene glycol and/or the ion-exchanger Nafion was demonstrated. The use of Nafion coating ensured the highest analytical performance in synthetic samples (low detection limit of 0.064 μM, sensitivity of 8 3 103 nA/μmol L cm2, Michaelis-Menten constant of 67.1 μmol/L for catechol), as well as in real samples. In another study [113], a tyrosinase-based biosensor was constructed by the immobilization of the enzyme onto screen-printed electrodes using various methods such as crosslinking with glutaraldehyde, entrapment with polyvinyl alcohol, and cross-linking with glutaraldehyde and human serum albumin. The best analytical performance in terms of the lowest detection and quantification limits (1.5 μM and 5.1 μM catechol, respectively) were obtained for the biosensor prepared via cross-linking of tyrosinase with glutaraldehyde. This biosensor was successfully applied in the determination of trolox equivalent antioxidant capacity of infusions prepared with various medicinal plants and the results were compared with a well-established method, namely, the DPPH spectrophotometric method. The incorporation of the enzyme within carbon-paste electrodes is a versatile and lowcost approach applied successfully in the construction of biosensors. The use of ionic liquids in conjunction with carbon paste demonstrated higher stability and sensitivity of the obtained electrodes. As an example, a laccase-biosensor for rosmarinic acid determination was fabricated using carbon paste and 1-N-butyl3-methylimidazolium hexafluorophosphate (BMIPF6) ionic liquid [108]. The quinone derivative produced in the enzymatic reaction was electrochemically reduced at 10.2 V versus Ag/AgCl and detected by using square-wave voltammetry. The optimized biosensor displayed a linear response toward rosmarinic

acid in the concentration range of 0.99
65.4 μM, a very low detection limit of 0.188 μM, and very good sensitivity and stability. The laccase-biosensor was used in the determination of rosmarinic acid in plant extracts with very good recovery values ranging from 96.1% to 105.0% and good agreement with the data obtained via capillary electrophoresis method at the 95% confidence level was observed. Besides these analytical performances, the preparation procedure based on the incorporation of the enzyme within a carbon paste with ionic liquid has several advantages like low cost, simplicity, rapidity, and renewability of the obtained laccase-biosensor. Another biosensor construction approach consists in the entrapment of the enzyme within a polymer membrane aiming to improve the stability of the biosensor. For instance, a laccase-based biosensor was built by immobilization of the enzyme within a polyvinyl alcohol polymer membrane onto graphite screen
printed electrodes [109]. The immobilization procedure ensured very good stability and excellent analytical performance in terms of detection limit of 0.5 μM caffeic acid and sensitivity of 24.9 nA/μM. The laccase biosensor was successfully applied in the determination of phenolic content of tea infusions with minimal sample preparation by using chronoamperometric detection mode and standard addition analytical protocol. The results obtained with the laccase biosensor were compared with those gathered by the Folin-Ciocalteu spectrophotometric method. The proposed laccase biosensor has several advantages like simplicity, rapidity, sensitivity, and direct analysis of real samples. Alternatively, the incorporation of enzymes within conducting polymers matrix during the electrochemical polymerization of the corresponding monomers onto various electrode substrates represents an efficient and simple biosensor construction procedure [103,114
116].

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In this case, the immobilization of the enzyme occurs mainly via electrostatic interactions between the positively charged backbone of the conducting polymer and the negative charges carried out by the enzyme at a given pH value. Even if there is some diffusion limitation within the polymer matrix of the reactive species, the rapidity and simplicity of this preparation method ensured its wide use in the development of enzyme-based electrochemical biosensors. Selected applications of tyrosinase,

horseradish peroxidase, and laccase-based biosensors in the determination of various antioxidants and mainly polyphenols in plants, foods, and beverages are presented in Table 4.2. The data displayed in Table 4.2 point out to the peculiar analytical performances of the enzyme-based electrochemical biosensors, namely, the very low detection limit values and the wide linear response ranges that can be achieved. The possibility to directly measure real samples with no (or minor)

TABLE 4.2 Enzymatic-Based Biosensors Used in the Determination of Various Antioxidants Linear Limit of Antioxidant Range (μM) Detection (μM) References

Enzyme

Immobilization Procedure

Sample

Tyrosinase

Nafion coated sonogel-carbon electrode

Beers and industrial wastewaters

Catechol




0.064

[105]

Tyrosinase

Multiwalled nanotube ionic liquid-chitosan coated on ITO electrode

Red wines

Phenol

10
80




[106]

Tyrosinase

On diazonium-functionalized Tea screen-printed gold electrodes

Catechol

0.1
22

0.1

[107]

Tyrosinase

Cross-linking with glutaraldehyde

Medicinal plants

Catechol

Up to 136

1.5

[113]

Tyrosinase

Cross-linking onto polypyrrole

Environmental water

Phenol




1.71

[117]

Tyrosinase

Entrapment within PEDOT

Synthetic sample Catechol

20
300

12.9

[116]

Horseradish peroxidase

Immobilization on selfassembled monolayers

Wine and tea

Catechin

Up to 25

2

[118]

Laccase

Incorporation with ionic liquid into carbon paste

Plant extracts

Rosmarinic acid

0.99
65.4

0.188

[108]

Laccase

Entrapment in polymer membrane onto graphite screen
printed electrodes

Tea

Caffeic acid

0.5
130

0.5

[109]

Laccase

Physical adsorption

Plant extract

Caffeic acid

Up to 10

0.56

[119]

Laccase

Covalent immobilization on dendrimers

Tea

Catechin

0.1
10

0.05

[120]

Laccase

Entrapment in nanocomposite Salvia officinalis multiwall carbon nanotubesand Mentha chitosan piperita extracts

Rosmarinic acid

0.9
12.1

0.23

[121]

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pretreatment of the sample constitutes the main appealing feature of electrochemical biosensors. Furthermore, the low-cost instrumentation, possibility of miniaturization, and rapidity and simplicity of the measurements represent other advantages of electrochemical biosensors compared to conventional analytical methodologies.

4.4 BIOSENSORS FOR THE ANALYSIS OF POLYPHENOLS IN BEVERAGES As it has been previously mentioned, polyphenols possess an important role in the antioxidant capacity. This antioxidant capacity has a strict relationship with several health benefits. Thus the determination of polyphenols, not only in plants but in the very common products made by them, is a critical factor. This section tries to show several approaches which can be used to solve this topic out. Some of the principles used have been briefly mentioned in the previous section; however, the approach will be slightly different, being mainly focused on beverages such as beers, wines, teas, juices, and others.

4.4.1 Beverages: Role of Antioxidant Capacity for Healthcare Purposes In the last decades the interest for relating the health benefits of beverages and the prevention of diseases is increasing. Numerous research studies show that moderate consumption of a certain beverages (e.g., wine, beer, coffee, juice, teas) produces positive benefits in the human body. However, these drinks are highly consumed, as shown in Fig. 4.3. Tea and beer seem to be the most-consumed beverages in the world. One of the most representative studies of this topic is the “French paradox.” In 1992 Renaud and De Lorgeril [122] pointed out the relation between wine consumption and health from the study of the mortality statistics of seventeen countries, such as Japan, France, Switzerland, Spain, Italy and Ireland, among others. In some of these countries (i.e., France) the intake of lactic fats was very high, although the mortality caused by heart disease was much lower than could be expected. Wine was considered as being responsible for longevity due to its high antioxidant capacity and its medium-high consumption. Grapes and, consequently wines, are wealthy sources of antioxidants. It has been demonstrated that the antioxidant presence in

FIGURE 4.3 Worldwide consumption of the most popular beverages in 2016.

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4.4 BIOSENSORS FOR THE ANALYSIS OF POLYPHENOLS IN BEVERAGES

this fruit has protective properties against several diseases [123]. They also participate in the slowing of thrombosis procedures due to the platelet aggregation inhibition, lipids peroxidation, oxidation of proteins of low density, and others. Moreover, their activity against oxidative diseases such as cellular aging, mutations and even cancer, is well-known [124]. Other beverages, such as teas, present anticarcinogenic effects, participating in the scavenging of reactive oxygen species (ROS) [125]. In addition, it has been demonstrated that they have an important role in the prevention of many diseases such as heart diseases and obesity [126]. Several procedures have been applied in the determination of the antioxidant capacity of beverages. The usual approach is the imitation of the radical scavenging, which can be found in the sample, between the native antioxidants and the ROS species. A chromophore reacts with a strong oxidant, obtaining a colored radical compound. The scavenging phenomenon leads to a decay in the concentration of the colored radical attributed to the antioxidant capacity. By spectrophotometric assays, a change of the signal in the UV-vis spectrum is found. Different radicals species can be used such as (2,20 -azino-bis(3-ethylbenzothiazoline-6-sulfonic acid)) ABTS [127], (dimethyl-4-phenylenediamine) (DMPD) [128], (2,2-diphenyl-1-picrylhydrazyl) (DPPH) [129], ferric reducing ability of plasma (FRAP) [130], and cupric reducing antioxidant capacity (CUPRAC) [131], among others. Fluorescence properties can be also used for this purpose, for example the oxygen radical absorbance capacity (ORAC) assay [132]. Some of these methods which have been applied to plants and food samples, have been previously commented in Section 4.3. A similar approach is the employment of DNA biosensors. These biosensors monitor the damage to DNA chains caused by radical species, analogous to the scavenging capacity of the compounds used in other methods. DNA biosensors will be explained deeply in Section 4.4.2.2.

121

Other alternative approaches focused on electrochemical devices have emerged. They can be classified according to two different principles. The first is based on the direct relation between an electrochemical signal and the antioxidant capacity of the sample [133] and this relation generates a new parameter, namely, the electrochemical index (EI). However, the main drawback of this procedure resides in the high amount of interferents found in complex real samples. The second principle is based on the important role of the polyphenolic content in the antioxidant capacity of beverages. A polyphenol index (Ip), usually calculated using certain standard polyphenol, is correlated with the antioxidant capacity of the sample. Wine, coffee, beer, and juice constitute a rich source of polyphenols. For instance, wine contains a large number of different polyphenols, which can be split in two groups, nonflavanoids and flavonoids. Nonflavanoids include polyphenols like gallic, caffeic, and ferulic acid. On the other hand, flavonoids contain tannins, anthocyanins, or flavanols such as catechins. All these polyphenols have been widely studied for different reasons. In the majority of cases, the antioxidant capacity of the beverages can be attributed to o-diphenols. Gallic acid, caffeic acid, and catechins, among others, have been widely studied [134
139]. Inside this group, some authors indicated the higher antioxidant capacity of gallic acid and caffeic acid over the others [140]. This study corroborates the use of these polyphenols as standard polyphenols to recalculate the Ip in electrochemical assays. Wine has the highest polyphenols content, although beer, tea, coffee, and juice also possess a significant amount of these polyphenols. Consequently, the study of these beverages is a very prolific field [97,141
147]. Table 4.3 shows the most relevant polyphenols and the beverages in which they mainly appear.

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TABLE 4.3 The Most Relevant Polyphenols in Beverages Polyphenol

Chemical Structure

Beverage

Quercitin

Juice and wine

Chlorogenic acid

Coffee

Caffeic acid

Coffee, wine, and beer

Catechin

Juice, wine, tea, and beer

Epicatechin

Juice, wine, and tea

(Continued)

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TABLE 4.3 (Continued) Polyphenol

Chemical Structure

Beverage

Epicatechingallate

Tea

Epigallocatechin- gallate

Tea

Gallic acid

Wine

(Continued)

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

TABLE 4.3 (Continued) Polyphenol

Chemical Structure

Beverage

Salycilc acid

Beer

Theaflavin

Tea

Theaflavin-3-gallate

Tea

Theaflavin-3,3’-digallate

Tea

(Continued)

125

4.4 BIOSENSORS FOR THE ANALYSIS OF POLYPHENOLS IN BEVERAGES

TABLE 4.3 (Continued) Polyphenol

Chemical Structure

Beverage

Vanillic acid

Beer and wine

4.4.2 Electrochemical Biosensing of Polyphenols in Beverages There are two main trends in polyphenols determination in these kinds of samples. The first is focused on the individual determination of these polyphenols. For this purpose, chromatography techniques [136,148] have an important role due to their power and good sensitivity. Other techniques such as electrophoresis [149] or mass spectrometry [150] can be used for their determination as well. However, all these procedures are very expensive and time-consuming. Besides, the individual determination of all polyphenols in beverages is, at best, a hard task. Some authors indicate that the contribution in the antioxidant capacity of some groups of polyphenols is higher than others, so the total polyphenolic content is not so important as the polyphenolic content of certain groups [151]. The second technique to determine polyphenols is based on the collective determination. Spectrophotometric methods stood out for their simplicity. One of the most representative methods in this category is the FolinCiocalteau (FC) method [152]. It has been wellknown for decades as a reference method for the determination of polyphenols content in

food samples. FC reagent is a mixture that contains sodium molibdate, sodium tungstate, and phosphates in basic media. The reaction of this mixture and polyphenols evolves into a chromophore. Thus this compound can be measured by using spectrophotometric assays. However, this reagent can interact with other reducers in the samples, as sugars, giving false information [153]. Therefore the main drawback of this method is the lack of selectivity [154]. Researchers are making great efforts in the development of different techniques to monitor polyphenols in real samples. The ideal technique should be low cost, selective, and sensitive. In addition, for the benefits of the population, the perfect technique should be easily applied in situ. Real-time control of antioxidant capacity of food samples would have a direct and positive impact in people health. These requirements can be found in electrochemical biosensors. First, the antioxidant capacity of polyphenols is related with their capacity to donate electrons, so these substances are good candidates for electrochemical-sensing purposes. Second, biosensors contain a biological recognition element, providing selectivity, good sensitivity, and very quick responses. Third, these kinds of devices can be easily

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miniaturized and used in online analyses. They can also be considered low cost in comparison with other techniques like HPLC-MS, among others. Different biological recognition elements have been successfully employed in biosensors such as enzymes, antibodies, DNA chains, microorganisms, and others. The most-used biosensors are based on enzymatic compounds. 4.4.2.1 Enzymatic Biosensors These devices have been commented in Section 4.3. because of their utility in the plant extracts analysis. However, there is also a great number of studies related to the application of these kinds of biosensors in beverages [101,137,143,144,153,155
161]. Enzymatic biosensors owe their popularity in this sector to several factors. First, the biological recognition element, enzymes, provides selectivity toward the analyte, in this case polyphenols. This selectivity plays an important role in the analytical determinations of complex real samples such as wine, juice, or beer. Second, the enzymatic catalysis usually improves the signal provided by the biosensor. Third, the immobilization of the enzyme in the transducer can be easily carried out by using several physical and chemical methods as reported in the literature. The most common methods are physical adsorption [162], entrapment [163], covalent

binding [164], and affinity immobilization [165], among others. Most enzymes used as biological recognition elements in enzymatic biosensors belong to the family of oxidases. Their activity has been briefly commented on Section 4.3, although here we will clarify some additional issues. These enzymes are mainly based on the oxidation of several substrates. Molecular oxygen is used as electron donor and hydrogen as electron receptor. Peroxidase is also frequently applied in biosensing with the use of H2O2 as receptor, whereas another different molecule has the role of electron donor. A large number of substances can act as donor, even the target analyte. These enzymes, together with laccase and tyrosinase, constitute the biological receptor in biosensor devices for beverage assays. Laccase is an oxidase enzyme which possesses four copper atoms to perform its enzymatic reactions. This is mainly focused on the oxidation of orthodiphenols and paradiphenols to their respective quinones. In Fig. 4.4 the usual oxidation process performed by laccase is illustrated. The substrate, a p-polyphenol, is oxidized to its quinone form which can be reduced again to the original polyphenol by the application of a certain potential, resulting in an electrochemical reduction response. This signal can be correlated with the concentration of the polyphenol. Besides, many FIGURE 4.4 Scheme of the reaction catalyzed by laccase enzymes.

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4.4 BIOSENSORS FOR THE ANALYSIS OF POLYPHENOLS IN BEVERAGES

FIGURE 4.5 Tyrosinase reaction scheme. (A) monophenolase activity.

interferences are avoided since a reducing potential is used instead of an oxidative one. It is important to clarify that beverages possess a number of substances that can be easily oxidized, such as sugars among many others. On the contrary, interferents susceptible to reduction are less abundant. Tyrosinase is another oxidase frequently used in biosensors applied to beverages. This oxidase has two copper atoms to carry out its reactions. First, tyrosinase may o-hydroxilate monophenols to o-diphenols (monophenolase activity). This enzyme can also oxidize orthodiphenols to quinones (cresolase activity). A typical reaction of tyrosinase can be found in Fig. 4.5. It should be noted that o-diphenols obtained in the first step could be oxidized in the second step, as shown in Fig. 4.5. Thus the signal obtained will include the component of the native monophenols which are transformed afterwards into o-diphenols, as well as the component from the orthodiphenols contained in the sample. Therefore this enzyme has two different substrates and for this reason it is possible to assume that selectivity is lower than that obtained with laccase enzyme.

activity;

and

(B)

cresolase

FIGURE 4.6 Peroxidase enzyme reaction scheme.

Third, the basis of peroxidase enzymatic catalysis is the electron transport from several electron donors (polyphenols, in this case) to H2O2. The core of this enzyme is a heme group. These biosensors can be used to determine H2O2 as well as polyphenols. The most used is HRP. Fig. 4.6 shows a typical reaction of peroxidase enzymes. As seen in Fig. 4.6, the enzyme as electron donor has the same product as the previous enzymes, although H2O2 is required. This is the main difference between this enzyme and the others. Thus biosensors based on peroxidase will need a previous addition of H2O2 to work properly. On the other hand, the principle is similar for the other enzymes; the quinone is reduced onto the surface of the electrode under a specific potential value.

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The control of the current will serve to monitor the concentration of the polyphenol. Despite of the selectivity of the recognition element, it is well-known that the real substrate of enzymes entails a family of compounds. Therefore these biosensors are usually focused on the collective determination of polyphenols in a real sample. There are several authors who have pointed out the relationship between the antioxidant capacity and the collective polyphenol amount determined with these kinds of biosensors [138,151]. Thus they can be considered as a feasible alternative to the classical methods in order to determine antioxidant capacity in beverages. Table 4.4 lists some of the most relevant enzymatic biosensors reported in the literature over the past few years. There are several interesting aspects concerning the information collected in Table 4.4. The first thing to point out is the predominant use of laccase enzymes, although tyrosinase is also frequently used. This fact can be justified by the higher selectivity provided by these enzymes in comparison with peroxidase-based biosensors. In contrast to peroxidase, a cosubstrate like H2O2 is not necessary to perform the enzymatic activity in laccase and tyrosinase-based devices, providing simplicity. Regarding selectivity, laccase only reacts with o- and p-diphenols, considered as the most relevant polyphenols due to their great antioxidant properties. Therefore laccase has been used in very different samples such as tea, wine, beer, and others, demonstrating its high versatility and potential. Tyrosinase is also widely used as biological recognition elements, obtaining good results in spite of their lack of selectivity. In order to highlight the results obtained with the relevance of these devices, some papers are focused in the correlation of tyrosinase biosensors and antioxidant capacity determined by employing a reference method [129].

Last, very good results have been obtained by using peroxidases. However, the use of an additional compound such as hydrogen peroxide constitutes the main drawback for online studies. Regarding the analytical technique employed, the most used is amperometry, which is based on the application of a fixed potential and the recording of the current obtained. The main advantage of this technique is the enhancement of the sensitivity of the analytical measurement compared with cyclic voltammetry. In contrast, by using this technique it is not possible to discriminate the analytical signal. However, the purpose of these biosensors is not the individual determination of a specific analyte, but the collective determination of a group of substances with similar properties. Voltammetric techniques, such as DPV and CV, are more selective for analytes, but they are less reported than amperometric techniques for polyphenol determination. There is a large number of electroactive compounds in real matrices, so several peaks can be overlapped. Despite this disadvantage, very good results in terms of sensitivity and limits of detection can be found when using voltammetric techniques [144,166,169,173]. The most-used standard phenolic compounds for the reported assays are catechol, caffeic acid, and gallic acid. The selection of the standard phenolic compound can also depend on the type of sample, although a general trend has not been observed. Gallic and caffeic acid are commonly chosen as representative of polyphenols in samples of wine and beer, but catechol is usually focused on tea samples. Different polyphenols, such as catechin and hydroquinone, are gaining importance as well. Hydroquinone is an isomer of catechol and it is widely employed in electrochemical sensing. Concerning other analytes, catechin and epicatechin (catechin isomer) are polyphenols which can be found in very high

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

TABLE 4.4 Features of the Most Relevant Enzymatic Amperometric Biosensors Used for the Determination of Antioxidant Capacity in Beverages Biological Recognition Element

Transducer

Technique Analyte

Limit of Detection (M) Linear Range (M)



29





28

25

Tyrosinase

GCE PEDOT RGO Fe2O3

DPV

CAT

7 10

4 10
6.2 10

Laccase

GCE SPCE

ChA

CAT

1 1026



5 1028
6 1025

Tyrosinase/ Laccase

GCE ERGOMWCNTs/GO

ChA

CAT

5 1027/ 3 1027

1 1026
3.4 1024/ 1 1026
3 1024

Laccase

GE AgNPs/ ZnO NPs

CV

Guayacol

  5  10

Tyrosinase

CPE

ChA

HQ

1.6 1026

28






Tyrosinase

SPE CoAlSO4 LDH/ AuSPE

ChA

Mixture of polyphenols

Brasica Napus hairy roots

CPE Fc-MWCNTsMO

ChA

CA/Resveratrol 1.1 1027/ 1.1 1027

Laccase/ Tyrosinase

ITO/CS-MWCNTs

ChA

Rosmarinic acid/CA/GA





  1  10

 
5  10

25



24



2 1025
1.2 1024



26

0
2.4 10

 

3.3 1027
3.8 1024/ 2.2 1027
3.3 1024



4 1027
6.4 1026/ 4 1027
7.4 1026/ 1.6 1026/8.1 1026

2.5 1027/ 2.88 1027/ 1.55 1027



 

27

 

  

 

 







Pretreatment Employed




Extraction, filtration

Green tea [166]

13.4

Dilution, filtration

Tea, black [167] tea

310

Dilution, filtration

Fruit juice [168]

5.57

Dilution

Red wine

[169]

42.6

Dilution

Red wine

[161]

Sample

References




Green tea [160]




Wine, tea

[170]

148/152/33 Extraction, filtration, dilution

Plant extracts

[171]

Laccase

US PES

ChA

CA

8.8 10

5 10
3.5 1025

0.102

Dilution

Wine

[172]

Tyrosinase

CPE/MWCNTSs/ Naffion

ChA

Trolox










Dilution

Wine

[129]

Laccase

SPCE/Pt Nps/ RGO

CV/ChA

CA

9 1028

2 1027
2 1026

2147.38

Dilution

Tea

[158]

Laccase

CPE

DPV

GA/CAT/RUT








Dilution

Honey

[144]

Laccase

Pt-Ag

ChA

CA




Dilution

Wine

[135]

2.29/2.53/ 2.84

Extracted

Coffee

[159]

Dilution

Wine

[153]






26

Sensitivity (μA/mM)








  

27

  

26



25

Laccase

Au/Au-DTSP/Au- ChA MTPS

HQ

9.1 10 / 8.9 10
7/ 2.5 10
7

3 10
1.5 10 / 3 1026
1 1025/ 9 1027
2 1025

Laccase

SPEs-MWCNTs

GA

0.1/0.3 (ppm)

0.1
17/0.1
18 (ppm)

ChA

 

(Continued)

TABLE 4.4 (Continued) Biological Recognition Element

Transducer

Technique Analyte

Limit of Detection (M) Linear Range (M)

Sensitivity (μA/mM)

Pretreatment Employed

Sample

References

2.944 1026 (μA/ppb)

Filtered

Wine

[173]

3.64

No treatment Must and [174] wine

217

Diluted

130

No treatment Wine

[156]

Degassing, dilution

Beer

[175]

Degassing, dilution

Beer

99.45/ Degassing, 12.75/ dilution 11.00/ 89.06/28.13

Beer

Degassing, dilution

Beer

[139]

Dilution

Wine

[138]

0.009/0.007 (μA/ppm)



PPO

GCE

DPV

CATE

1.76 (ppb)

Peroxidase

SPE ChitosanMWCNTs

ChA

Gluconic acid

2.6 1026



4 1026
6.2 1024

Tyrosinase

SPE

ChA

CATE

3 1028

5 1028
1.5 1025



ITO-MWCNTs/ ChA ITO poly(GVPB)-gMWCNTSs/ITOpoly(HEMA)-gMWCNTs

CATE/GA




Tyrosinase

SNGC

CA/FA/GA/ CATE/EPI

1.43 1026/
/3.38 1026/ 3.30 1027/ 4.2 1027

Peroxidase

Laccase

SNGC

SNGC

ChA

ChA

CA/FA/GA/ CATE/EPI

CA/FA/GA/ CATE/EPI





 5  10


3.5  10

25

Tyrosinase

ChA

40
200 (ppb)

6 1028/ 1.6 1027/ 4.1 10
6/ 1 1026/ 1.6 1026

        2  10
5.15  10 / 2  10
4  10 / 2  10
8.36  10 / 2  10
2  10 / 2  10
3.2  10 4  10
2  10 / 4  10
2  10 / 1  10
2.2  10 / 4  10
3  10 / 4  10
8  10





  

24

2 1027/ 3.2 1027/ 9.4 1025/ 1.4 1027/ 1.8 1027



   

   

6 1027
2.45 1025/ 0.72/
/3.6 1026
5.8 1025/
/0.08/ 1.57/1.23 6 1027
1.03 1025/ 6 1027
2 1025 27 27

26

26

25 27

24

26

27

26

28

26

28

26


7

26

28

26

28

26

8.55/4.32/ 0.024/ 15.925/ 11.192

Laccase

SNGC

ChA

GA










Laccase

SNGC

ChA

GA

0.11 (mg/L)

50
200 (mg/L)

3.65 1027 (A L/mg)



Tea

[137]

CA, caffeic acid; CAT, catechol; CATE, catechin; ChA, chronoamperometry; CPE, carbon paste electrode; CS, chitosan; CV, cyclic voltammetry; DPV, differential pulse voltammetry; DTSP, 3,3’-Dithiodipropionic acid di(N- succinimidyl) ester; EPI, epicatechin; ERGO, electrochemical reduced graphene oxide; FA, ferulic acid; Fc, ferrocene; g, glucose; GA, gallic acid; GCE, glassy carbon electrode; GE, gold electrode; GO, oxide; GVPB, 4-vinylphenylboronate; HEMA, 2-hydroxyethylmethacrylate; HQ, hydroquinone; ITO, indium tin oxide; LDH, layer double hydroxide; MO, mineral oil; MTPS, (3-mercaptopropyl)-trimethoxysilane; MWCNTs, multiwalled carbon nanotubes; Nps, nanoparticles; PEDOT, poly(3,4ethylenedioxythiophene); PES, polyethersulfone membranes; PPO, polyphenol oxidase; RGO, reduced graphene oxide; RUT, rutin; SCE, screen printed electrode; SCPE, screen printed carbon electrode; SNGC, Sonogel-carbon.; US, universal sensor.

4.4 BIOSENSORS FOR THE ANALYSIS OF POLYPHENOLS IN BEVERAGES

concentration in samples such as wine and beer. Last, other polyphenols such as resveratrol can be used as well. The importance of resveratrol in human health has been previously demonstrated [149]. As reported in Table 4.4, wines, beers and teas are the most commonly analyzed samples. This trend is similar to the one found in the consumption graph, (Fig. 4.1). Beer and tea are highly consumed due to cultural factors or the health benefits attributed to these beverages, as widely exposed until now. Wine is also highly consumed, but not as much as beer and tea. However, there has been a great amount of research regarding this sample. The lower worldwide consumption level cannot be correlated with a smaller economic impact. In 2018 the International Organisation of Vine and Wine reported that 2017 wine global market generated benefits of about 34.2 billions of US dollars, which can be considered a very significant amount since it supposes a rise of 4.8% compared with 2016. In terms of antioxidant capacity, wines have the highest index. Therefore higher benefits of health could be expected under responsible and advised consumption of wines by health authorities. The usual pretreatments applied to the real sample are dilution, filtration, extraction, and degassing. Extraction is only required in tea samples due to the different solubility of tea compounds. Dilution is almost always needed to improve the sensitivity of the biosensors. Filtration is not common, but in some cases with juice or tea it can be performed to avoid some suspended material. Degassing is required in beer in order to remove the CO2, which hinders the correct application of the electrochemical procedure. It should be noted that the best biosensor would involve the least treatment possible to easily perform in situ and online analyses. The large number of approaches dealing with the transducer can be also noticed. Transducers convert chemical information into an electrical signal. Ceramic, carbon, and

131

screen-printed electrodes are widely used for these purposes, although indium and tin oxides popularity has risen over the past few years. All these transducers are commercially available. Other handmade electrodes can be excellent candidates as immobilization matrices. In this way, carbon-based ceramic electrodes obtained by using sol
gel process have emerged due to their excellent conductivity as well as a high surface area. A new type of ceramic electrode, namely, sonogel-carbon, emerged in the past decade as an excellent alternative for biosensor transducers. Several characteristics such as low residual current, renewable surface, and wide operational range can be attributed to this material [176]. All the materials mentioned above can be easily modified to improve their features, leading to an enhancement of electrical conductivity. However, higher conductivity is not the only aim for the modification of these materials. For instance, it has also demonstrated that metal nanoparticles, such as AuNps, improve the direct electron transfer between the biological recognition element and the transducer, increasing the sensitivity of the measurements [177]. On the other hand, conducting polymers can be used as the supporting element to improve the immobilization of the enzyme onto the electrode surface [116]. This immobilization process enhances the robustness of the biosensor obtained. Metal nanoparticles, carbon nanomaterials, and conducting polymers are frequently reported as modifiers. The modified transducers have better analytical performance such as limit of detection, limit of quantification, and sensitivity. In Table 4.4 very good sensitivities can be appreciated [158,168,175]. It is noteworthy to mention that correlation studies with antioxidant capacity methods are much more useful than having an excellent limit of detection in these applications. The procedure used should be as selective as possible. FC is not a good example for reasons previously stated. However, methods based in radical scavenging offer a more

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4. ELECTROCHEMICAL BIOSENSORS FOR ANTIOXIDANTS

suitable approach. In this line of reasoning, some assays have obtained very good relations [129,135,139,157] using methods such as DPPH, ABTS, and others. However, some authors keep using FC for the correlation in spite of the lack of selectivity this method shows. It is also possible to use several enzymatic biosensors simultaneously for sample analysis. According to this idea electronic tongues stood out. In general, an electronic tongue is a battery of amperometric sensors (biosensors, sensors, or a mixture) which are applied simultaneously to one sample. These devices provide a large number of different responses that require statistical treatment to extract useful conclusions [134,178]. The main disadvantage of electronic tongues is the timeconsuming treatment of the data, although it leads to a more complete characterization of the sample than that obtained with single biosensors. 4.4.2.2 DNA Biosensors In the previous sections, the relation between antioxidant compounds with biological harm has been mentioned. Biological damage is mainly caused due to ROS, obtained as by-products in breathing procedures. This group contains peroxides, superoxides, and hydroxyl radicals, among others, which affect tissues and even DNA and can lead to several diseases. Using this approach DNA biosensors are considered as an excellent alternative to determine the antioxidant capacity in beverages since they mimic the process and interactions in the human body under oxidative stress. DNA can be immobilized onto the surface of the transducer, using the genetic material as biological receptor. The nucleobases (adenine or guanine) are damaged by the presence of ROS, resulting in a decay of the electroanalytical signal. The addition of a polyphenolic sample removes ROS species, improving the response obtained. Consequently, this improvement can

FIGURE 4.7 Scheme of Fenton reaction in the human body.

be related to the antioxidant capacity of the sample [97]. Fig. 4.7 The most accepted reaction to study the degradation of DNA is the Fenton reaction, represented in Eq. (4.4). Fe21 1 H2 O2 -Fe31 1 OH2 1 OH 

(4.4)

Fe (II) is used to promote the generation of hydroxyl species. Although it has been demonstrated that Cu (II) and also Cr (II) can also perform this generation, iron is the most considered due to its abundance in living beings and having the highest ratio of reactions. On the other hand, hydroxyl radical participates in reactions with other species as it can be noticed in Eqs. (4.5) and (4.6). OH  1 H2 O2 -H2 O 1 OH2

(4.5)

HO2 1 Fe31 -Fe21 1 O2 1 H1

(4.6)

In real conditions, Fenton reactions are not simple because of the existence of several cofactors and enzymes. For instance, the

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

4.4 BIOSENSORS FOR THE ANALYSIS OF POLYPHENOLS IN BEVERAGES

concentration of H2O2 is controlled by superoxide enzyme (SOD) and catalase. A scheme of a possible Fenton reaction adapted for living beings is presented in Fig. 4.5. Several oxidants and reducers contained in the environment participate in the formation of DNA  . The molecule resulting can be repaired in some cases, although it is possible to have hard damage in its structure. This damage can be due to the modification of more than 20 different nucleotides. Among these, 8-oxoguanine is the product mostly studied and it is used as a clinical biomarker for oxidative damage to DNA. A typical oxidation of guanine base to this biomarker is shown in Fig. 4.8 Within the biosensor, different strategies to detect DNA damage can be applied. The most common ones are based on the interaction between DNA and the transducer. • Studying the oxidation signal of DNA and its modifications. • Application of electrochemically active mediators. • Measuring possible changes in the properties of charge transfer of the DNA layer deposited onto the surface of the electrode. DNA biosensors in antioxidant fields are based on the abovementioned strategies. Their electrochemical signal can be excellently correlated with other reference analytical methods such as DPPH, ABTS, Trolox, and others. The antioxidants selected are added to the solution

133

at the same time as the Fenton reagent. The signal is compared with the signal in the absence of antioxidants. It is important to point out that damage in DNA is mostly irreversible, so these biosensors are, in general, for single use. These biosensors involve complex procedures to ensure their robustness, so they are not suitable for field work. On the other hand, good reproducibility can be achieved [142]. The most relevant DNA biosensors currently used are listed in Table 4.5. A comparison of the data displayed in Table 4.5 and Table 4.4 shows that DNA biosensors are less frequently applied to beverages than the enzymatic biosensors. This is due to factors previously mentioned such as less robustness or single use. However, there are several biosensors which provide very good results. In the first column the biological recognition elements are shown. Two trends are clearly identified. The first uses directly the nucleotide base, such as guanine, adenine, and others. The second trend is the use of some biological sample containing this nucleotide, for instance, salmon or herring sperm. In the latter, the DNA is directly dissolved in a buffer solution and used afterwards. It seems that the best results are obtained for the biosensors using the purified nucleotide base. One feasible reason to use another genetic compound such as herring and salmon sperm is their low cost. In contrast with enzymatic biosensors, transducers are not so much modified in this

FIGURE 4.8 Guanine oxidation to 8-oxoguanine.

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

TABLE 4.5 Features of the Most Relevant DNA-Based Biosensors Used for the Determination of Antioxidant Capacity in Beverages Biological Recognition Element

Transducer

Technique Employed

Guanine

GCE

SWV

AA/GA/ 0.29/0.09/0.06/ CA/COU/ 0.08/0.07 Resveratrol

0.5
2.5/0.1
0.5/ 0.4
0.8/0.31
0.73/ 0.10
0.5

Adenine

GCE

SWV

Deoxyadenylic acid

CPE

Thymus DNA

Sensitivity (μA/ppm)

Pretreatment Employed

2.82/9.33/ 8.76/9.20/ 11.8




AA/GA/ 0.99/0.08/0.07/ CA/COU/ 0.27/0.10 Resveratrol

2
6/0.11
0.44/ 0.4/7.38/ 0.1
0.5/0.1
1/0.1
0.5 11.9/3.81/ 8.78




SWV

AA/GA

2.7 (AA)

10
30

CPE

DPV

GA/CA




Salmon sperm DNA

GCE/Polyvynil alcohol

CV

-

Salmon sperm DNA

SPCE Chitosan/Nafion CV, EIS




Deoxyadenylic acid

CPE

SWV

AA/GA/ 0.23 CA/p-COU

Herring sperm DNA

GCE/ Poly L-glutamic acid doped silver hybridized membrane

CV

AA




Guanine

GCE/ Graphene nanoribbon

SWV

AA

0.05

Analyte

Limit of Detection (ppm) Linear Range (ppm)

Sample

Reference

Flavored water

[179]

Filtered and diluted

Tea

[97]

167.53/ 2 1028
5 1028/ 7.5 1028
7.5 1027(M) 95.90 (μA/ mM)

Dried and extracted

Black [180] and green tea













White wine

[181]













Black tea, coffee

[182]

1
20

3.04

Filtered

Juice

[183]

1 1026
5 1025(M)




Filtered

Juice

[184]

0.1
4

4.16

Centrifuged and filtered

Juice

[185]







7.7 (AA)







AA, ascorbic acid; CA, caffeic acid; COU, coumaric acid; CPE, carbon paste electrode; CV, cyclic voltammetry; DPV, differential pulse voltammetry; EIS, electrochemical impedance spectroscopy; GA, gallic acid; GCE, Glassy carbon electrode; SPCE, screen printed carbon electrode; SWV, square wave voltammetry.

135

4.5 CONCLUSION

case. One possible reason is the difficulty to immobilize DNA onto the surface of the electrode. Nevertheless, some good results have been obtained using polymers and nanomaterials, such as polyvinyl alcohol [181] and silver nanoparticles [184]. Several membranes have been used in order to protect and provide robustness to these kinds of biosensors (chitosan, polyvinyl alcohol, etc.). Otherwise, metallic nanoparticles are used for the same reasons as for enzymatic biosensors, that is, enhanced conductivity and electrocatalytic effect. The lack of transducer modification reflects the difficulties to work with DNA instead of enzymes. In addition, the relevance of these kinds of biosensors relies on their ability to measure antioxidant capacity in an efficient way, very similar to real body conditions, hence higher sensitivities or limits of detection are not required. Voltammetric techniques such as SWV and CV are frequently used due to their ability to monitorize the DNA harming. Anodic SWV is applied in the detection of several compounds related to DNA damage. For instance, the analysis of thioridazine, an intercalator of these oxidative procedures, has been reported in the literature and as shown in Table 4.5 [182,186]. In addition, destruction of guanine and adenine nucleotide can be monitorized by using a similar procedure. Other compounds can be tested as electrochemical probes in order to evaluate the deterioration of DNA chain. Another approach based on the drop of the electrochemical response is the employment of EIS. Nyquist plots provide information about the resistance of the electrochemical system, which is related to the degradation phenomenon [187]. Regarding analytes used for electrochemical sensing, several similarities with Table 4.3 can be appreciated. The main compounds used are GA, CA, and resveratrol, among others. Ascorbic acid is also frequently used, since it is one of the most important benchmark analytes

reported in the literature. Beer and wine are much more commonly studied using enzymatic biosensors than DNA devices, since the application of these biosensors involves more simple matrices like teas and juices. Other complex samples contain more interfering species and protective membranes are necessary to avoid fouling processes. However, many difficulties arise due to the presence of these membranes. This is why, as shown in Table 4.4, pretreatments such as filtration, centrifugation, and dilution are frequently applied as well. To summarize this section, polyphenols have a critical role in the antioxidant capacity of highly consumed beverages. Consequently, there is a wide range of techniques to determine them individually and collectively. In this final Section, biosensors have been thoroughly discussed according to several advantages such as quick response, sensitivity, simplicity, robustness, limit of detection, low cost, and the possibility of in situ analyses. Two different approaches can be observed, using enzymes as biological recognition elements or the use of a DNA compound. In most cases enzymatic biosensors are more frequently used than the others. However DNA biosensors show a response that can be directly related with the antioxidant capacity. Hence the development of DNA-based biosensors is a very interesting and growing field that can expect a great raise in next years.

4.5 CONCLUSION As highlighted throughout this chapter, antioxidants and healthcare are intimately related to each other, since the former help us to prevent significant diseases affecting the population all around the world. This is why it is essential to possess a complete set of fast, simple, and versatile tools that are useful for the determination of these antioxidant

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

136

4. ELECTROCHEMICAL BIOSENSORS FOR ANTIOXIDANTS

substances. Due to the advantages of electrochemical methods based on the use of biosensors, in terms of simplicity, rapidity, sample treatment, sensitivity, and limits of detection and quantification compared to classical analytical methodology, electrochemical biosensors are potentially improved alternatives for that purpose. In this chapter, the use of electrochemical biosensors in the assessment of various antioxidants in plants, food, and beverages was reported, paying special attention to the major achievements over the past decade. The preparation of oxidase enzymes-based electrochemical biosensors by various immobilization procedures has been thoroughly discussed from the point of view of the obtained analytical performances. The large number of selected applications commented on within this chapter clearly demonstrates the feasibility and utility of the electrochemical biosensors for the successful electroanalyses of antioxidants in real samples having complex matrix compositions. On the other hand, these results show that great attention has been devoted to the improvement of the analytical performance of biosensors in terms of limit of detection, linear response range, sensitivity, and stability. It was demonstrated that the enzyme immobilization procedure as well as the immobilization matrix play major roles in the improvement of the overall analytical performances. The reported limit of detection, linear response range, and sensitivity values underline the competitiveness of the electrochemical biosensors among other analytical methodologies. New directions in the development of biosensors for antioxidant determination may be related to the use of multienzymatic systems, more stable and size-reduced immobilization platforms, application of chemometrics tools in experimental data treatment, and the development of disposable biosensors.

4.6 Acknowledgments S. Lupu gratefully acknowledges the grant from the Romanian National Authority for Scientific Research, CNCS
UEFISCDI, Project number PN-II-ID-PCE-2011-30271 (298/06.10.2011). J. J. Garcı´a-Guzma´n and D. Lo´pezIglesias acknowledge ESF funds, Sistema de Garantı´a Juvenil depending on Ministerio de Empleo y Seguridad Social of Spain and Junta de Andalucı´a for their employment contracts. The Spanish research group thanks Junta de Andalucı´a and the Institute of Research on Electron Microscopy and Materials (IMEYMAT) for their financial support.

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C H A P T E R

5 Electrochemical Immunosensors for Rapid Detection of Breast Cancer Biomarkers Ankita Sinha1, Dhanjai2,3,*, Samuel M. Mugo4, Huimin Zhao1, Jiping Chen3 and Rajeev Jain5 1

Key Laboratory of Industrial Ecology and Environmental Engineering (Ministry of Education, China), School of Environmental Science and Technology, Dalian University of Technology, Dalian, P.R. China, 2Department of Mathematical and Physical Sciences, Concordia University of Edmonton, Edmonton, AB, Canada, 3CAS Key Laboratory of Separation Science for Analytical Chemistry, Dalian Institute of Chemical Physics, Chinese Academy of Sciences, Dalian, P.R. China, 4Department of Physical Sciences, MacEwan University, Edmonton, AB, Canada, 5School of Studies in Chemistry, Jiwaji University, Gwalior, India

5.1 INTRODUCTION Cancer is a life-threatening disease that affects people worldwide. It is a multistage disease which involves complex genomic or epigenomic alterations progressively through various stages. These changes disturb the cellular signaling of the body system and lead to tumorigenic transformation and malignancy. Currently more than 200 types of cancers have been identified and the disease is considered as the leading cause of deaths all over the world [1].

Breast cancer is a very common type of cancer occurring in women and contributes to about 23% of all cancer cases worldwide. It is a serious health concern for women as it accounts for the second largest number of deaths among all cancers [2]. Breast cancer is frequently diagnosed among women in 140 of 184 countries worldwide and represents one in four of all cancers in women. Since 2008, worldwide breast cancer incidence has increased by more than 20% and mortality has increased by 14% (www.bcrf.org). As per the World Health Organization (WHO),

*

Corresponding authors

Advanced Biosensors for Health Care Applications DOI: https://doi.org/10.1016/B978-0-12-815743-5.00005-6

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© 2019 Elsevier Inc. All rights reserved.

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5. ELECTROCHEMICAL IMMUNOSENSORS FOR RAPID DETECTION OF BREAST CANCER BIOMARKERS

approximately 508,000 women died in 2011 due to breast cancer (Global Health Estimates, WHO 2013). Further, according to the report, almost 50% of breast cancer cases and 58% of deaths occur in less-developed countries (GLOBOCAN 2008). The incidence rates vary worldwide from 19.3 per 100,000 women in Eastern Africa to 89.7 per 100,000 women in Western Europe (www.who.int). According to the Breast Cancer Research Foundation, about 1.7 million new breast cancer cases were diagnosed in 2012. The American Cancer Society reported approximately 231,840 new cases of invasive breast cancer and 40,290 breast cancer deaths among US women in 2016 (American Cancer Society Breast Cancer Facts & Figures 2016). The new breast cancer cases increased to 252,710 and deaths increased to 40,610 in 2017 (American Cancer Society Breast Cancer Statistics, 2017). In 2018 the number of new cases of breast cancer further increased to 266,120 and deaths increased to 40,920 (American Cancer Society Cancer Statistics, 2018). As per the report, from 2005 to 2014, the overall breast cancer cases increased among Asian/Pacific Islanders (1.7% per year), non-Hispanic blacks (0.4% per year), and Hispanic (0.3% per year) women. However, it was observed stable in non-Hispanic white and American Indian/ Alaska Native (AI/AN) women. The early stages of cancer development pose maximum potential for therapeutic intervention and about 90% of deaths related to cancer occur from metastasis of the primary cancer tumor [3]. Therefore detecting premalignant or premetastatic malignant tumors is crucial to achieve effective treatment of cancer [4]. The molecules which undergo major alterations during various stages of cancer are recognized as biomarkers and are classified into three categories, namely, diagnostic, prognostic, and predictive. According to the National Cancer Institute, a biomarker is defined as “a biological molecule found in

blood, other body fluids, or tissues that is a sign of a normal or abnormal process or of a condition or disease” [5]. Diagnostic biomarkers are related to the detection of the disease while prognostic biomarkers provide information about the course of reoccurrence of the disease. Further it can be used to indicate the aggressiveness of the tumor. Predictive biomarkers measure the response to the undergoing treatment [1,2]. Various progressive phases of cancer development are marked by the changes in the level of the specific biomarkers in the cell. Further, these tumor markers are considered as one of the most valuable tools for early cancer detection, classification, staging, and progression monitoring. Therefore identification of reliable biomarkers can be highly useful for detection, prognosis, and treatments of cancer. Development of costeffective, reliable, and powerful screening strategies for cancer biomarkers is crucial due to the prevalence of the disease and high possibility of its reoccurrence. Further, simultaneous measurements of biomarkers are required to avoid false positives which can be raised from population variations in expression of a single biomarker during diagnosis. For cancer diagnosis, various techniques such as mammography, biopsies, sonography, magnetic resonance imaging, molecular breast imaging, and thermography have been applied. Apart from these, the traditional biomarker-based expression techniques such as enzyme-linked immunosorbent assay (ELISA), polymerase chain reaction, radioimmunoassay, and immunohistochemistry methods have also been utilized for cancer biomarker detection [2]. However, these techniques often suffer from technological limitations such as long-reaction time, high cost, and consumption of expensive reagents. Further, the advanced liquid chromatography
mass spectrometry based proteomics are too expensive and complicated to detect cancer biomarkers in routine clinical diagnostics. Therefore to detect biomarkers, the focus

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

5.1 INTRODUCTION

has been inclined toward simple analytical techniques which are capable of sensitive and parallel detection of biomarkers rendering high sensitivity, selectivity, and point-of-care testing. Recently, various modern biotransducer systems have been developed for breast cancer biomarker detection. For example, electrochemical transducers are one of the most common biotransducer systems which are based on the biochemical reactions between the biomarker and electrode surface. Additionally, light absorption and emission-based optical biotransducers, fluorescence resonance energy transfer (FRET) biotransducers, quantum dots
based FRET, electroluminescence (ECL) and photoluminescence biotransducers, surface plasmon resonance biotransducers, colorimetric biotransducers, and mass change
based biotransducers are other methods used [2]. In this chapter, we will only focus on electrochemical affinity
based biotransducers (immunosensors). Electrochemical immunosensors are powerful analytical tools which detect a highly specific binding event between antibody (Ab) and antigen (Ag) by the formation of a stable complex. These sensors collectively combine the inherent specificity of immunoreactions and simplicity of commercially available physical transducers [6]. These sensors are mainly comprised of two main classes, namely, potentiometric and amperometric immunosensors, where potentiometric methods include various forms of voltammetry such as linear sweep (LSV), differential pulse (DPV), square wave (SWV), and stripping. The immunosensors work either as direct (unlabeled) or indirect (labeled) sensors. Direct sensors are able to detect the physical changes during the immune complex formation, whereas indirect sensors employ signal generating labels which are more sensitive when incorporated into the complex. Recently, the impedimetric and capacitive immunosensors are also gaining immense attention due to their direct mode to analyze

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Ab
Ag interaction. However, they are less sensitive for clinical applications [7,8]. The indirect mode of electrochemical immunosensors runs by labeling a tracer antibody (Ab2) with an electroactive species (enzyme/nanoparticles/quantum dots) and then allowing the tracer to bind with the target analyte, which can be through a primary antibody or capture antibody (Ab1). This complex is immobilized on an electrode surface and the concentration of the target biomarker is quantified by applying a potential and measuring the resulting current. Basically, the electrochemical immunosensing quantitatively records the electrical signal resulting from the binding activity of the Ab2 and target analyte (Ab1 captured biomarker). This mechanism is clearly illustrated in Fig. 5.1A [9]. Further, the applied potential drives the redox reaction of labeled electroactive species and the resulting current is proportional to the concentration of the Ab-bound analytes [7,8]. Electrochemical immunosensors offer remarkable performances for detection of cancer biomarkers. However, this chapter focuses on the detection of only breast cancer biomarkers. Advanced state-of-the-art electrochemical immunosensing strategies developed for clinically relevant breast cancer biomarkers have been well discussed and contribute significantly to their biomedical and therapeutic monitoring.

5.1.1 Breast Cancer Biomarkers Certain specific biomolecules such as cell surface protein receptors, mutated genes, or miRNAs are considered as clinically relevant biomarkers whose varied expression in the tumor cell indicate the stages of cancer progression. Like other cancer biomarkers, breast cancer biomarkers are also categorized into major classes of stage-dependent biomarkers and overexpressed biomolecule-based

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5. ELECTROCHEMICAL IMMUNOSENSORS FOR RAPID DETECTION OF BREAST CANCER BIOMARKERS

(A)

Enzymatic substrate Enzyme

Detection antibody Electroactive product

Analyte Capture antibody

eElectrode Transducing

Response

Signal Detector

(B) Prognostic biomarkers (ER, PR, HER-2, BRCA1, Ki67, osteopontin, mammoglobin, sirtuins, autoantibodies) Cancer stage based biomarkers

Therapeutic biomarkers (ER, PR, HER-2, CA15-3, CA27.29, Ki67, miR-21, CTC) Diagnostic biomarkers (HER-2, CEA, BRCA1, miR-21, 155, 222)

Breast cancer biomarkers

Glycoproteins MUC1, HER-2, CEA, EpCAM, EGFR

Biomolecule based biomarkers

DNA BRCA1, BRCA2

Micro RNA miR-21, 16, 27a, 150, 191

FIGURE 5.1 (A) Working principle of an electrochemical immunosensor [9]. (B) Classification of breast cancer biomarkers [2]. Reprinted with permission from I.H. Cho, J. Lee, J. Kim, M. Kang, J.K. Paik, S. Ku, et al., Current technologies of electrochemical immunosensors: perspective on signal amplification, Sensors (2018) doi:10.3390/s18010207.

biomarkers [2]. Based on this, a classification scheme of breast cancer biomarkers is shown in Fig. 5.1B. Cancer stage
based diagnostic biomarkers are regarded as of higher significance than those of prognostic and therapeutic

biomarkers. However, some biomarkers such as human epidermal growth factor receptor 2 (HER-2), estrogen, and progesterone are recognized as both diagnostic and prognostic kinds of biomarkers. Furthermore biomolecule-based

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5.2 ELECTROCHEMICAL IMMUNOSENSING OF BREAST CANCER PROTEIN BIOMARKERS

biomarkers include various protein and nucleic acid (genomic DNA, epigenomic DNA, mRNAs, and miRNAs) biomarkers. However, we will only consider protein-based breast cancer biomarkers in this chapter. Altogether, every biomarker is very specific in its function and information it reveals at each progressive stage. Detection of breast cancer protein biomarkers by electrochemical immunosensing technique has been widely explored over the past few years. The potential of biomarkers as target molecules through biosensing techniques contributes to multiple sets of information for breast cancer diagnosis, prognosis, and treatment. Further, the studies establish novel stateof-the-art detection strategies toward breast cancer biomarkers for clinical applications.

5.1.2 Signal Amplification Strategies in Electrochemical Immunosensors Electrochemical immunosensing employs nanostructured surfaces, nanoparticle labels, enzyme labels, and magnetic beads which offer highly sensitive biomarker detection through the sandwich immunoassay approach [7]. As per the sandwich immunoassay scheme, the transducer surface is functionalized with capture antibodies (Ab1) to bind with the biomarker targets. Further a tracer enzyme/nanoparticle-labeled antibody (Ab2) is added which is complexed with captured analyte protein. The labeled antibodies are used as signal amplifiers which enhance the sensitivities of the detection processes. For example, use of magnetic beads in immunoassays as substrates for Ab1 or for target Ags facilitates the reaction kinetics significantly compared to bulk solid transducer surfaces. Another possible way to achieve amplified signals is accomplished by loading a large number of biomolecules or multienzymes onto nanoparticle probes (e.g., carbon nanotubes (CNTs), polymers, gold nanoparticles (AuNPs), and

151

magnetic beads). A systematic illustration of magnetic beads and various labeled bioconjugates based immunosensing strategies are shown in Fig. 5.2A and B, respectively [7].

5.2 ELECTROCHEMICAL IMMUNOSENSING OF BREAST CANCER PROTEIN BIOMARKERS In normal cells, the tumor markers exist at very low concentrations, however, this level changes upon tumor formation. Therefore highly sensitive immunoassays have been developed that are capable to detect even small changes in the levels of biomarkers. This section accounts for different electrochemical immunoassays toward the detection of protein-based breast cancer biomarkers and a summary is provided in Table 5.1.

5.2.1 Cancer Antigen Out of the variety of cancer-related Ags (CA) that are being explored for clinical diagnoses of cancer cells, CA15-3, CA27.29, and CA549 are the major biomarkers which show their elevated levels in breast cancer tumor cells as .25 kU/L, .38 kU/L, and .10 kU/L, respectively. CA15-3 is a breast-associated mucin whereas CA27.29 and CA549 are soluble forms of MUC1 glycoprotein and highmolecular weight glycoprotein, respectively. Among these, CA15-3 is the most widely studied biomarker (elevated in breast carcinoma with distant metastasis) through electrochemical immunosensing [10
22,63,64]. For example, label-free immunosensing methodologies based on an N-doped graphene (N-GR) biosensor were developed [10,11]. The anti-CA15-3 Ab was attached to the N-GR surface through amidation reaction between the carboxylic group of N-GR and the amino group of Ab. These immunosensors, by incorporating a highly

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5. ELECTROCHEMICAL IMMUNOSENSORS FOR RAPID DETECTION OF BREAST CANCER BIOMARKERS

(A)

Capture antibody to protein of interest Surface functionalized magnetic bead

Electrochemical measurement

(B)

Tracer antibody labeled with enzyme

Interaction with protein analyte

Enzyme substrate

Signal

Carbon WE Magnetic bar

a

b

c

Protein Ab1 Labeled secondary antibody d

e

f

g

h

i

Sensor surface Blocking agent

FIGURE 5.2 (A) Scheme for sandwich immunoassay using magnetic beads. (B) Sandwich immunoassay-based signal amplification strategies using bioconjugates; (a) Ab2-enzyme, (b) Ab2-nanoparticle, (c) Ab2-biotin-strepatavidin-enzyme, (d) Ab2-CNTs-enzyme, (e) CNT-(PDDA-AP)4-PDDA-PSS tag, (f) multienzyme-Ab2-nanoparticle, (g) Ab2-nanoparticleQdots, (h) Ab2-MB-multienzyme clusters, and (i) MB-AuNPs-Ab2-multienzyme [7]. Reprinted with permission from B.V. Chikkaveeraiah, A.A. Bhirde, N.Y. Morgan, H.S. Eden, X. Chen, Electrochemical immunosensors for detection of cancer protein biomarkers, ACS Nano, 6 (2012) 6546
6561.

conductive N-GR modified electrode, exhibited significantly increased electron transfer and high sensitivities toward CA15-3. Further, a polyglutamic acid/carbon nanotubes (PGA/ CNTs/anti-CA15-3/Ab) based biosensor was prepared for label-free detection of CA15-3 biomarker. The current change in response to the immunological reactions and the varying concentrations of CA15-3 was recorded as the analytical signals [12]. Furthermore, recently a copper sulfide and reduced graphene (CuS-rGO) nanocomposite-based, labelfree immunosensing was reported toward detection of CA15-3 biomarker [13]. The anti-CA15-3 Ab was immobilized on CuS-rGO modified screen printed graphite electrode

(SPGE) and electrochemical measurements were carried out in 1.0 mM catechol (in 0.1 M phosphate buffer). Addition of CA15-3 Ag resulted in its complexation with Ab and also reduced the catalytic activity of CuS-rGO due to the oxidation of catechol. The current response decreased with the increase of CA15-3 Ag concentration and a detection limit of 0.3 U/mL was obtained. These label-free immunosensors were capable of recording the direct response of the change in electrochemical signals during immunoreactions at the electrode surface which delivered real-time and in situ detection benefits. In current time, sandwich-type immunoassays have been more frequently applied for biomarker detection.

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

TABLE 5.1 Electrochemical Immunosensing of Breast Cancer Biomarkers Biomarker

Method

Biosensor

Label

LOD

Linear range

Refs.

CA15-3

DPV

N-GR/GCE

Label-free

0.012 U/mL

0.1
20.0 U/mL

[10]

CA15-3

DPV

N-GR/GCE

Label-free

0.017 U/mL

0.1
20.0 U/mL

[11]

CA15-3

DPV

PGA/CNTs/CPE

Label-free

0.025 U/mL

0.1
30.0 U/mL

[12]

CA15-3

DPV

CuS/rGO/SPGE

Label-free

0.3 U/mL

1.0
150.0 U/mL

[13]

CA15-3

DPV

PPy/ P-1,5DAN/SPE

MB/HRP

0.02 U/mL

0.05
20.0 U/mL

[14]

CA15-3

Chronocoulometry GO/ITO

Gal/Tyr

0.1 kU/L

0.1
100.0 U/mL

[15]

CA15-3

Amperometry

MP/HRP

6.0 μU/mL

10.0
1000.0 μU/mL

[16]

DμID

25

CA15-3

DPV

AuNPs-GR/GCE

HRP/liposomes

5.0 μU/mL

2.0 3 10

CA15-3

DPV

GO/Au

MWCNTs/ferritin

0.01 U/mL

0.05
100.0 U/mL

[18]

CA15-3

DPV

AuNPs/Fc-ERGO/GCE

Label-free

0.015 U/mL

0.05
2.0 U/mL

[19]

CA15-3

ECL

AuNPs/GR/GCE

PdNCs/PEI/PSRu

0.003 U/mL

0.01
120.0 U/mL

[20]

CA15-3

DPV

GO/SPCE

Magnetic SiNPs/GO 21

24

2.8 3 10

U/mL

23

10

to 4.0 U/mL

to 200.0 U/mL

[17]

[21]

CA15-3

SWV

IL/GR/GCE

f-TiO2/Cd

0.008 U/mL

0.02
60.0 U/mL

[22]

EGFR

EIS

AuNPs-protein G/Au

Label-free

0.34 pg/mL

1.0 pg/L to 1.0 μg/mL

[23]

EGFR

DPV

Apt/MB

AuNPs

50.0 pg/mL

1.0
40.0 ng/mL

[24]

HER-2

EIS

AuNPs/SPCE

Label-free

0.01 ng/mL

0.01
100.0 ng/mL

[25]

HER-2

EIS

AuNPs/MWCNTs/CILE

Label-free

7.4 ng/mL

10.0
110.0 ng/mL

[26]

HER-2

EIS

AuNPs/SPGE

Label-free

6.0 ng/mL

0
40.0 ng/mL

[27]

HER-2

LSV

AuNPs/SPCE

STP/ALP

4.4 ng/mL

15.0
100.0 ng/mL

[28]

HER-2

DPV

Fe3O4NPs/GCE

AuNPs/Fe3O4NPs

2.0 3 1025 ng/mL

5.0 3 1024 to 50.0 ng/mL [29]

HER-2

DPV

Fe3O4NPs/Au

Label-free

0.99 ng/mL

0.01
100 ng/mL

[30]

HER-2

Amperometry

Inkjet-printed Au

STP/HRP

12.0 ng/mL




[31]

HER-2

Amperometry

SPCE

HRP

1.0 μg/mL

1.0
200.0 μg/mL

[32]

HER-2

DPV

SPCE

ALP

6.0 ng/mL

0
15.0 ng/mL

[33] (Continued)

TABLE 5.1 (Continued) Biomarker

Method

Biosensor

Label

LOD

Linear range

Refs.

HER-2

SWV

AuNPs/GCE

Hydrazine/AuNP/ apt

0.04 pg/mL

0.0001
10.0 ng/mL

[34]

HER-2, CA15-3

LSV

Bi-AuNPs/SPCE

ALP

2.9 ng/mL,5.0 U/ mL

0
50.0 ng/ mL,0
70.0 U/mL

[35]

HER-3

EIS

Au

Label-free

0.28 pg/mL

0.4
2.4 pg/mL

[36]

VEGFR-2

DPV

Chi/rGO/Th/GCE

HRP

0.28 pM

0.4
86.0 pM

[37]

CEA

DPV

PTGR/GCE

HRP/PtNWs

5.0 pg/mL

0.01
60.0 ng/mL

[38]

CEA

DPV

ERGO/GCE

AuNPs/Cu-MOF

0.03 ng/mL

0.1
80 ng/mL

[39]

CEA

DPV

PD/AuNPs/GCE

TMB

10.0 pg/mL

0.05
10.0 ng/mL

[40]

CEA

DPV

AgNPs/PPD/MB/MPCE

AuNPs/GO/HRP

1.0 pg/mL

0.01
40.0 ng/mL

[41]

CEA

Amperometry

AgNPs/SnO2/rGO/GCE

PdNPs/V2O5/ MWCNTs

0.17 pg/mL

0.0005
25.0 ng/mL

[42]

CEA

DPV

Fe3O4NPs/GR

HRP/AuHS

1.0 pg/mL

0.01
80.0 ng/mL

[43]

CEA

SWV

Th/AuNPs/Den/Au

MWCNTs/GOx/ HRP

4.4 pg/mL

0.01
50.0 ng/mL

[44]

CEA

DPV

AuNPs/PB/PEDOT/GCE

Label-free

0.01 ng/mL

0.05
40.0 ng/mL

[45]

CEA

CV

GR/Chi/fc/GCE

Fe3O4NPs/AuNPs

0.4 pg/mL

0.001
30 ng/mL

[46]

CEA

FET-SWV

HRP/anti-CEA/AuNPs/ ZnONPs/Au




0.01 ng/mL

0.1
200.0 ng/mL

[47]

BRCA-1

Amperometry

Th/PVPP/GR/GCE

SBA-15/HRP

4.9 pg/mL

0.01
15.0 ng/L

[48]

CD-146

Amperometry

rGO/TEPA

Au/PdNPs/TiO2

1.6 pg/mL

0.005
20.0 ng/mL

[49]

CD-105

CV

AuNPs/Au

PtNPs

0.9 μg/mL

1.3
200.0 μg/mL

[50]

IL-6

DPV

Magnetic bead

AgNP/TiP

0.1 pg/mL

0.0005
10.0 ng/mL

[51]

IL-6

Amperometry

Magnetic bead

PolyHRP/STP

0.4 pg/mL

1.75
500.0 pg/mL

[52]

IL-6

Amperometry

Au compact disk

STP/HRP

10.0 fg/mL

10.0
1300.0 fg/mL

[53]

IL-6

SWV

GR

PDDA/BN/AuNCs

1.3 pg/mL

0.005
100.0 ng/mL

[54]

IL-8

Amperometry

SPCE

MWCNTs/HRP

8.0 pg/mL

8.0
1000.0 pg/mL

[55]

TN-C 392

p53

EIS

ITO

HRP

SWV

AuNPs/SPCE

HRP/GO

7.0 μg/mL

0.2
12.5 μg/mL

[56]

0.01 nM/L

0.02
2.0 nM/L

[57]

25

5.0 fg/mL

10
10 ng/mL

[58]

Label-free

0.3 pg/mL

1.0 pg/mL to 1.0 μg/mL

[59]

IL/AuNPs/GR/GCE

ALP/fc

0.04 pg/mL

0.1
80 pg/mL

[60]

DPV

PTC-NH2/GCE

AuNPs/STP/ALP

3.9 fg/mL

0.01
100.0 pg/mL

[61]

CV

SWCNTs/GCE

Label-free

0.026 ng/mL

0.1
1.0 ng/mL

[62]

MMP-9

SWV

GRNR/SPCE

PS/PD/Cd

MDM-2

EIS

Au

APE-1

CV

APE-1 tPA

21

3

Abbreviations: AuNPs, gold nanoparticles; AuHS, nanogold hollow microspheres; AuNCs, gold nanoclustures; AgNPs, silver nanoparticles; Apt, aptamer; ALP, alkaline phosphatases; BN, boron nitride; CNTs, carbon nanotubes; CuS, copper sulfide; CPE, carbon paste electrode; CILE, carbon ionic liquid electrode; CLPE, carbon ionic liquid paste electrode; Chi, chitosan; CV, cyclic voltammetry; DPV, differential pulse voltammetry; DμID, disposable microfluidic device; Den, dendrimer; ERGO, electrochemically reduced graphene oxide; EIS, electrochemical impedance spectroscopy; Fc, ferrocene; Fe3O4NPs, iron oxide nanoparticles; f-TiO2, functionalized titanium dioxide; Gal, β-galactosidase; GCE, glassy carbon electrode; GO, graphene oxide; GR, graphene; N-GR, nitrogen-doped graphene; GRNR, graphene nanoribbon; GOx, glucose oxidase; HRP, horseradish peroxide; ITO, indium tin oxide; IL, ionic liquid; LSV, linear sweep voltammetry; MB, magnetic bead; MP, magnetic nanoparticles; MWCNTs, multiwalled carbon nanotubes; MPCE, magnetic carbon paste electrode; MOF, metal organic framework; N-GR, nitrogen-doped graphene; PGA, polyglutamic acid; PPy, polypyrrole; P-1,5DAN, poly-1,5-diaminonaphthalene; PdNCs, palladium nanocages; PEI, poly(ethylenimine); PS, polystyrene; PtNWs, platinum nanowires; PD, polydopamine; PPD, o-phenylenediamine; PB, Prussian blue; PEDOT, poly(3,4-ethylenedioxythiophene); PVPP, poly-vinylpyrrolidone; PDDA, poly-diallyl dimethylammonium chloride; PTC, 3,4,9,10-perylene tetracarboxylic dianhydride; SPGE, screen printed graphite electrode; SPCE, screen printed carbon electrode; SiNPs, silica nanoparticles; SnO2, tin oxide; STP, streptavidin; SWV, square wave voltammetry; TMB, 3,30 ,5,50 -tetramethylbenzidine; Th, thionine; Tyr, tyrosinase; TiP, hollow titanium phosphate sphere; TEPA, tetraethylene pentamine; V2O5, vanadium pentaoxide; ZnO, zinc oxide.

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For example, an electrosynthesized bilayer polymer
based sandwich immunosensor was developed using magnetic beads as nanocarriers [14]. The polymer bilayers of polypyrrole (PPy) and poly-1,5-diaminonaphthalene (P-1,5DAN) were utilized as a conducting base for mouse monoclonal anti-CA15-3 Ab1 on the SPE surface which were coupled by antiCA15-3 Ab2 and horseradish peroxidase (HRP) labeled streptavidin-coated magnetic beads (Fig. 5.3). The biosensor exhibited amplified signal response toward CA15-3 Ag with a detection limit of 0.02 U/mL. Along with the high performance of the prepared immunosensor, the study focused on the use of conducting polymers and magnetic beads as nanocarriers in immunoassays for signal amplification. In another study, a sandwich-based electrochemical sensor for CA15-3 was designed using a capture Ab and anti-CA15-3 Ab1 immobilized at graphene oxide (GO)-modified indium tin oxide (ITO) electrodes [15]. A mixture of two secondary antibodies

(Ab2) conjugated with tyrosinase (Tyr) and β-galactosidase (Gal) in two-enzyme schemes was utilized to amplify the signal via redox cycling. In the presence of CA15-3 biomarker, Gal converted phenyl β-D-galactopyranoside into phenol while Tyr converted phenol into catechol and then to o-benzoquinone which was further reduced onto the electrode surface (Fig. 5.4). Chronocoulometric analysis allowed the detection of CA15-3 with a low detection limit of 0.1 kU/L. Currently, microfluidics-based immunoassays offer advanced biosensing strategies toward biomarker detection. Multiple advantages such as smaller quantities (few nanoliters) of reagents and samples, large surface-to-volume ratio, and highly automated fluid handling facilitate Ab
Ag interactions with improved reproducibility and selectivity [63,64]. Taking benefits of the microfluidics-based sensing technique, a sandwich immunoassay was performed toward CA15-3 Ags. A disposable microfluidic immunoarray device (DμID) was

FIGURE 5.3 Schematic preparation of P(1,5DAN)PPyNW-based immunosensor for CA15-3 biomarker detection [14]. Reprinted with permission from V.A. Nguyen, H.L. Nguyen, D.T. Nguyen, Q.P. Do, L.D. Tran, Electrosynthesized poly(1,5-diaminonaphthalene)/polypyrrole nanowires bilayer as an immunosensor platform for breast cancer biomarker CA 15-3, Curr. Appl. Phys. 17 (2017) 1422
1429.

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the study, the bioconjugated paramagnetic particles suggest their crucial role in designing a low cost, portable, and highly sensitive microfluidic immunoarray for detection of clinically important cancer biomarkers.

5.2.2 Epidermal Growth Factor Receptor

FIGURE 5.4 Signal amplification scheme using a twoenzyme system (tyrosinase and β-galactosidase) for electrochemical immunosensing toward CA15-3 biomarkers [15]. Reprinted with permission from S. Park, A. Singh, S. Kim, H. Yang, Electroreduction based electrochemical enzymatic redox cycling for the detection of cancer antigen 15-3 using graphene oxide modified indium 2 tin oxide electrodes, Anal. Chem. 86 (2014) 1560
1566.

fabricated using the layer-by-layer technique for immobilizing mouse monoclonal MUC1 primary antibody Ab1 on SPE surface [16]. A bilayer of poly-diallyl dimethylammonium chloride (PDDA) and AuNP-modified surface was used for the modification of Ab1 which was then sandwiched by HRP-polyclonal Ab2-labeled magnetic nanoparticles. The high sensitivity of the sensor was attributed to the use of massively decorated paramagnetic particles and to their bioconjugates, with a large quantity of HRP-labeled Ab2 which amplified the detection signals significantly. The amplification strategies integrated with the nanostructured immunosensors and microfluidics resulted in remarkable sensing performance toward detection of CA15-3 with a limit of detection (LOD) of 6.0 μU/mL. The DμID was applied for the detection of CA15-3 in real samples of breast cancer patients and was found applicable for its clinical applications. Through

Epidermal growth factor receptor (EGFR) or human epidermal growth factor receptor 1 (HER-1) is another biomarker which is overexpressed ( . 75.3 μg/L) in breast cancer tumors. It is a cell transmembrane glycoprotein present at the cell surface [65]. On binding to transforming growth factor (TGFα), it triggers cell differentiation and proliferation by activation of tyrosine kinase [66]. Nowadays, impedimetric methods have received particular attention as an extremely sensitive labelfree electrochemical technique for studying biorecognition events at the electrode surface. These impedimetric biosensors are generally based on measuring the dielectric properties of the system as the interface changes which are intrinsic to the biorecognition event. Further, it necessarily does not require any specific labels or any electroactive moieties, which makes the fabrication of sensors very easy. For sensitive detection of EGFR, a label-free impedimetric sensor was fabricated using an anti-EGFR Ab immobilized at an AuNP-protein G modified Au electrode [23]. The changes in the impedance values were found to be proportional to EGFR concentrations in the range from 1.0 pg/L to 1.0 μg/mL with a LOD value of 0.34 pg/mL. The superior analytical activity of the EGFR immunosensor was attributed to the synergetic effect between the AuNPs and the protein-G scaffold which helped in the proper orientation of Ab at the electrode surface. The matrix effect from mouse brain tissue homogenate was also evaluated and the immunosensor showed excellent recoveries ranging from 98.3% to 115%. The study

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5. ELECTROCHEMICAL IMMUNOSENSORS FOR RAPID DETECTION OF BREAST CANCER BIOMARKERS

showed possible opportunities of the developed sensor toward clinical diagnosis and prognosis of breast cancer tumors. Further, a sandwich-based immunosensing strategy was developed for EGFR where a biotinylated EGFR-specific aptamer was immobilized at streptavidin-coated magnetic beads, which were used as capture probe [24]. Anti-EGFR Ab conjugated AuNPs was used as the detection probe. In the presence of EGFR, an immunocomplex was formed and signals were recorded by DPV in the EGFR concentration range of 1.0
40.0 ng/mL with a detection limit of 50.0 pg/mL.

5.2.3 Human Epidermal Growth Factor Receptor 2 and 3 HER-2 is an important prognostic marker which is overexpressed in about 30% of

invasive breast cancers. It is encoded by the ERBB-2 gene. In the absence of tumors, the normal HER-2 levels in blood range of 2
15 μg/L [65]. In patients with highly aggressive breast cancer tumors, the increased levels of HER-2 extracellular domain (ECD) in serum are associated with a poor prognosis making the detection of HER-2 highly important at a very early stage. Immunoassay-based different electrochemical methods have been reported for HER-2 detection [25
35]. For example, in a recent study, a highly sensitive and low-cost impedimetric immunosensor based on single chain fragment variable Ab fragments (scFv) was developed for quantitative detection of HER-2 employing AuNPs/screen printed carbon electrode (SPCE) electrode as transducer [25]. Systematic fabrication of the immunosensor using HER-2 specific scFv on SPCE is shown in Fig. 5.5. The AuNPs facilitated fast

FIGURE 5.5 Representation of immunosensor fabrication using HER-2 specific scFv on SPCE [25]. Reprinted with permission from S. Sharma, J.Z. Rodriguez, R. Saxena, R.O. Kennedy, S. Srivastava, Ultrasensitive direct impedimetric immunosensor for detection of serum HER-2, Biosens. Bioelectron. 106 (2018) 78
85.

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electron transfer and offered a biocompatible surface for immobilization of anti-HER-2 scFv in an oriented manner resulting in an improved Ag binding efficiency. The antiHER-2 scFv modified SPE immunosensor offered simple changes in the impedance at the electrode surface in a wide range of biomarker concentrations from 0.01 to 100 ng/mL with a detection limit of 0.01 ng/mL. In another example, a novel impedimetric immunosensor was fabricated using nanoconjugates of multiwalled carbon nanotubes (MWCNTs) and AuNPs at a carbon ionic liquid electrode (CILE) for detection of HER-2 [26]. CILEs offered multiple advantages for electrochemical sensing such as low cost, easy preparation, wide electrochemical windows, antifouling and easy regeneration of the electrode surface along with high conductivity of AuNPs and MWCNTs. The modification of AuNPs/MWCNTs/CILE was followed by self-assembling of the carboxyl-stabilized colloidal AuNPs at the electrode using 1,6-hexanedithiol (HDT) as a cross-linker and resulted in the construction of AuNPs-HDT/ AuNPs/MWCNTs/CILE sensors. The colloidal AuNPs increased sensitivity of the immunosensor by providing a compatible surface for the immobilization of monoclonal HER-2 Ab1 and for the determination level of HER-2 ECD in the serum samples of several breast cancer patients. Moreover, in another sandwich immunoassay, monoclonal antihuman
HER-2 antibodies (Ab1) were immobilized on the AuNP-modified SPE surface and streptavidin
alkaline phosphatase (STPALP) conjugate labeled detection antibodies (Ab2) were added for the analysis of immunoreactions [28]. The detection of the Ab
Ag interactions was made possible by using an enzymatic substrate 3-indoxylphosphate and silver (Ag1) ions where Ag1 were reduced to metallic silver (Ag0) and analyzed by an anodic stripping LSV. Further, taking advantage of iron

159

oxide nanoparticles (Fe3O4NPs), a sandwichtype immunosensor was developed for HER-2 cancer biomarkers using anti-HER-2 (Ab1)/ Fe3O4NPs modified GCE and colloidal AuNPs/ Fe3O4NPs labeled anti-HER-2 (Ab2) as tracer [29]. Moreover, in a label-free immunosensing anti-HER-2 (Ab1)/Fe3O4NPs modified Au electrode was applied for measuring Ab
Ag interactions [30]. The studies suggested the great significance of Fe3O4NPs in immunoassays toward detection of breast cancer biomarkers which not only provided a compatible matrix for biomolecule immobilization, but also increased the sensitivities of the biointeractions and facilitated the electron transfer processes. In an ultra-advanced immunoassay procedure, a disposable inkjet-printed immunosensor was developed for HER-2 detection [31]. The electrochemical set-up utilized an array of eight inkjetprinted Au working electrodes in combination with an onboard counter and chlorinated inkjetprinted Ag/AgCl quasi reference electrodes. The sandwiched immunoassay involved antiHER-2 (Ab1) immobilization on the working surface and further detection of biorecognition using STP and polyHRP-labeled detection antibody (Ab2). The inkjet-printing technique offered a cost-effective immunoassay with the benefits of high sensitivity, low detection limit, rapid response, and minimal sample usage. Most biomarkers are associated not only with one cancer, but also various cancers. Therefore multibiomarker analysis is essential. The simultaneous detection of biomarkers such as CA15-3 and HER-2 ECD is valuable for the identification of high-risk breast cancer. Justifying the fact, a disposable immunosensor was developed for the simultaneous detection of CA15-3 and HER-2 ECD using dual SPE electrodes [35]. The AuNPs modified electrodes were individually coated with captured Ab1 of monoclonal antihuman CA15-3 and monoclonal antihuman
HER-2 ECD, respectively shown in Fig. 5.6A. The biointeractions of Ab
Ag was

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5. ELECTROCHEMICAL IMMUNOSENSORS FOR RAPID DETECTION OF BREAST CANCER BIOMARKERS

self-assembled 4-aminothiophenol monolayer modified Au electrode [36]. The immunosensor exhibited a linear dynamic range of 0.4
2.4 pg/mL with a detection limit of 0.28 pg/mL with high sensitivities.

5.2.4 Vascular Endothelial Growth Factor Receptor

FIGURE 5.6 (A) Immunosensing scheme for simultaneous detection of HER-2 ECD and CA15-3 biomarkers and (B) corresponding LSV plots [35]. Reprinted with permission from R.C.B. Marques, E.C. Ramaa, S. Viswanathan, H.P.A. Nouws, A.C. Garcı´a, C.D. Matos, et al., Voltammetric immunosensor for the simultaneous analysis of the breast cancer biomarkers CA 15-3 and HER-2 ECD, Sens. Actuat. B Chem. 255 (2018) 918
925.

analyzed using ALP-labeled Ab2 in each case. The signals were recorded as the peak current intensity of the enzymatically generated metallic silver by LSV (Fig. 5.6b). HER-3 is a membrane-bound protein encoded by the ERBB-3 gene and highly expressed in breast cancer. In the absence of tumors, its normal level in blood is in the range of 0.06
2.5 μg/L [65]. For its detection, an impedimetric immunosensor was developed by immobilizing anti-HER-3 Ab on a

Vascular endothelial growth factor receptors (VEGFRs) are specific tyrosine kinase receptors divided into subtypes of VEGFR-1, VEGFR-2, and VEGFR-3. Among the three types, VEGFR-2 triggers most of the angiogenic functions. VEGFR-2 is an important biomarker for breast cancer where its blood concentration level . 15 ng/L is indication of presence of breast cancer tumors [65]. For detection of VEGFR-2, a sandwich immunoassay was designed by immobilizing anti-VEGFR-2 Ab1 at chitosan/rGO/thionine modified GCE [37]. An HRP-labeled Ab2 was used as the detection Ab, which catalyzed the oxidation of thionine by H2O2. VEGFR-2 was quantified by DPV in the linear concentration range of 0.4
86.0 pM with a detection limit of 0.28 pM.

5.2.5 Carcino Embryonic Antigen Carcino embryonic Ag (CEA) is a type of glycoprotein produced by cells of the gastrointestinal tract during embryonic development. The normal level of CEA in blood is 2.5 μg/L among nonsmokers and 5 μg/L among smokers [65]. An elevated level of CEA represents the progression or recurrence of tumors cells including breast cancer. For detection of CEA, various electrochemical immunosensors have been developed [38
47]. In a displacementbased immunoassay, a biosensor for CEA was developed where anti-CEA Ab1 was immobilized on Pt/thionine (Th)/GR modified GCE [38]. During immunoreactions, Ab1 captured Ag was coupled with star-like Pt

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161

FIGURE 5.7 (A) Preparation of HRP/ anti-CEA/AuNPs/ZnONPs/Au sensing platform for CEA detection and (B) electrochemical cell for injection flow analysis [47]. Reprinted with permission from P. Norouzi, V.K. Gupta, F. Faridbod, M.P. Hamedani, B. Larijani, M.R. Ganjali, Carcinoembryonic antigen admittance biosensor based on Au and ZnO nanoparticles using FFT admittance voltammetry. Anal. Chem. 83 (2011) 1564
1570.

nanowires (PtNWs) and HRP-labeled secondary antibodies (Ab2) which catalyzed the reduction of H2O2. In the presence of the target biomarker, the NW/enzyme complex was displaced from the sensing interface, causing a reduction in the generated current which was analyzed by DPV. Furthermore, taking advantage of the high conductivity and large surface area of metal organic frameworks (MOFs), a sandwich immunoassay was developed using AuNPs/Cu-MOF nanoconjugate labeled Ab2 as the detection probe while electrochemically reduced graphene oxide (ERGO) was utilized as the capture probe for Ab1 [39]. This new class of signal probes marked a difference from traditional probes toward sensitive detection of biomarkers with high analytical performances. Another sandwich-based platform was fabricated for CEA biomarkers using Th immobilized AuNPs encapsulated dendrimer as the capture probe [44]. The Ab2 labeled with two enzymes, HRP and glucose oxidase

(GOx), was immobilized onto the surface of MWCNTs and utilized to enhance the electrocatalytic reduction of H2O2 which was monitored by SWV. Further, in advanced electrochemical sensing, a fast Fourier transform admittance biosensor was developed for CEA detection [47]. The anti-CEA Ab was immobilized on the Au electrode which was coated with highly conductive bilayer film of zinc oxide nanoparticles (ZnONPs) and AuNPs. The positively charged HRP enzyme was used to block the active sites of the sensing interface to prevent nonspecific binding. In the absence of CEA, the enzyme catalyzed the reduction of H2O2 and subsequently the oxidation of ferrocyanide to generate current. The presence of the target biomarker decreased the turnover rate of the enzyme due to blocking the active center of the enzyme which led to a further decrease in the current signal. The systematic mechanism of FET immunosensor fabrication and CEA biomarker detection is shown in

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5. ELECTROCHEMICAL IMMUNOSENSORS FOR RAPID DETECTION OF BREAST CANCER BIOMARKERS

Fig. 5.7. The admittance biosensor applied fast Fourier transform continuous SWV which generated sensitive, selective, and highly rapid response for detection of CEA biomarkers. The technique was applied as a detector in a flow injection system. The admittance reduction current decreased linearly in two concentration ranges of CEA between 0.1
70.0 ng/mL and 70.0
200.0 ng/mL with a low detection limit of 0.01 ng/mL.

5.2.6 Breast Cancer Type 1 and 2 Susceptibility Proteins Breast cancer type 1 and 2 (BRCA-1 and BRCA-2) are antioncogenes in women especially who are predisposed to breast cancer genetically. These are normally expressed in breast cells and other tissues during their involvement in the repair of chromosomal damage. Generally, malfunctioning of BRCA-1 and BRCA-2 occurs due to mutation in their corresponding gene which eventually results in the increased risk of breast cancer [67]. Using techniques of electrochemical biosensing, a sandwich-based amperometric immunosensor was fabricated for detection of BRCA-1 [48]. An anti-BRCA-1 antibody (Ab1) was immobilized at Th/poly-vinylpyrrolidone (PVPP)/GR modified GCE surface while the detection antibody (Ab2) was anchored at the surface of mesoporous silica SBA-15 in which HRP was entrapped uniformly. Signal amplification was successfully achieved by treating the signaling tag with the [BMIM]BF4 ionic liquid (IL) which enhanced the electrochemical activity of the enzyme and facilitated ion transport. The synergistic effect of IL, SBA-15, Ab2, and HRP enzyme collectively enhanced the sensitivity of the immunosensing within the dynamic concentration range of 0.01
15.0 ng/L and a detection value of 4.9 pg/mL.

5.2.7 Cluster of Differentiation 146 Antigen and 105 Antigen (CD-146 and CD-105) Cluster of differentiation 146 Ag (CD-146) is a cell adhesion molecule which belongs to the immunoglobulin superfamily. It is identified as a melanoma and breast cancer progression marker. The normal level of CD-146 in blood serum of healthy individuals is generally 309 μg/L [68]. For detection of CD-146, a sandwich-based amperometric immunosensor was fabricated in which anti-CD-146 antibody (Ab1) was immobilized at rGO-tetraethylene pentamine (TEPA) modified GCE [49]. This modification provided a large number of amino groups on the electrode surface to enhance the loading capacity of antibodies. The secondary Ab was modified with TiO2 colloidal sphere and Au/Pd nanoparticles and assay was accomplished by measuring the amperometric response toward electrocatalytic reduction of H2O2. Furthermore, Cluster of differentiation 105 Ag (CD-105) is a transforming growth factor receptor (TGF) β1 and β3 in vascular endothelial cells. The level of CD-105 is found highly upregulated in patients ( . 0.9 μg/L) affected with metastatic breast cancer [69]. Involving sandwich-based immunosensing strategies, a voltammetric biosensor was prepared using anti-CD-105 antibody (Ab1) which was immobilized onto the surface of AuNPmodified Au electrode [50]. The detection probe was prepared by immobilizing a thionine acetate modified anti-CD-105 antibody (Ab2) with PtNPs. The complexation reaction of Ab
Ag was monitored using CV in the linear CD-105 concentration range of 1.3
200.0 μg/mL.

5.2.8 Interleukin-6 and -8 Interleukin-6 (IL-6) is considered as a breast cancer biomarker as its upregulated levels

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( . 10
75 ng/L) indicate the presence of breast cancer tumor cells [70]. For the detection of IL-6, a sandwich-based voltammetric immunosensor was prepared using magnetic beads labeled anti-IL-6 antibody (Ab1) captured within the wells of a microtiter plate positioned at the top of a magnet [51]. AgNPtitanium (AgNP-TiP) phosphate spheres were used to label the secondary antibody (Ab2). DPV was applied for the interrogation of IL-6 level in the linear range of 0.0005
10.0 ng/mL with a detection value of 0.1 pg/mL. In a related approach, poly-HRP-STP conjugate was used as the label for detection antibody (Ab2). The immunoreaction was monitored by amperometry after addition of a mixture of hydroquinone and H2O2 [52]. Using a similar principle, a sandwich-based amperometric immunoassay was developed using an array of eight gold compact disks integrated into a microfluidic device [53]. The capture anti-IL-6 antibody (Ab1) was immobilized at Au electrodes while STP-HRP-labeled Ab2 was used as the detection probe. A mixture of hydroquinone and H2O2 was then injected into the microfluidic device to generate the amperometric response with a detection value of 10.0 fg/mL. In another study, boron nitride (BN)/ PDDA/gold nanoclustures (AuNCs) nanocomposite was used as a signal tag for secondary antibody (Ab2) while GR was used as capture probe for Ab1 [54]. The biomarker was quantified using SWV with a detection limit of 1.3 pg/mL. The elevated levels of interleukin-8 (IL-8) ( . 13 ng/L) can also be used for the diagnosis of breast cancers [65]. For the detection of IL-8, a sandwich-based amperometric immunoassay was designed using SPCE as an immobilization base for Ab1 while MWCNTs/ HRP was used as signal tags for Ab2 for enhanced electrochemical response of the immunoreactions [55]. IL-8 was quantified simultaneously with prostate specific Ags with a detection limit of 8.0 pg/mL.

163

5.2.9 Other Important Biomarkers Biomarkers such as tenascin-C (TN-C), p53 tumor suppressor protein (p53), matrix metallopeptidase 9 (MMP-9), murine double minute 2 (MDM-2), apurinic/apyrimidinic endonuclease 1 (APE-1), and tissue plasminogen activator (tPA) are other significant biomolecules that are considered as breast cancer tumor markers [65]. For example, TN-C is a glycoprotein which is overexpressed in breast cancer tumor cells. An impedimetric assay for TN-C was developed using an anti-TN-C antibody (Ab1) at the Au electrodes. TN-C Ag was captured by Ab1 and the nonoccupied binding sites were identified HRP-conjugated antibody (Ab2) which was specific to the Ag binding fragment [56]. The enzyme catalyzed the conversion of soluble 3-amino-9-ethyl carbazole into insoluble 3-azo9-ethyl carbazole, which was precipitated on the electrode surface and significantly increased the interfacial resistance. Further, p53 is a potent transcription factor that plays an important role in controlling cellular responses to stress factors. Loss of p53 function results in tumor formation and gene mutation. Moreover, p53392 regulates the oncogenic function of mutant forms of the p53 protein in breast cancer. The normal level of p53 in blood is 150.0 ng/L and an irregular level is used for diagnosis of breast cancer tumors [65]. In a sandwich-based voltammetric immunosensor for p53392, an anti-p53392 antibody (Ab1) was immobilized at AuNPs/SPCE while HRP-GO conjugate labeled secondary Ab was used for signal amplification [57]. After a sandwich immunoreaction, the HRP-p53392-Ab2-GO was captured onto the electrode surface which produced an amplified electrocatalytic response by the reduction of enzymatically oxidized Th in the presence of H2O2 (Fig. 5.8). The increase in the response current was found to be proportional to the

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FIGURE 5.8 Signal amplification by multistep redox reactions using HRP/GO bioconjugate labeled secondary antibody (Ab2) and AuNP/SPCE as immobilizing surface for capture antibody (Ab1) for detection of p53392 antigen [57]. Reprinted with permission from D. Du, L. Wang, Y. Shao, J. Wang, M.H. Engelhard, Y. Lin, Functionalized graphene oxide as a nanocarrier in a multienzyme labeling amplification strategy for ultrasensitive electrochemical immunoassay of phosphorylated p53 (S392), Anal. Chem. 83 (2011) 746
752.

p53392 concentrations in the linear range of 0.02
2.0 nM with the detection limit of 0.01 nM monitored by SWV. In addition, MMP-9 is a gelatinase subgroup of the matrix metalloproteinases which degrades type IV collagen in the basement membrane and helps to separate the epithelial cells from the underlying stroma. Elevation of the serum level of MMP-9 ( . 189 μg/L) is generally observed in breast cancer patients [65]. For MMP-9 detection, a sandwich-based voltammetric immunosensing platform was developed taking an anti-MMP-9 antibody (Ab1) at graphene nanoribbon (GRNR) modified SPCE [58]. A polystyrene spheres/ polydopamine/cadmium nanohybrid (PS/ PD/Cd21) was utilized to label the secondary antibody (Ab2) which was used to monitor immunoreactions by SWV with a detection limit of 5.0 fg/mL. MDM-2 is a protein which is encoded by the MDM-2 gene and acts as a negative regulator for the p53 tumor suppressor. It is overexpressed in different kinds of tumor cells including breast cancer [65]. A label-free impedimetric immunosensor for MDM-2 was constructed using anti-MDM-2 antibody (Ab1) immobilized at Au electrode using 1,4-phenylene diisothiocyanate as a linker [59]. The impedance changed upon binding of the biomarker to the electrode surface in the wide dynamic range of 1.0 pg/mL to 1.0 μg/mL.

The immunosensor exhibited a good performance with a detection limit of 0.29 pg/mL in comparison with ELISA for the analysis of the MDM-2 in the cancerous mouse brain tissue homogenates. Moreover, the biosensor had high selectivity against EGFR with long-term stability and reproducibility. Additionally, APE-1 is a multifunctional protein in the DNA base excision and repair pathway. High levels of APE-1 indicate tumors of various cancer cells including breast cancer. A cyclic voltammetry-based immunosensor was fabricated for detection of APE-1 using [BMIM]BF4 IL/AuNPs/GR nanoconjugate as Ab1 immobilization surface [60]. Further, ALP-fc labeled Ab2 was used for signal amplification. The IL/AuNPs/GR significantly facilitated the high loading of ALP as a carrier for fc-Ab2 and accelerated electron transfer. ALP catalyzed its substrate ascorbic acid 2-phosphate (AA-p) and resulted in the production of ascorbic acid (AA) which, in turn, was oxidized and catalyzed by the Fc/Fc1 coupling. In a related approach, the ALP enzyme was utilized to catalyze the oxidation of AA to amplify the signal via a triple-signal amplification strategy as shown in Fig. 5.9 [61]. The first step included the labeled biotinylated ALP on the nickel hexacyanoferrates nanoparticle-decorated Au nanoparticles (Ni
AuNPs) for the catalysis of ascorbic acid 2-phosphate (AA-P) into AA. Further the

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165

FIGURE 5.9 Triple-signal amplification scheme for detection of AEP-1 antigen: (A) stepwise fabrication of ALP/STP/ Ab2/Ni
AuNCs bioconjugates; (a) absorption of NiNPs, (b) Ab2 loading; (c) blocking with STP, and (d) binding ALP. The inset (B) shows the molecular structure of PTC-NH2 [61]. Reprinted with permission from J. Han, Y. Zhuo, Y. Chai, Y. Xiang, R. Yuan, Y. Yuan, et al., Ultrasensitive electrochemical strategy for trace detection of APE-1 via triple signal amplification strategy, Biosens. Bioelectron. 41 (2013) 116
122.

signal was amplified by electrochemical oxidation of AA because of the catalysis of Ni
AuNPs. Moreover, due to AuNP-modified STP, the stoichiometry of ALP was increased, which further helped in signal amplification. Further, a CV-based, label-free immunosensing strategy was developed for tPA biomarkers which plays a key role in cellular processes such as angiogenesis. A SWCNTs modified GCE sensor was used to immobilize anti-tPA antibodies through 1-ethyl-3-(3 dimethylaminopropyl) carbodiimide hydrochloride (EDC)/ N-hydroxy-succinimide coupling [62]. The immunosensor showed high performance toward biomarkers with a very low detection limit of 0.026 ng/mL. Altogether, the simple assay protocols facilitated the production of a new range of immunodiagnostic sensors for breast cancer diagnosis.

5.3 FUTURE PROSPECTS AND CHALLENGES Developing biosensors for breast cancer biomarkers has attracted much attention in recent years. However, biomarker discoveries and the development of diagnostic tools for early stage breast cancer detection are still in their infancy. Although electrochemical immunoassays have been very effective biotransducers, the biomarkers identified for breast cancer are required to benchmark their specificities, sensitivities, and performance against the developed diagnostic standards. To ensure the reproducibility of the electrochemical immunoassays, biorecognition processes at electrode surface must be carefully controlled in multiple aspects. For example, fabrication of biosensors should be accurately

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handled in terms of size/shape/number of nanomaterials or signal tags. The signal amplification strategies need to resolve the nonspecific adsorption issues which controls the detection limits in electrochemical immunoassays. Despite advancements, designing highly specific and true biospecific systems for breast cancer biomarkers is still very challenging as the biomarkers show their elevated and upregulated levels in tumor cells of various other cancers. Therefore a major issue for electrochemical biomarker sensors is to develop reliable strategies for simultaneous detection of multiple biomarkers in complex biological samples to avoid false positives. For this, combined detection of biomarkers using electrochemical transducers for cancer detection is highly anticipated with a follow-up on cohorts of women with, or without, clinical history of breast cancer to establish the true diagnostic values of biomarkers. The motivation of multianalyte analyses of biomarkers based on lab-on-chip diagnostic devices is highly encouraging [7]. Integration of electrochemical and electroluminescence immunoassays into a microfluidic system (e.g., paper-based microfluidics) manifests a strong sensing base for biosensor construction particularly for diagnostic and biomedical applications [71]. The development and progress of these advanced detection devices will help in rapid clinical diagnoses of cancer, possibly at early stages. However, detection strategies proposed for cancer biomarker detection necessarily require standardization of the pre- and postanalytical protocols such as sample processing, storage, and optimization of experimental conditions for true validation of the assays and more genuine performance of the developed biosensor.

5.4 CONCLUSION With the remarkable progress in nanotechnology and biosensor techniques, novel electrochemical strategies have been developed for

the detection of breast cancer biomarkers. The reliable electrochemical detection devices have been integrated with the principles of immunology to host specific Ag
Ab reactions at the transducer surface and to detect the biomarkers quantitatively. Electrochemical immunoassays have enabled selective detection of breast cancer biomarkers in different biological fluids like blood, urine, saliva, and others with great analytical performances. Significant progress in signal amplification has been made by using multienzyme labeling on carbonaceous materials, conducting polymers, magnetic beads, and metal nanoparticles, which further improves the detection sensitivity of the immunoassays. However, important gaps and challenges must be addressed and implemented before commercialization of the developed biosensors for actual clinical practices toward the detection of breast cancer biomarkers.

Acknowledgment This research was supported by the Chinese Academy of Sciences (CAS) through the President’s International Fellowship Initiative for Postdoctoral Researcher (2015PM002).

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C H A P T E R

6 Functionalized Advanced Hybrid Materials for Biosensing Applications Oana Hosu, Anca Florea, Cecilia Cristea and Robert Sandulescu* Analytical Chemistry Department, Faculty of Pharmacy, Iuliu Ha¸tieganu University of Medicine and Pharmacy, Cluj-Napoca, Romania

6.1 INTRODUCTION The Oxford Dictionary defines “hybrid” as “a thing made by combining two different elements” (noun) or something “of mixed character; composed of different elements” (adjective). Referring to advanced materials, generally “hybrid” means increasingly sophisticated materials, such as miniaturized, recyclable, environment friendly, energy efficient, reliable, and inexpensive materials. In this context, advanced hybrid materials represent a very large and heterogeneous class of materials, starting with molecular and supramolecular assembled materials, polymers, or nanosized objects to nanostructured and hybrid architectures with inorganic, organic, or biological characteristics, having or combining particular properties that cannot be found in other types of materials. It is then easy to understand that the attempt to conceive a rigorous and comprehensive classification of hybrid materials is a very difficult task.

A classification attempt could be as follows: inorganic
inorganic, inorganic
organic, organic
organic, multifunctional hybrid materials, and biohybrid materials. The design and development of functionalized advanced hybrid materials involve a large range of synthesis strategies, such as achievement of colloidal and titanium-oxo clusters, self-assembly of inorganic or organic compounds, ionic liquid
based hybrid materials, chemical or electrochemical polymerization, pre- or postsynthetic functionalization, molecularly imprinting, metal
organic frameworks, block copolymer templating, molecular or block-assembling bottom
up processes, molecular and supramolecular hybrid organic
inorganic interfaces, silicate and silicabased biohybrids, and even whole cell
based hybrid materials [1]. Inorganic compounds, endowed with high chemical and thermal stability that allows their application under different operating conditions can be obtained by inexpensive processes

*Corresponding author.

Advanced Biosensors for Health Care Applications DOI: https://doi.org/10.1016/B978-0-12-815743-5.00006-8

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and be easily deposited in the form of thin or thick films by different techniques. On the other hand, organic compounds are characterized by synthetic versatility and reactivity, which make it possible to modulate the molecular structure. This could be an advantage if the enhancement of the selectivity toward a target analyte is followed. Nevertheless, the drawbacks of single inorganic or organic sensing materials, namely, high-operating temperature and low selectivity for inorganic sensing materials, and poor chemical stability and mechanical strength for organic sensing materials, could restrict their practical application. All these individual drawbacks could be fulfilled by the addition of inorganic particles into the organic materials to form hybrids [2
5]. Generally, a simple approach is to use the inorganic material as a protective one while the organic matrix is dispersed by different techniques, or both materials could present a synergistic effect, as both components are involved in the sensing mechanism. For sensor applications over the past few decades, many reports have been published based on the use of the organic or inorganic hybrid sensing materials, gas sensors, humidity sensors, ultraviolet sensors, strain sensors, and other sensor development [6]. These new hybrid materials become attractive for many new electronic, optical, magnetic, or catalytic applications since their properties or performances can be improved. A chemical sensor is a device that transforms chemical information into an analytically useful signal [7]. The sensors contain two basic functional units, namely a receptor part which transforms the chemical information into a form of energy that may be measured by the transducer part capable of transforming the energy carrying the chemical information about the sample into useful analytical signals. To promote selectivity, the transducer is usually functionalized with various receptors [8].

A particular class of sensors is represented by biosensors which according to IUPAC are “a self-contained integrated device, which is capable of providing specific quantitative or semiquantitative analytical information using a biological recognition element (biochemical receptor) which is retained in direct spatial contact with an electrochemical transduction element” [9]. The biological recognition element may be based on a chemical reaction catalyzed by, or on an equilibrium reaction with, macromolecules that have been isolated, engineered, or present in their original biological environment [9]. Biosensors are based on the same principles as chemical sensors. Another definition of a biosensor establishes that it is an integrated device that consists of a biological recognition species in direct contact with a transduction element. Therefore biosensors can be classified according to the biological recognition element (immuno, enzymatic, DNA, and whole-cell biosensors) or the signal transduction method (optical, mass-based, electrochemical, and thermal biosensors). To summarize, a biosensor combines a biological recognition element with a suitable transduction method so that a meaningful signal can be achieved when binding, or some reaction occurs between that element and a target species [7]. Several examples of hybrid nanomaterials will be presented in this chapter with a focus on the biosensor’s designs.

6.2 ADVANCED INORGANIC HYBRID MATERIALS Advanced inorganic nanohybrids such as colloidal clusters, titanium-oxo clusters, alloy and metal oxide hybrids, self-assembled nanorods, and mesoporous-silica hybrid materials are presented, emphasizing their ability to be used in sensor development. Some of these materials could find other interesting health

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6.2 ADVANCED INORGANIC HYBRID MATERIALS

applications like energy storage or power sources in implantable cells and batteries, drug delivery systems, or nanorobots (Janus particles) [10,11].

6.2.1 Colloidal Clusters Recent studies on clusters of anisotropic colloidal particles were reviewed by Morphew and Chakrabarti [12], emphasizing their selfassembly into finite supracolloidal structures, mimicking the symmetry of molecular structures through the colloidal interactions due to the presence of one or more anisotropy attributes. The recent progress and future prospects in the field of new functional materials fabrication through the assembly of colloidal building blocks, the strategies to synthesize and assemble such colloids, and the major limitations and challenges were recently reviewed by Ravaine and Duguet [13]. The design and fabrication of colloidal particles able to selfassemble and organize themselves via directional and specific interactions due to mutual recognition prompted great and increasing interest. These new colloidal particles, named patchy particles, exhibit site-specific engineering of their surface, which is characterized by an increasing number of patches with multiple functionalities. Thus one-patch (also called Janus particles), two-patch (identical or different), three-patch, four-patch, five-patch, and six-patch microparticles and nanoparticles were described, emphasizing the shape of a micelle and a vesicle of Janus particles, the formation of small clusters of Janus amphiphilic particles in the presence of 1 mM KNO3, and of a colloidal macromolecule resulting from the self-assembly of monodimpled spheres. The main strategies to achieve patchy particles have been summarized. The self-assembly of multishell structures from particles with two types of patches allows not only a better understanding of the formation of virus

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capsids (protein shells with specific sizes and structures designed to encapsulate the viral genome), but also the interactions between the constituent proteins [14]. Silver nanoparticle (AgNP) colloids synthesized in water by using potato starch as the reducing and stabilizing agent were used to detect ascorbic acid by surface-enhanced Raman scattering (SERS) in aqueous solutions. The intensity increase of the Raman peak at 1386 cm21 (assigned to deformation modes of the starch structure) with the ascorbic acid concentration increase is related to a decrease of the gap between dimers and trimers of the AuNP clusters produced by the presence of ascorbic acid in the colloid. The 0.02 mM LOD was obtained for ascorbic acid in the range of 0.02
10 mM [15].

6.2.2 Titanium-Oxo Clusters Titanium and zirconium oxo-cluster structures, synthesis procedures, and the organic functionalization of titanium-oxo
based nanobuilding blocks allowing the development of Lego-like chemistry and the assembling of a large variety of structurally well-defined complex hybrid architectures were presented in reviews by Rozes et al. [16] and Schubert et al. [17]. A graphene
titanium dioxide (GO
TiO2) nanocrystal hybrid was achieved by a two-step method in which TiO2 was first coated on GO sheets by hydrolysis and crystallized via a hydrothermal treatment in the second step. The novel, advanced hybrid material based on GO
TiO2 nanocrystals showed threefold superior photocatalytic activity to other TiO2 materials and, therefore, are very promising for applications in lithium-ion batteries [18]. In a recently published review, Zhu et al. [19] described the synthesis and the photocatalytic applications of some novel Ti-based metal
organic frameworks (Ti
MOFs), with unique structural properties, high thermal and chemical stability, and very promising optoelectronic

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and photocatalytic behaviors. Ti
MOFs proved to be photocatalysts with high sunlight harvesting efficiency and photocatalytic performance, finding many applications in storage and energy fields, photocatalytic redox and polymerization reactions, organic pollutant degradation, water splitting, and sensors.

6.2.3 Alloy and Metal Oxide Hybrids Remarkable advancements in functional material technologies have accelerated the use of novel electrocatalytic nanomaterials with specific sizes, shapes, and self-assembled architectures as new materials, and their hybrids and composites, for diverse applications [20,21]. Intense and ongoing research is currently focused on improving the analytical parameters of the carbon-based sensors [carbon nanotubes (CNTs), graphene] to fulfill the requirements of practical biosensors in terms of very low detection limits and high selectivity and to make them commercially attractive for analytical sensing systems [21,22]. One effective strategy to meet these requirements is blending carbonaceous materials with metal oxide nanostructures to form hybrid composite architectures, which could improve sensor performance [23]. To date, a large number of papers have been reported for the use of hybrid composites based on GO- or CNTmetal oxides. It is clear that the use of carbonaceous nanomaterials as a support to effectively disperse metal oxide NPs improves catalytic properties of the hybrid through the increase of the surface area and, moreover, through the synergistic effect of the physical properties of the obtained hybrid material [24
26]. When graphene or CNT are integrated with semiconductor metal oxides, they promote photogenerated electrons through π
π bond interactions and significantly hinder the charge recombination in the semiconductor
oxide materials

[27]. Composite sensors involve metal oxides like SnO2, WO3, ZnO, Co3O4, TiO2, Fe2O3, Fe3O4, CuO, MnO2, Cu2O, NiO, Al2O3, Sb2O3, and ZnFe2O4 [22]. For example, when GO-NPs hybrids are used as electro or photocatalysts, an enhanced catalytic activity, more active surface area, and good stability are demonstrated compared to monometallic NPs or GO counterparts [27]. The hybridization of metal oxides with carbonaceous nanomaterials has yielded very interesting and varied applications. The synergistic effects including (1) bifunctional and (2) electronic effects between two metallic species have an effective impact on the electrocatalytic activity of the nanocomposite electrode surface. To date, several platinum (Pt)-based catalysts with improved electrocatalytic performances have been studied such as Pt
Ag [28], Pt
Au [29], Pt
Pd [30], Pt
Ru [31], and others. Among these materials, Pt
Pd bimetallic catalysts have captured much interest because of its unique characteristics in contrast to their individual counterparts. A new nanocomposite containing Pt
Pd nanoflowers promoted with 2D nanosheet structured cuprous oxide (Cu2O) supported on reduced graphene oxide (rGO) (PtPd
NFs/ Cu2O
NSs/rGO) was developed for methanol electrooxidation [32]. Another bimetallic nanoparticle biosensor (Pt
PdNPs
MWCNTs) was developed by Chen et al., for the amperometric detection of glucose showing a wide linear range (6.2 3 1025
1.4 3 1022 M) and a low detection limit of 31 μM [33]. Nickel hydroxide has shown extraordinary promise for hybrid supercapacitors because of its high-theoretical capacity and low cost, but is usually limited by its low-power density and complex synthetic processes. By a self-assembly, environment-friendly solvothermal-precipitation method [34], a 3D porous Ni (OH)2 was synthesized with enhanced electron transportation and ion diffusion properties. An AuNPs
Fe3O4 nanocomposite modified electrode for trace level detection of arsenic was employed [35].

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Good electrical conductivity and fast electron transfer on the electrode surface contribute to improving the sensitivity of immunosensors [36,37]. Li et al. [38] developed a hybrid electrochemical immunosensor using a combination of fullerene
palladium platinum alloy
graphene nanoribbons (n-C60
PdPt/N
GNRs) as the sensing matrix for proprotein convertase subtilisin/kexin type 9 (PCSK9), a potential cardiovascular disease biomarker. Pd
PtNPs were used as a signal enhancer and catalyzer for the reduction of H2O2, thus promoting electron transfer [33]. Furthermore, in order to increase the loading capacity of Pd
Pt nanoparticles, amino-functionalized fullerene (n-C60) was introduced as a carrier owing to its large specific surface area, good biocompatibility, better hydrophilicity, and abundant amino groups for further modification [39,40]. A facile, one-step hydrothermal method to decorate SnO2 nanostructures on few-layered graphene was demonstrated for dopamine detection. The prepared nanohybrids exhibited an increase in current with respect to an increased analyte concentration over a wide range of 5 3 1029
2 3 1027 M with an LOD of 6.3 nM [41]. In bioanalytical applications, the interesting unique properties of AuAg alloy NPs, including superior conductivity, excellent catalytic activity, and high chemical stability [24,26,42], have led to the potential development of electrocatalysts for electrochemical sensors [43
45]. In this context, a high-quality, graphene-encapsulated AuAg alloy nanohybrid was successfully fabricated for serotonin showing a very low detection limit (1.6 nM) in a wide linear detection range (2.7 3 1029
4.8 3 1026 M) [46].

6.2.4 Self-Assembled Inorganic Nanorods The synthesis, functionalization, self-assembly, and sensing applications of Au nanorods, with unique optical and electronic properties

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which are dependent on their shape, size, and aspect ratio, are discussed in detail by Vigderman et al. [47]. Thu, single-crystalline and pentahedrally twinned rods are synthesized by wet chemistry methods and polycrystalline rods are synthesized by templated deposition. Their functionalization, which is necessary for different applications, depends on the surfactant coating and is achieved by covalent and noncovalent procedures. The significant progress made in the control of the nanorods assembly into various arrangements has a huge impact on the measurable properties of the rods, making it particularly applicable toward sensing a variety of analytes. The controlled self-assembly of inorganic nanoparticles can be achieved by deliberate surface modifications, which allow recognition between host and guest molecules. The structures and properties of some typical host molecules such as cyclodextrins, cucurbit[n]urils, calix [n]arenes, and pilla[n]arenas, and their supramolecular complexes with guest molecules, which are frequently used for nanoparticle assembly are described by Li et al. [48]. The dynamic selfassembly of nanoparticles lead to special assembly behaviors such as, stimulus-responsive or reversible self-assembly, and the self-assembly systems involving host
guest interactions have potential applications in sensing and catalysis. The colloidal TiO2 nanorods self-assembled into nine arrays, three of these arrays being highly ordered superlattices over large areas, which was recently reported for the first time by Zhang et al. [49]. The facile self-assembly procedure is driven by the density-driven phase evolution and entropic depletion attraction and can be used for the design and fabrication of various materials into self-assembled nanoparticle superlattices, which can be easily tailored for the fabrication of various functional materials and a wide range of devices. Carbon-doped golden wattle-like TiO2 microspheres with remarkable visible light photocatalytic activity by the growth and

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self-assembling of TiO2 single crystal nanorods simultaneously with in situ carbon doping were reported by Xu et al. [50]. The strategies for nanodrug fabrications, their applications in tumor therapy, and the challenges and perspectives of peptide-mediated assembled nanodrugs for potential clinical translation were recently reviewed by Li et al. [51]. The authors describe how peptides and peptideconjugates are designed for controlling and mediating the formation of the assembled nanodrugs, the rational design of intermolecular interactions between drugs and peptides, in vitro and in vivo drug delivery, and their antitumor therapeutic effects. The hydrophobic effect of the supramolecular self-assembly of peptides and drugs is responsible for both colloidal and circulation stability in vivo against dilution and blood-flow shearing, without any consequence on drug function. The drug itself is one of the building blocks for the supramolecular assembly and nanostructure formation, which assures its remarkable stability and high-loading efficiency in a controlled and tunable way.

6.2.5 Mesoporous-Silica Hybrid Materials Generally, silica-based materials are electroinactive and insulators, and for electrochemical applications their electrical conductivity must be improved by conductive carbon nanomaterials or metal nanoparticles. Despite this primary “handicap,” silica is a biocompatible platform for the immobilization of most bioelements and is being employed in the development of many inorganic
organic biohybrids or conductive biocomposites. Furthermore, the ordered mesoporous materials recently achieved led to the development of many sensors and biosensors based on mesoporous silica [52
54] with improved electroanalytical performance in comparison to those fabricated with the nonordered silica materials [55].

The electrochemical applications of core
shell and nanostructured mesoporous nanoparticles and oriented mesoporous-silica thin films were recently reviewed in an excellent paper by Walcarius [52]. Silica can act as a core (SiO2@shell) supporting a reactive shell for metal nanoparticles (SiO2@AuNPs) or amino-functionalized mesoporous silica shell (SiO2@mSiO2NH2). Silica can be a shell as well (core@SiO2) around another metal oxide or a metal core which can be modified with different functional groups, decorated with metal NPs and bioelements, or simply acting like a protective shield. Mesoporous-silica nanoparticles proved to be very good nanocarriers for many compounds including bioelements (e.g., enzymes, antibodies), and by the incorporation of metal nanoparticles (e.g., AgNPs, AuNPs) and carbon nanomaterials (e.g., graphene) in the porous silica beads, accelerated and improved electron transfer rates were obtained [53,56]. Silica-coated magnetic beads provide large surface areas for immobilizing aptamers, and silica beads decorated with AuNPs improve the signal amplification of sandwich-type affinity sensors (e.g., immunosensors, genosensors) [57]. The reproducible detection of the mutated apolipoprotein E gene associated with Alzheimer’s disease [58] has been reported with a graphene and mesoporous-silica hybrid nanomaterial
based electrochemical.

6.3 ADVANCED ORGANIC
INORGANIC HYBRID MATERIALS Coupling inorganic materials with organic compounds leads to attractive composite materials with new remarkable properties that open the way for next-generation sensors. Various nanosized materials such as CNTs, graphene, metal materials, and magnetic particles were functionalized with (or incorporated into)

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organic matrices leading to materials with unique properties, for example, increased surface area
to-volume ratio, improved electronic properties, and remarkable chemical and physical properties. Depending on the linking manner, organic
inorganic hybrid materials have been classified in two distinct classes: Class I, in which organic and inorganic compounds are connected via weak bonds (hydrogen, van der Waals, or ionic bonds); and Class II, involving strong bonds between the components (covalent or ionic-covalent). The organic components can be introduced in the hybrid material as network modifiers (e.g., triethoxydibenzoylmethane, polyvinylmetoxysilane, diphenylphosphine-propyltrimethoxysilane, metalloporphyrin, and porphyrin derivatives) or network formers (e.g., pyrrole derivatives, trimethoxysylilpropylacrylate). The latter can be introduced in two ways: (1) using already presynthesized functional monomers linked with the inorganic component through embedding using a proper solvent (Class I) or chemical grafting (Class II); or (2) in situ generation via photo- or thermal-polyaddition, most frequently in the presence of an initiator, chemically, or electrochemically induced polymerization, polycondensation, atomic transfer radical polymerization, and others [59]. A plethora of methods can be employed to synthesize organic
inorganic materials, all of them based on the closely mixing organic molecules or networks (e.g., organic polymers) with inorganic compounds, such as graphene, CNTs, metal oxides, or metal
oxo polymers. A perfect example is the coronary stent, which is used in coronary heart disease and is placed in the arteries to keep them open. The stent is made from a platinum chromium alloy (PtCr) embedded in two layers of drug
polymer coating. The inner layer consists of a primer-polymer layer which helps in the attachment of the bare metal and the outer part consisting of the drug matrix. The latter contains a semicrystalline

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copolymer, for example poly(vinylidene fluoride-co-hexafluoropropylene) (PVDF-HFP) that contains 40-O-(2-hydroxyethyl)-rapamycin (everolimus), an active pharmaceutical ingredient with immunosuppressant activity. The applications of some of the most common composite materials in sensor development are presented next.

6.3.1 Carbon
Organic Hybrid Materials Among carbon materials, CNTs and graphene are the most widely used in the development of biosensors. Two-dimensional graphene has the shape of a honeycomb lattice conferred by the arrangement of a sheet of sp2-bonded carbon atoms lattice. In addition to having remarkable mechanical strength, graphene exhibits good biocompatibility and excellent electrical conductivity [60]. Moreover, graphene properties can be easily tuned to the desired application, for example it can be converted to graphene oxide (GO). GO exhibit many hydrophilic groups (e.g., hydroxyl, epoxy, carbonyl, carboxyl) increasing aqueous dispersability, and the possibility of functionalization with other molecules [61]. CNTs consist of rolled graphene sheets and since their discovery [62], CNTs attracted enormous attention in the development of biosensors due to their many advantages, such as good thermal and electrical conductivity, great mechanical flexibility, and excellent stability. Due to their large active surface and high stability, CNTs serve as great support for the immobilization of organic molecules [63]. Carbon materials have been widely employed in conjunction with polymeric platforms, offering a great design variability and versatility to a variety of sensors for the analysis of different analytes. Either prepared in situ directly on the transducer surface or ex situ by direct blending, many of the polymers used

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for biosensor development are conductive polymers [64,65]. With a multitude of advantageous chemical and physical properties that can reversibly or irreversibly change under the influence of external stimuli, polymers are the most commonly used materials in sensor fabrication [66]. Such nanocomposites can be used directly to increase sensitivity or selectivity toward the detection of certain analytes or to efficiently immobilize various biorecognition elements, for example, antibodies and aptamers. For instance, Jijie et al. recently elaborated an immunosensor based on GO/poliethyleneimine platform coupled with electrochemical detection system that was applied to detect Escherichia coli. With an interesting functionalization approach, namely electrophoretic depositing, gold electrodes were coated with thin layers of nanocomposite material. Reduced GO increased the electroactive area as well as the electrocatalytic properties, while polyethyleneimine provided a high number of amino groups to covalently link anti
E. coli antibodies. Based on hindering of the oxidation of a redox probe in DPV measurements upon antigen binding to form an immune complex the LOD of 10 cfu/mL was obtained in the linear range 101
104 cfu/mL [67]. Van Chuc et al., reported an immunosensor using polyaniline
graphene hybrid composite for sensitive detection of atrazine, an

environmental pollutant. The graphene film was prepared by thermal chemical vapor deposition and then attached onto PANIcovered microelectrodes. Providing signal amplification and increasing electron transfer properties the platform led to a low LOD of 43 3 10212 g/L [68]. A PANI
graphene composite was also employed by Barton et al. as immobilization platforms for antibodies. In this case, the fabrication approach was simpler as PANI was readily electrodeposited onto graphene
screen printed electrodes. The sensor exhibits high selectivity and good sensitivity for the detection of low molecular weight, poorly water soluble endocrine disruptor chemicals, with low detection limits (pM) [69]. Benvidi et al. combined poly(L-glutamic acid) and MWCNTs to provide an increased number of carboxyl groups for immobilization of aptamers toward tetracycline. MWCNTs were dropcasted onto glassy carbon electrodes (GCE) followed by potentiodynamic electrodeposition of the polymer. The sensor exhibited remarkable features with an LOD of 3.7 3 10217 M and excellent selectivity toward tetracycline [70]. A nanoplatform combining graphene, cyclodextrin, and tyrosinase was achieved by layerby-layer deposition and optimized by using square wave voltammetry (Fig. 6.1) [71]. The sensor was used to determine dopamine in pharmaceutical products, human serum and urine.

FIGURE 6.1 The development of the graphene/β-cyclodextrin biosensor with tyrosinase [71].

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A pyrrole monomer functionalized with cyclodextrin was synthesized by Cosnier et al. [72] which could be electropolymerized to cover electrodes with composite materials. Following the entrapment of tyrosinase, the sensor was applied to detect catechol and dopamine with very good performances. Therefore, the electropolymerization method represents a promising tool to fabricate miniaturized sensors suitable for in vivo determinations of biochemicals. Recently aptamers and antibodies have been widely replaced with synthetic tailor-made receptors, for example, molecularly imprinted polymers (MIPs), due to their high stability in different environmental conditions, low cost, ease of synthesis, and high selectivity. The most simple and common approach to develop electrochemical sensors based on MIP is direct electropolymerization onto electrodes, which results in a thin adherent layer. For example, simultaneous detection of uric acid and tyrosine as diagnostic markers for various diseases has been achieved by Zheng et al. with a novel MIP
GO nanocomposite. The hybrid material was prepared by a simple potentiodynamic electrodeposition of poly(2-amino-5-merapto1,3,4-thiadiazole) onto MWCNTs drop-casted on GCE providing a higher surface for analyte binding. The sensor exhibits an LOD of 3.2 nM and 46 nM for uric acid and tyrosine, respectively, and was applied for the detection of these analytes in serum and urine [73]. Yin et al. exploited the self-polymerization property of dopamine in water to construct an MIP sensor for the detection of sunset yellow on MWCNT-modified GCE. Combining the excellent electrocatalytic activity provided by MWCNTs with the highly selective cavities in the superficial polymer layer the sensor exhibited a detection limit of 1.4 nM, great selectivity, good stability, and reproducibility [74]. Even though most applications of MIPs include small molecules, imprinting large

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molecules, such as proteins or cell is also possible. Huang et al. [75] reported on the selective sensing of glycoprotein ovalbumin. The research group used boronic acid functionalized GO to covalently link glycoproteins via ester linkage. Surface imprinting was obtained by polymerization using the sol-gel method, employing organic silanes and ovalbumin as the template. The sensor outperformed previously reported sensors for ovalbumin detecting concentrations as low as 2 3 10211 mg/mL and included biological fluids applications. A remarkably low detection limit was also obtained by Liu et al. by combining MIPs with graphene for the detection of testosterone by EIS [76], demonstrating the synergistic effects of carbon nanomaterials and biomimetic polymers. A carbon
polymer hybrid platform can be successfully employed as a sensing platform without requiring the integration of a biorecognition element. Hosu et al., fabricated a simple sensor combining the advantages of electroactive polymers, for example, poly(methylene blue) and CNTs. Nanostructured polymer films were prepared in deep eutectic solvents by cyclic voltammetry either on top or beneath the CNT layer on GC electrodes. The resulting hybrid material demonstrated superior electronic conductivity and was applied for the analysis of ascorbic acid and paracetamol reaching LODs of 1.7 and 1.6 μM, respectively (Fig. 6.2) [77].

6.3.2 Metal Nanoparticles: Organic Composites and Metal
Organic Frameworks Novel metal nanomaterials have been fabricated recently with various sizes, shapes, compositions, and structures since all these properties are closely related to their physicochemical properties. Having enhanced surface areas, excellent electrical conductivity, and

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FIGURE 6.2 Hybrid nanocomposite based on poly(methylene blue) electropolymerized in choline chloride deep eutectic solvent (PMBDES) and CNTs [77].

high-binding capacity for organic compounds and biomolecules, metal nanomaterials are widely used to fabricate high-performance biosensors. The surface plasmon phenomena property of metal particles has been widely employed to develop optical sensors [78]. Metal particles can be chemically synthesized and then incorporated into composite materials, or, in the case of electrochemical sensors they can be readily electrodeposited directly onto the electrode surface. Combining nanostructures made up from different materials can bring about great performances of biosensors. Doping polypyrrole with AuNPs led to an MIP layer with increased sensitivity toward the detection of 3nitrotyrosine, an oxidative stress biomarker, by facilitating the charge transfer process and increasing the number of active binding sites [79]. Another biomimetic sensor employing AuNPs was fabricated by Tertis et al. for the

sensitive determination of dopamine. AuNPs were electrogenerated simultaneous with the polymer polythioaniline in a single step by cyclic voltammetry. The sensor allowed the direct detection of low dopamine concentrations with high selectivity and without the need of a redox mediator [80]. The same group elaborated an AuNPs
polypyrrole NP composite as a modifier for GCE that allowed the selective detection of another neuromediator, serotonin. Combining the two materials, an increased surface area together with an enhanced catalytic effect was obtained leading to an analytical response 320-fold higher than on nonmodified electrodes [81]. The group also employed an AuNPs
polypyrrole composite as an immobilization platform for IL6 aptamers to develop a label-free aptasensor for the sensitive determination of IL6. The presence of the conductive polypyrrole and the gold provided good conditions to link the

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thiolated aptamer. The impedimetric response showed a wide linear range from 1 pg/mL to 15 mg/mL with an LOD of 0.33 pg/mL [82]. An impedimetric sensor was also designed by Fusco et al., who developed an AuNPsfunctionalized PANI and MWCNT composite integrated in an immunosensor detection of 2,4-dichlorophenoxyacetic acid. The composite material was used to covalently attach antibodies, leading to an LOD of 0.3 ppb [83]. Another hybrid material consisting of graphene, gold, and polypyrrole for sensing applications was reported by Cernat et al. The sensor was applied for the fast detection of pyoverdine. The hybrid material helps to improve the electron transfer allowing the detection of pyoverdine with an LOD of 0.33 μM [84]. Although gold is one of the most widely used metals in sensors using composite materials, other metals are also gaining increasing interest. For example, molybdenum disulfide (MoS2) is a layered transition metal dichalcogenide, but the drawback of MoS2 is its low-electronic conductivity, thus, it is usually combined with more conductive materials. Yang et al. combined MoS2 with PANI by oxidizing aniline on the MoS2, leading to a platform with a high electrocatalytic surface area, the resulting composite material giving better responses for the electrochemical detection of DNA molecules [85]. Metal
organic frameworks (MOFs) (coordination polymers, coordination networks) consist of metal ions
clusters bridged by organic ligands. MOFs proved remarkable potential in biosensing due to their structural flexibility, porosity, controllable synthesis, tailorable pore size or walls, and exposed active sites [86,87]. Due to the fact that they are biodegradable, biocampatible, and have low cytotoxicity, MOFs are particularly advantageous for biological applications [87]. MOFs are also advantageous for developing fluorescent sensors, since the metal and ligands (e.g., lanthanides, aromatic, or conjugated π moieties) can generate fluorescence that can be tuned by the

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interaction among the components [88]. Moreover they can be loaded with various guest molecules. All these advantages promote the use of MOFs as quenchers in DNA or RNA sensing, small biomolecule sensing, and of MOF
enzyme composites in enzyme-activity sensing and label-free biosensing. In optical biosensing applications, signals can be achieved by using MOFs as fluorescence quenchers toward the fluorophores of analytes. It is also possible to prepare MOFs with fluorescent properties that change depending on the environment or guest species [87]. For electrochemical biosensing applications, MOF composites (e.g., carbon-based materials, metal NPs, heteropolyacids, enzymes) with improved electrocatalytic activities are often synthesized and applied for different analytes. Two approaches may be undertaken for their synthesis, namely in situ one-pot methods involving combining precursors of both materials, and multistep synthesis in which precursors of MOFs are mixed with presynthesized functional materials [89,90]. MOFs exhibit low-electronic conductivity, electroreactivity, and poor water solubility. The assembly with conductive materials, for example, CNTs and graphene, improves MOFs’ electrical conductivity and mechanical strength, making them attractive for the development of electrochemical biosensors with improved characteristics [86]. For example, Yadav et al. demonstrated the advantage of incorporating silver in zinccontaining MOF-5 to enhance electrochemical oxidation of accumulated nitrophenols (pollutants) at GCE. Analyte accumulation on the hybrid materials was demonstrated as Langmuir binding constants of 15
40 3 103 M21, and current enhancement with more than an order of magnitude were achieved [91]. Metal-modified MOFs have also been reported by Shu et al. who decorated Ni/NiO/MOFs with AuNPs and employed the hybrid material in a nonenzymatic glucose sensor. Ni/NiO/MOFs have

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been prepared by calcinating Ni
MOF and linking AuNPs through electrostatic adsorption. Amperometric measurements on modified GCE showed high sensitivity with an LOD of 0.1 μM [92]. Wang et al. combined AuNPs with Cu
MOFs to construct a paperbased biosensor for microRNA detection. The MOF
hybrid material was used to label DNA strands as a signal amplification strategy. The prepared DNA strand/AuNPs/Cu-MOF is immobilized onto electrodes via chain hybridization and the released target was available for many cycles leading to increased signals. A remarkably low LOD of 0.35 fM for the detection of miRNA-155 was achieved with this strategy, demonstrating the use of MOF
hybrid materials for signal amplification in electrochemical biosensors [93]. As noted, graphene is a carbon material that possesses excellent electrical conductivity and exhibits numerous functionalities (hydroxyl, epoxy, carboxylic, oxygen). Its electrocatalytic activity has been explored in MIP-based sensors. A sandwich immunosensor for galectin was constructed by Tang et al. [94] based on N-doped graphene nanoribbons/AuNPs/ MOFs signal amplification system. The target was captured between a primary antibody immobilized on electrodes modified with this MOF hybrid and a secondary antibody labeled with rod-like Au-Pt-methylene blue. Following DPV measurements, an LOD of 33.33 fg/mL was achieved. Another signal amplification strategy using MOFs was reported by Yu et al. [95]. Sensitive determination of Pb21 was achieved based on promotion of nucleic acid cleavage by the Pb21DNAzyme in which Pt-Pt@Fe-MOFs were used as tags to trigger the signal. The designed sensor exhibits an LOD of 2 pM in the range of concentrations from 0.005 to 1000 nM. MOFs have been also employed in synthesizing artificial receptors based on the molecular imprinting technique. Guo et al. [96] developed several MOF-based MIP sensors for

the detection of various analytes. The general protocol is presented for TNT imprinted polymers (Fig. 6.3). The imprinted microporous
MOF was deposited directly onto the surface of gold electrodes via electrodeposition of poly(p-aminothiophenol)-linked gold nanoparticles using trinitrotoluene as the template. After removal of the template from the resulting polymeric composite, specific recognition cavities for TNT were achieved. The rebinding of the target was monitored by LSV measurements using a redox probe [96]. This protocol has been optimized and adapted for the indirect determination of other compounds such as tetracycline [97], gemcitabine [98], estradiol [99], and glyphosate [100]. All sensors exhibit good responses in real complex matrices with wide linear ranges (nM-fM) and low LODs in the fM range (Table 6.1).

6.3.3 Magnetic Materials Widely employed for in vitro biomedical diagnostics and biotechnology, magnetic nanoparticles (MNPs) have also gained increasing interest in recent years in the development of biosensors due to their unique properties such as particular physicochemical properties, large surface area, high-mass transference, biocompatibility, and facile production. The nature of these magnetic materials (e.g., magnetite, iron, nickel, cobalt), their size, core, and coating have a strong influence on their toxicity and biocompatibility. The size and shape also contribute to the magnetic properties of the nanoparticles. Magnetite (Fe3O4) and maghemite (γ-Fe2O3) are more widely used, being more biocompatible (less toxic) and less prone to oxidation [101]. Many methods can be used for the synthesis of magnetic particles depending on the intended application, including physical methods, wet chemical methods, and microbial

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FIGURE 6.3 General protocol for developing metal
organic frameworks (MOF)
based molecularly imprinted polymers (MIPs) for electrochemical detection of TNT [96].

TABLE 6.1 Metal
Organic Framework (MOF)
Based Molecularly Imprinted Polymer (MIP) Sensor Performances for the Electrochemical Detection of Various Analytes Template

MIP Slope

MIP R2

NIP Slope

NIP R2

Linear Range

LOD (fM)

LOQ (fM)

Refs.

TNT

0.835

0.976

0.182

0.997

4.4 fM
44 nM

0.044

4.4

[96]

Tetracycline

0.022

0.982

0.004

0.885

0.2 fM
2.2 nM

0.224

0.2

[97]

Gemcitabine

0.133

0.980

0.022

0.915

3.8 fM
38 nM

0.003

3.8

[98]

17β-Estradiol

0.403

0.972

0.016

0.650

3.6 fM
3.6 nM

0.036

3.6

[99]

5.9

[100]

Glyphosate

0.018

0.971

0.005

0.802

5.9 fM
5.9 nM

26

5.9 3 10

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methods. New methods have been elaborated to obtain magnetic particles of different sizes and shapes [102] [120]. In order to prevent irreversible aggregation and to promote dissociation (and, thus, higher stability) the surface of MNPs is usually coated with surfactants or polymers that prevent cluster growth after nucleation and offer attractive forces against the particles or with lipid-like compounds (e.g., liposomes) [103]. MNPs possess superparamagnetic properties below 50 nm size and exhibit their best performances at 10
20 nm, which makes them a suitable choice in magnetic fields when a fast response is needed [104]. MNPs can be easily integrated into various transducers with the advantage of being able to attract and bind analytes from the sample in a very simple manner by an external magnetic field [103] or they can be easily used as labels. The advantages of a biosensing strategy involving MNPs include enhanced sensitivity, lower LODs, improved stability, less noise, faster analyses, and reduced influence of the matrix effect due to improved washing [105]. Magnetic/polymer materials have been commonly used to fabricate many different sensors over the past few years and several relevant examples with different applications are described further. An interesting approach was undertaken by Eftimie Totu et al. who developed a sensor based on magnetic nanoinclusions for the early diagnosis of periodontal disease based on monitoring the alterations in inorganic ions (e.g., sodium). For this purpose the group developed several sodium selective membranes with magnetic nanoinclusions which allowed Na detection by electrochemical methods in the range of 3.1 3 1025 and 1021 mol/ dm. Due to the very small dimensions the designed sensors can be applied to measure ions directly from the gingival fluid without the need to collect a high amount of fluid for analysis [106].

Zaibudeen et al. [107] developed a magnetic nanofluid
based nonenzymatic sensor for the sensitive and selective detection of urea. The nanofluid used as optical probe is an oil-inwater magnetic nanoemulsion containing superparamagnetic iron oxide nanoparticles of sizes B10 nm which, under a constant magnetic field, form a one-dimensional Bragg diffraction grating with a fixed interparticle spacing. The probe showed a large wavelength shift in the visible wavelength range of 880
600 nm in the presence of urea, due to complexation of urea with the functional moieties that dramatically changes the electrostatic repulsion between emulsion droplets or conformation of adsorbed polymer (Fig. 6.4). The functional moieties used here were sodium dodecyl sulfate, poly acrylic acid, and poly (ethyleneoxide)-block-poly(propyleneoxide)block-poly(ethyleneoxide) copolymer. The sensor detected urea in the gram per liter range of concentrations [107]. Magnetic nanoparticles functionalized with palladium nanoparticles on a polymeric matrix was applied due to its catalytic effects in a novel electrochemical sensor. The Pd@PCA-bPEG-Fe3O4 was immobilized on the surface of a GCE for the electrochemical reduction of hydrogen peroxide showing good electrocatalytic activity toward the reduction of H2O2 in the neutral phosphate buffer solution. A linear dependency of the current with the H2O2 concentration was achieved from 0.1 to 50.0 μM by DPV [108]. Magnetic silica particles were modified with copolymers containing guanidine using the sol
gel technique and employed in a bioluminescent sensor for bilirubin detection by Timin et al. The surface of these materials was functionalized by a protein extracted from Japanese eel (UnaG) which exhibits fluorescence upon bilirubin binding. The detection was achieved via the bright fluorescence exhibited by the silica-polymer-UnaG material in an aqueous bilirubin solution [109].

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FIGURE 6.4 Working principle of the magnetic nanofluid sensor for urea optical detection [107].

Magnetic core MIP particles have been widely employed for the detection of a wide range of analytes allowing easy capture and preconcentration of the analytes, thus leading to low detection limits and high selectivity. For example Ruiz-Cordova et al. developed a magnetosensor for the detection of the allergenic 1-chloro-2,4-dinitrobenzene [110]. The magnetic nanoparticles were synthesized via the coprecipitation method and were then encapsulated in a hydrophobic polymeric matrix via the miniemulsion method. The resulting particles were covered with an MIP layer via precipitation polymerization in a solution containing the

analyte, methylene diphenyl diisocyanate-4, bisphenol A, and phloroglucinol. The particles were added to the sample to be analyzed allowing to load a high amount of target analyte and then were captured on a working electrode with an integrated magnet. The analyte was directly detected by DPV in concentrations as low as 6 μM with good reproducibility (RSD 2.7%). Zamora Galvez et al. [111] prepared magnetic MIPs for the detection of tributyltin water pollutant by functionalizing Fe3O4 NPs prepared from iron chloride with amino propyl triethoxysilane, which allowed controlled

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polymerization of ethyleneglycol dimethacrylate around tributyltin hydride template molecules. Afterwards the prepared MIP Fe3O4 NPs were incubated with the sample to bind the analyte and were placed on screen-printed electrodes for electrochemical measurements and kept in place on the working electrode with a magnet placed underneath. A low LOD of 5 pM was obtained by impedimetric measurements, which was suitable for real sample applications being below the recommended values [111]. Wu et al. prepared a magnetic fluorescent probe of CdTe QDs/nanoFe3O4@MIPs for the optical detection of malachite green in fish samples. The composite material included CdTe QDs and Fe3O4 nanoparticles which were covered with MIP layers. The method used for the preparation was reverse microemulsion. The fluorescence of the prepared material was quenched by malachite green from the samples allowing to achieve a detection limit of 0.014 μmol/L due to efficient magnetic separation and enrichment with the analyte. The strategy was successfully applied for the detection of malachite green in fish samples with a good average recovery of 105.2% [112]. The core
shell method was used by Uzuriaga-Sanchez et al. to synthesize magnetic MIP particles for the selective detection of biotin in milk-based food samples. Molecular modeling was performed to choose the best monomer with high-binding affinity for the target (i.e., acrylic acid). Radical polymerization was employed involving azo-bisisobutyronitrile as a radical initiator and ethyleneglycol dimethacrylate as a cross-linking agent. The resulting magnetic particles were used for HPLC analysis of biotin in milk samples with good performances [113].

6.3.4 Functionalized Clays and Silica The production of synthetic materials such as polymers, mesoporous silica, and MOFs generates a series of problems (issues or

challenges) related to pollution and waste management. Starting materials or substrates with important capabilities in the immobilization or entrapment of biomolecules for sensor design are abundant in nature. Due to their structures and chemical composition as well as chemical accessibility of their reactive sites, clay minerals are suitable for this purpose. In order to extend their applications, they could be easily modified with organic compounds. Among the clay minerals, the smectite family is largely used in the nanocomposite field, but more recently kaolinite, which is abundant in nature [114], has been exploited in functional chemistry. Even that it was found to be difficult to modify, the aluminol functions confined in the interlayer spaces if made accessible could react with organic moieties to produce functional organic
inorganic nanohybrid materials, such as polymer-based nanocomposites [115]. Several examples of nanohybrids used for biosensors are described by Tonle et al. who succeeded to modify kaolinite by grafting with 3-aminopropyltriethoxysilane or with imidazolium and to use these new materials in sensors for heavy metal detection [116,117]. The individual features of organic and inorganic units that interact at the molecular scale assign to organic
inorganic hybrids a broad range of properties, which could be relevant for many applications [60]. Biohybrids, which represent a growing field of research of advanced functional materials design, can be employed in these organic
inorganic hybrids. These biohybrids, called bioorganoclays, are prepared by intercalation of biomolecules (lipids or proteins) [118
120], their main application being as fillers in the preparation of bionanocomposites or as biointerfaces for the adsorption of biological species. The use of biofillers in polymer matrix improves the biocompatible character of the material and brings new features, such as barrier properties in reinforced bioplastics. Some large molecules, such as polypeptides and proteins, could also intercalate into montmorillonite and other

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smectite clay minerals, thereby producing biohybrid materials. Montmorillonite, being a phyllosilicate, is characterized by a colloidal particle size, high specific surface area, and large cation exchange capacity. Each silicate layer is formed by the repetition of a central octahedral alumina sheet sandwiched by two tetrahedral silica sheets. The preparation of zein
montmorillonite biohybrids was reported by Alcantara et al. focusing on the control of solubilized zein for an effective intercalation of the protein into Na-montmorillonite. Zein
Na-montmorillonite biohybrids were tested as strengthening fillers of other biopolymer matrices for developing “fully” ecofriendly bioplastics [121]. Silica is another interesting material which is highly modifiable, with pore sizes, pore volumes, pore organization, and pore wall functionalities [122,123] being easily tunable. It represents another starting point for developing hybrid nanomaterials. Interest in the studies of silica particles, which bear functional groups, lies in achieving control over its stability to hydrolysis and especially over its availability and chemical reactivity. The silica matrix shows mechanical stability due to the organic bridging moieties present in the organosilica host. The nitrogen atoms contained by moieties of organosilanes could reversibly be protonated contributing to enhance the ionic conductivity of the proton conducting ionogels. When the silica matrix is combined with ionic liquids like 1-methyl-3-(3sulfobutyl) imidazolium p-toluenesulfonate, which has a high ionic conductivity, the newly obtained nanohybrid material yields stable organosilica matrix materials. The combination of ionic liquid with a suitable organosilica host gives access to robust and flexible ionogels that may be interesting as membrane materials in sensors, batteries, or fuel cells [124]. As previously mentioned, the mesoporous-silica hybrid materials have had many applications in the field

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of electrochemical biosensors, immunosensors or genosensors, and electrochemiluminescence (ECL) platforms, and are very well presented in a review by Walcarius [52]. Thus core
shell nanoparticles were reported with the silica core used as a support for the conductive polymer (SiO2@polyaniline) [125] and mesoporoussilica micro- and nanospheres are increasingly employed for the development of biosensors based on antibodies, aptamers, nucleic acids, or MIPs. Walcarius et al. [126] reported the use of surfactant templating and electrochemically assisted self-assembly for the achievement of highly ordered silica films with mesopores perpendicular to the electrode surface. After the surfactant molecules removal by calcination or solvent extraction, uniform thin membranes formed of 2
3 nm in diameter hexagonally-packed channels were obtained [127]. Besides having excellent molecular sieving properties (nonaferrocenyl dendrimer was able to pass a 2.9-nm pore, but was totally excluded by 2.0 nm channel diameter [128]), the negatively charged walls of the silica membranes also showed charge permselectivity, anions being rejected and cations being accumulated by electrostatic interactions, while small neutral species freely diffused across mesochannels [129].

6.3.5 Ionic Liquid Hybrid Materials Over the past decade, there has been an increasing emphasis on the topic of “green” chemistry because of its minimized adverse environmental effect. Enormous attention was focused on the use of ionic liquids (ILs) for applications such as catalysis, biotechnology, electrochemistry process technology, and analytics [130], which is highlighted in the number of papers that has grown exponentially since then. Nevertheless, their use as “green” solvents is often challenged. In this sense, deep

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eutectic solvents (DESs) are fostered as greener alternatives to ILs because of their similar characteristics, but DESs are cheaper and easier to produce (lower cost of the raw materials), less toxic, and often biodegradable [131,132]. Thus DESs are prepared by mixing a salt and a hydrogen bond donor, hence, the hydrogen bonds with the anion of the salt, which could be organic or inorganic [133]. DESs are defined as a mixture of compounds that have a melting point significantly lower than that of either individual component becoming liquids at room temperature [132]. The most common DESs are based on choline chloride, carboxylic acids, and other hydrogen bond donors (e.g., urea, citric acid, succinic acid, glycerol) [134]. ILs and DESs media have been applied for the electrochemical synthesis and deposition of conducting polymers for hybrid sensing systems [77,133,135
139]. Zhu et al. [139] developed an impedimetric DNA sensor based on ionic liquid-carbon paste electrode entrapped by a rod-like Bi2S3/PANI nanocomposite. The ionic liquid-carbon paste electrode exhibited excellent electronic conductivity (detection range from 10215 to 10211 M and LOD of 4.37 3 10216 M), large surface area, and good biocompatibility for DNA immobilization and hybridization detection. Another PANI-based electrochemical platform was synthesized in a 1:2 mixture of choline chloride/1,2-ethanediol. The film exhibited nanoparticulate morphology, high reversibility, and excellent conductivities [136]. Nanoparticle-shaped film formation also occurred in the case of poly(methylene blue) synthesized in choline chloride
ethylene glycol eutectic mixtures and it was proven that the DES films exhibited a 15fold higher sensitivity for ascorbate detection compared to those prepared in aqueous media [135]. Other electrochemical sensors were developed by Pratish et al. by electropolymerizing EDOT in different DES media

with application as biomarker sensors [137,138]. Due to the ongoing research in this hot topic, future perspectives will be certainly envisaged on the development of greenchemistry sensing approaches for a wide range of molecules over the next few years.

6.4 ADVANCED ORGANIC HYBRID MATERIALS Another hot topic of research which developed rapidly in the 1990s is nanostructured materials. In this regard, ordered porous polymer materials attracted attention from the fields of physics, materials, and chemistry due to their advantages like slight weight, ordered row of pores, high specific surface area, great adsorption capacity, and uniform and tunable pore sizes. These materials have great prospects in biomedicine (biofilms, bioreactors, etc.) and separation techniques, allowing the growth of templates with specific shapes. For these reasons, order polymers have aroused much interest in the field of polymer science and are being intensively used in targeted drug delivery and sensing [140]. Currently, ordered polymers are classified into two types, namely membranes and interconnected repeating cells (channels). The inherent properties and conformational movements of the polymers combined with organic molecules and macromolecules may lead to molecular self-assemblies or nanostructures with potential application in different area. Therefore in nanotechnology, molecular self-assemblies are products of a “bottom
up” approach. An emerging application of those molecular self-assemblies is in the field of drug delivery as a result of the noncovalent interactions between their building blocks. This is due to their capabilities to degrade back into individual monomers easily broken down by the in vivo environment [141].

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The main advantage in using nanostructured ordered polymers instead of traditional ones consist in uniform pores with large aspect ratio compared with random pore size distribution. Possessing also excellent adsorption capacity, they are used as building blocks in sensor development. Sitti et al. [142] ordered polymeric membranes for the design of a nanosensor containing a pair of interdigitated electrodes deposited on the surface of polymeric membranes. High sensitivity due to the water adsorption inside the nanopores was obtained after registering the resistance and capacitance variation between the electrodes at different humidity levels. The main application of these sensors is focused on in situ water leakage detection. Another example in which order polymers were used in sensor design was described by Stucky et al. [143]. They demonstrated that ordered polymer materials can construct an effective platform for optical chemistry sensor. Cartwright et al. [144] prepared multifunctional holographic polymers to be used as a general template for the encapsulation of recognition elements and developing an O2 sensor. Encapsulating the fluorophore ruthenium (II) within these nanostructured materials displayed a full detection range from 0% to 100% oxygen concentration, exhibiting stable, repeatable, and reversible performances with a fastresponse time. Conducting polymer-based nanohybrids proved their role as innovative transducers for high-performance chemical and biological sensing devices. Synthetic strategies of the conducting polymer-based nanohybrids are classified into four groups: (1) impregnation, followed by reduction; (2) concurrent redox reactions; (3) electrochemical deposition; and (4) seeding approach [145]. Sometimes the electrical and optical properties of conducting polymers are similar to those of inorganic semiconductors or metals.

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For this reason, conducting polymers have been widely used to design chemical and biological sensors. Another important property that makes them suitable for various sensor transducers consists in their high sensitivity to their surrounding environment. If we compare the conducting polymers with other materials such as metals or inorganic materials, the polymers are easily synthesized and processed as they possess chemical and structural diversity, low weight, and high flexibility. Due to properties like good electrical conductivity, affinity towards biomolecules and biocomptibility, different strategies have been developed for the direct and indirect detection of various biomolecules, such as hydrogen peroxide, glucose, dopamine, DNA, and proteins at modified electrodes with conducting polymers [146]. The most used organic conductive polymers include poly(acetylene)s, poly(pyrrole)s, poly (thiophene)s, poly(terthiophene)s, poly(aniline) s, poly(fluorine)s, poly(3-alkylthiophene)s, poly(tetrathiafulvalene)s, poly(naphthalene)s, poly(p-phenylene sulfide), poly(p-phenylene vinylene)s, and others [147,148]. Conducting polymers are known for their compatibility with biological molecules in neutral aqueous solutions and, in some cases, the polymer itself can be modified to bind biomolecules to a biosensor [149]. Another advantage of conducting polymers is that the electrochemical synthesis allows direct deposition of a polymer film on the electrode substrate followed by biomolecule immobilization. It also allowed the deposition over defined areas of the electrodes. It is, thus, possible to control the spatial distribution of the immobilized biomolecules and film thickness. Conducting polymers can act as electron promoters. The use of conducting polymers doped with other organic compounds and used for biosensor design was extensively presented in the review by Shim et al. [150]. Polypyrrole (PPy) is one of the conducting polymers that has been extensively studied

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and used due to its advantages in biosensor design. PPy has interesting properties such as electrical versatility from insulator to nearly metallic, easy functionalization, high solubility in aqueous environment, biocompatibility, low electropolymerization potential, as well as stability under ambient conditions, making this material a promising tool for electrochemical sensor design [151
153]. A nice example of PPy nanotubes functionalized with carboxylic groups was described by Jang et al. and the obtained conducting polymer nanohybrids are being used as substrates for molecular probes and DNA carriers [154]. Functionalized PPy nanotubes decorated with a photoluminescent molecule, pyreneacetic acid, and silica nanoparticles employed as the linker was used to attach single-stranded DNA (ssDNA) with a terminal amino group. The silica nanoparticles with surface amino groups were first attached to the PPy nanotubes and then pyreneacetic acid was linked with the silica nanoparticles. All these modifications lead to an enhanced surface area of the sensors, indicated by higher sensitive responses compared to the unmodified PPy nanotubes. The PPy nanotube hybrids were characterized with IR spectroscopy, XPS, and electron microscopy. Another frequently used polymer is polyaniline (PANI), which has many attractive processing properties [155,156]. Because of its rich chemistry, PANI is one of the most studied conducting polymers. It was used to develop organic
organic materials and their application in sensing was described by Tian et al. who reported stable multilayer films using PANI/mercaptosuccinic-acid-capped gold (MSAG) nanoparticle hybrids by the layer-bylayer method [157]. The PANI layer was easily stimulated and its electroactivity was shifted to neutral pH by the use of modified MSAG under zero polyelectrolyte conditions. Moreover, a PANI hybrid layer was functionalized in order to introduce the aminoterminated DNA catcher probes, which can be used for monitoring their hybridization

with different DNA target strands by electrochemical methods. The electrochemical impedance spectroscopy (EIS) was used to investigate the surface-charge density of the PANI hybrids and the researchers observed an increase from the hybridization process with target DNA, and an increase of the electrostatic repulsions. Polyethyleneimine (PEI) is another polymer that by being embedded with cyclodextrines was successfully used in sensor design. PEI is a cationic polymer and has been used over the past decade for the entrapment of several molecules in various biosensors. The numerous advantages of this polymer are used in biosensor design, the main benefits being the capacity of retaining the biomolecule at the surface of the electrode without stressing it with any supplementary electropolymerization process, and its partial solubility in water [158
160]. The simultaneous detection of ascorbic and uric acids from urine samples without any previous separation was described by Fritea et al. [161] and was achieved with a sensor based on β-cyclodextrin entrapped in a polymeric film of PEI. Beyond doubt, the new and improved chemical and physical properties of conducting polymer nanohybrids will be further exploited for designing high-performance sensors. Due to the advances in nanohybrid syntheses, it is expected that the use of organic nanohybrids based on conducting polymers will also evolve, giving to the sensor technology the possibility to improve their sensitivity and selectivity.

6.5 OPTICAL MULTIFUNCTIONAL ADVANCED HYBRID MATERIALS Optical detection biosensors are the most diverse class of biosensors because there are many optical processes to be followed, such as absorption, fluorescence, phosphorescence,

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refraction, dispersion, and others. [162]. Moreover, different properties can be triggered during the spectroscopic measurements such as polarization, energy, amplitude, decay time, and phase. Usually, the detection method of optical biosensors is based on processes like absorption, fluorescence, or light-scattering events. One major advantage of optical biosensors occurs from their nonelectrical working systems. Thus in vivo applications can be monitored allowing simultaneous multiple analyte detection by scanning the spectra at different wavelengths [163]. A large variety of optical methods have been used in biosensors; however, fluorescence spectroscopy, surface plasmon resonance (SPR), and surfaceenhanced Raman spectroscopy are the most commonly used. Recently, optical ring resonators and photonic crystals have been under investigation, representing the emerging optical sensing technologies [162]. In the past few years, numerous papers reporting the development of new optical sensing devices have been mentioned in significant fields including environmental monitoring, food safety, and clinical analysis [164,165]. One of the most popular protocols for optical bioanalysis is still attributed to immunosensing strategies due to the advantages of their nondestructive operation mode and harmless visible radiation that reveals the antibody
antigen event as well as providing rapid signal generation and reading [166]. Different techniques including chemiluminescence (CL), ECL, fluorescence, surface plasmon resonance (SPR), and SERS [164] have been reported for optical sensing and would be further detailed. Some challenges could occur due to the harsh environmental conditions into which nanoparticle labels are created, which often cause the inactivation of biological targets. Two main strategies have been explored and developed to deal with these challenges: (1) hybrid composite surface functionalization; and (2) improved bioimmobilization procedures (physisorption, affinity interaction, entrapment in sol
gel

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matrices and covalent conjugation) [163]. Some examples of optical biosensors based on hybrid materials are presented in Sections 6.5.1
6.5.3.

6.5.1 Chemiluminescent and Electrochemiluminescent Materials CL methods have become very popular in clinical and biomedical fields due to their advantages such as producing nonradioactive wastes and having low detection limits in wide dynamic ranges [167
169]. Optical detection based on CL is an ideal method for miniaturization and POC biosensor development because of its inherent sensitivity and simplicity [170]. Moreover, ECL combines luminescent and electrochemical techniques. Thus the luminescence event occurs as a result of an electrochemical reaction in which the excited molecules generate photons during their relaxation [171]. CL immunoassays are able to determine the analyte concentrations by the intensity of the luminescence emitted by the immunoreaction event [172]. One of the most used CL-based immunosensors employ enzymes as labels for signal amplification (e.g., horseradish peroxidase for hydrogen peroxide [H2O2] determination) [173]. The sensitivity of the systems has been improved by following different amplification strategies based on bionanocomposite label probes. Gold nanoparticles (AuNPs) have already proven their excellent biolabeling properties for sensing in general. Moreover, as CL labels, AuNPs are frequently used because of their rapid and simple chemical synthesis, and high biocompatibility, as well as its ease of preparation for a wide range of particle sizes [173,174]. For example, AuNPs were functionalized on a luminol-biotin-nanoprobe to be further modified with the secondary antibody labeled with streptavidin horseradish peroxidase (StrepHRP) [175]. The CL signal was produced by the luminol oxidation which was reported to be 40-times higher than other previous work

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[176]. Thus the sensitivity is directly related to the number of CL labels functionalized on the secondary antibody [173]. In this context, two similar CL immunoassays based on AuNPs/ antibody/enzyme were designed for human IgG [174] and α-fetoprotein (AFP) [177], respectively. Another AFP immunoassay involved the use of a new CL enhancer (bromophenol blue) and was based on the use of an MBs-AuNPs hybrid material modified with HRP-labeled anti
AFP antibodies [178]. The detection limit was 1 order of magnitude lower than that obtained without using AuNPs and much lower than that typically achieved by ELISA, thus, proving the synergistic effect of the hybrid composite. When using 4-(40 -iodo) phenylphenol as a signal enhancer for AFP detection, a very low LOD of 5 pg/mL (0.008
0.3 ng/mL linear range) was obtained [172]. Yang et al. proved that mesoporous-SiO2 nanoparticles (MSN) and mesoporous-carbon nanospheres could be also used as carriers for nanoprobe preparation [179]. A CA-125 cancer biomarker immunoassay has been developed by utilizing the CL resonance energy transfer to graphene QDs. The biosensor showed a low LOD of 0.05 U/mL (with a linear range of 0.1
600 U/mL) in the buffer solution, and 0.08 U/mL when tested in buffer solution with blood plasma ratio of 1:1, respectively [180]. Since the first detailed ECL investigations described by Bard in 1965 [181] and later in the ECL study of nanoparticles with SiNPs pioneered by Ding in 2002 [182], the ECL of nanoparticles have received considerable attention in various fields from pharmaceutics and environmental analysis to amplified biosensing [171,183
185]. In order to obtain high sensitivity, efficient transduction techniques should be sought to monitor successful binding events between target analytes and bioelements. Due to several appealing features such as high sensitivity, rapidity and simplicity, reproducibility, and low background, ECL-based aptasensors and

immunosensors have attracted intensive research interest. Moreover, due to their high potential for miniaturization, in situ analysis, and multichannel detection, ECL biosensing has important applications in clinical diagnoses with promising advantages [172,186]. Different ECL tags have been used for labeling biomolecules such as, Ru complex, luminol and its derivatives, quantum dots, and dendrimerencapsulated palladium nanoparticles [187]. One of the most reported ECL reagents is tris (2,20 -bipyridyl)ruthenium(II), (Ru(bpy)321) because of its ability to generate light on reaction with the oxidized form of the complex. By its covalent incorporation in SiNPs, an enhanced ECL signal was obtained due to the nanostructured labeled matrix formation [188]. A complex of AuNPs and Ru(bpy)321-labeled aptasensor was designed for thrombin detection. This structural change resulted in an obvious ECL intensity increment due to the decreased quenching effect of the ferrocene label (Fc) to the ECL substrate [189]. Another, aptasensor reporting the ECL of Ru(bpy)321 based on an Fc-labeled aptamer showed a remarkable eightfold change in ECL emission with a LOD of 1 fM for adenosine detection [190]. Nafion-hybrid composite materials were imagined to improve the performance of the sensors having the ECL substrates immobilized at the electrode surface [184,185,191]. In ECL technique, the chemiluminescent signal is electrogenerated at the electrode surface; therefore, ECL suffers less from possible quenchers in solution that could affect the optical properties compared to fluorescent and chemiluminescent techniques [192]. Thus, in this context, ECL detection is more advantageous for sensing purposes.

6.5.2 Fluorescent Materials Fluorescent materials have been widely used for optical biosensing applications, contributing to an enhanced inherent sensitivity and selectivity which can be imagined

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measuring up to a single cell [21,162]. A fluorescence biosensing system can be applied for: (1) direct determination of a molecule with fluorescent properties; (2) indirect determination of a fluorescent label which will optically transduce the target molecule presence; and (3) fluorescence energy transfer (FRET) which uses a fluorophore to generate the optical signal [193]. Fluorescence immunoassays are one of the most common approaches in the field of optical biosensors; they benefit of the high selectivity of immunoreaction event combined with the high sensitivity of fluorescence detection. Because of its high efficiency as a luminescent material, β-NaYF4:Yb,Er has been used for different biolabeling sensing applications [172]. In this context, Liu et al. proposed an IgG immunoassay that was able to detect very low levels of protein (LOD 5 0.1 ng/mL) [194]. A nanocomposite based on AuNPs-coated MBs and albumin nanoparticles (AlbNPs) loaded with fluorescent dyes served as molecular nanocarriers to develop a magnetic fluorescent miRNA sensing system [195]. The AuNPs were used for quenching the fluorescence signal when in the vicinity of the fluorophores leading to an LOD as low as 9 fM for miRNA-21 [196]. Furthermore, signal amplification plays an important role in cellular imaging due to the unique electronic, optical, and biocompatible properties of nanoparticle probes [172,187] Cancer cells surface expressed with sialic acid and N-acetylglucosamine can be triggered with the aid of a new hybrid nanocomposite with excellent fluorescence and magnetic properties formed by the encapsulation of QDs and nano-γ-Fe2O3 in poly(styrene/acrylamide) copolymers functionalized with wheat-germ agglutinin [197]. By changing the functionalized molecule (e.g., lectine) other biotargeting fluorescent nanospheres were designed [198]. Another approach for in situ signaling of sialic acid
cancer cells was imagined using a hybrid

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composite based on polysialic acid embedded Au nanoprobe with fluorescent properties given by 3-(dansylamino) phenylboronic acid [199]. MBs-reinforced polymer nanocomposites have been reported as effective contrast agents for enhancing magnetic resonance imaging (MRIs) of tissue/organs [200]. Liver bioimaging was performed with the aid of a block copolymer (methoxy poly(ethylene glycol)-bpoly(ε-caprolactone)—mPEG-b-PCL) as a MRI contrast agent coupled with Mn-doped Fe2O3 NPs [201]. By contrast, Lin et al. [202] used the superparamagnetic properties of Fe3O4 NPs to form a nanocomposite with in situ polymerized polydopamine film as the theranostic agent with high fluorescence quenching efficiency and near-infrared absorption. 6.5.2.1 Photoelectrochemical Materials Photoelectrochemical (PEC) biosensing has recently been at the forefront of research due to the unique optical, electronic and catalytic properties of the new nanomaterials [203,204]. Combining light as an excitation source and current as a detection signal, different sensing strategies based on reactant determinant, electron transfer, or energy transfer can be employed. Among the nanomaterials used for electrochemical sensing some have had tremendous importance, these include noble metal NPs, metal oxide NPs, carbonaceous NPs, and hybrid composites of these materials [187,205]. Chemical reduction, UVphotoactivation, sonochemical, and laser pulse methods are some of the methods used for their syntheses [205,206] Further, some PEC sensors based on nanocomposites will be briefly presented. An SnO2 NPs-based PEC sensor was reported for the sensitive detection of adenosine triphosphate from cancer cell extracts [207]. By using TiO2 NPs, the synthesis of a hybrid water soluble porphyrin derivate was developed by Tu et al. [208], while the photoelectrochemistry of CdS NPs has been

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FIGURE 6.5 Biosensor based on β-cyclodextrin modified glucose oxidase attached through a tetrazine derivative to β-cyclodextrin modified polypyrrole film by host
guest interaction [211].

used for the detection of tyrosinase activity [209]. Moreover, a synergistic effect could be seen when Au was integrated in TiO2 nanotubes doped with acetylcholinesterase [210]. A hybrid material that exhibited both electrochemical and fluorescent properties was successfully used as a “bridge” to immobilize the β-cyclodextrin-tagged glucose oxidase (Fig. 6.5). Tetrazine derivatives were immobilized at the electrogenerated polypyrroleβ-cyclodextrin film through the host
guest interactions as reported by Fritea et al. [211]. The achievement of highly-organized microto nanostructures for a photosensitive electrogenerated trisbipyridinyl Ru(II) metallopolypyrrole film was reported by Fritea et al. [212]. For this purpose, polystyrene beads were combined with electropolymerized metallopolymer (poly-[Ru(II)-pyrrole]) resulting a highlyorganized honeycomb microstructured polymer with photosensitive properties. Another class of photoaffinity materials is represented by diazirines, which are

photoreactive entities that can photogenerate highly reactive species such as nitrene, triplet carbonyl state, or carbine under UV illumination. Cosnier et al. made an important contribution in this filed by reporting the synthesis and characterization of various photoreactive molecules [213
217]. For example, Hosu et al. [217] described the development of an original interfacial hybrid cross-linker molecule combining electrochemical and photochemical properties by substituting two functional groups, pyrene and diazirine [216,217]. The capacity of biomolecule immobilization by photografting two enzymes (glucose oxidase and polyphenol oxidase) was investigated (Fig. 6.6). 6.5.2.2 Luminescent Optical Labels Luminescent NPs have been widely investigated for sensing and imaging in biomedical and environmental fields. These NPs include fluorescent quantum dots (QDs), dye-doped silica, rare-earth-doped downconversion (DC) and upconversion (UC) NPs [196,218]. These

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an enhanced biocompatibility and thermal stability compared to the pristine polymeric fibers alone [223]. However, one significant problem can be the release of toxic components in QDs, which affects the surrounding biological environment [21,219]. Inspired by the PEC properties of TiO2 nanotubes arrays (TNA) and their bioimmobilization features, in situ generated CdS QDs on a TNA platform were developed for asulam detection [224]. Recently, a new class of luminescent optical labels reached the forefront of the research area. Lanthanide-doped UC NPs have become promising alternatives to QDs and organic fluorophores in the development of new sensing and imaging systems [194].

FIGURE 6.6 Pyrene-diazirine photoreactive platform for enzymatic sensor development.

luminescent NPs are usually bioconjugated to functional peptides, antibodies, and DNA with specific recognition ability for targeting imaging and selective detection. Nevertheless, MIPs have shown their contribution in the development of highly sensitive nanoplatforms by coating MIPs on luminescent NPs [218]. QDs have already proven their contribution in this field because of their optical properties, photochemical stability and molar-scale production, being increasingly investigated as filler in polymer matrices and contrast agents for applications in medical imaging [219
221]. While luminescent QDs work as energy transfer donors with an array of organic dye acceptors [204], Pons et al. [222] demonstrated the quenching effect of AuNPs for conventional dye donors. Semiconductor QDs are a novel generation of fluorescent labels (e.g., CdSe, CdTe) [196]. New hybrid nanocomposite fibers based on CdSe/ZnS QDs and poly(methyl methacrylate) were electrogenerated showing

6.5.3 Hybrid Materials Used for Surface Plasmon Resonance and SurfaceEnhanced Raman Scattering Sometimes labeling of molecules is not desired or possible and, moreover, requires an additional preparation step. An optical technique that reports the development of label-free sensors is SPR [193]. The first SPR biosensor was designed by Liedberg et al. in 1983 [225] and it gradually became one of the most used label-free tools with applications in a wide range of fields [172]. The SPR technique involves a physical optics phenomena produced by optical coupling of thin metal films [172]. However, small-molecular weight targets might raise some issues in the system sensitivity as a result of the small change in the refractive index corresponding to the binding event [226]. Therefore surface functionalization plays an important role in overcoming these challenges. Although, AgNPs showed the highest plasmon resonance within visible spectrum, the highest efficiency and stability for long term usage were attributed to AuNPs [227]. Typically, AuNPs show an intense red color characteristic to well-dispersed

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nanometer size (10
50 nm diameter) particles with a localized SPR effect [228]. By applying the excellent properties of AuNPs in the design of an immunoassay, the SPR signal could be grater enhanced. Liu et al. [229] reported a sandwich assay for human C4 concentration in the range of 0.05
5 μg/mL with much lower LOD (40 times) compared to the direct SPR assay. Another AuNPs-based SPR immunosensor was developed for prostatespecific antigen (PSA) [230]. By using AuNPs a greatly enhanced SPR signal was obtained of about 100-times higher, thus, proving their outstanding effect in amplifying immunosensor sensitivity [231]. A hybrid composite based on Fe3O4@Au-AFP secondary antibody was used as signal amplification to detect AFP by a SPR sandwich immunosensor [232]. Moreover, AuNPs can be applied in surface plasmon resonance imaging (SPRi) to overcome its oftenlimited sensitivity. Hu et al., [233] reported the use of AuNPs for SPRi signal amplification for the detection of several mycotoxins (aflatoxin B1, ochratoxin A, and zearalenone). In addition to their unique SPR effect, AuNPs can also be used as probes for singlemolecule SERS detection [204]. SERS triggers the vibrational and rotational transitions of a molecule while being light scattered. Fleischmann et al. [234] discovered the enhanced Raman signals of pyridine by using a silver electrode. Thereafter, many papers reported the use of SERS for sensitive singlemolecule detection and imaging [228]. The great enhancement at the contact of two or multiple metal NPs regarding the SERS signal was firstly proved by Kneipp et al. [235] and Nie et al. [236]. Some examples reporting the use of the SERS technique for medical imaging by the use of different nanomaterials are presented as it follows [187,237,238,239]. For example, Chen et al. [237] used multiplepolysaccharide-coated AuNPs functionalized with lectins to trigger the glycan at the cell surface by Raman light scattering. For the same

purpose, Song et al., [238] imagined multicore nanostructure satellites with enhanced sensitivity formed by Au nanoflowers (AuNF) and poly(N-acetylneuraminic acid)/AuNPs as nanoprobes to recognize the sialic acid at the cell surface. Thus the use of AuNF resulted not only in an increased surface area, but also proved their fingerprinting features for sensitive SERS medical imaging [187,239].

6.6 CONCLUSION Defined as something “of mixed character; composed of different elements,” having or combining particular properties that cannot be found in other types of materials, hybrid materials represent a very large and heterogeneous class of materials. This fact explains the great variety of synthesis and fabrication methods and the wide range of properties exhibited by hybrid materials. Even the classification into inorganic, organic
inorganic, and organic hybrid materials is not ideal, this attempt helped us to introduce some rigor in the presentation of this extremely large and varied group of materials. This chapter is limited to the hybrid materials applied in biosensing, thus other uses and employments in health or generally in the biomedical field were, at best, only collaterally mentioned. Inorganic hybrid materials have very few biosensing applications, playing, in general, either the role of the transducer (electrode) if they present good electrical features or as support for a multitude of functionalizations and the immobilization of a great variety of molecularly biorecognition receptors. Thus colloidal and titanium-oxo clusters, alloys, and metal oxide hybrids have mainly photocatalytic or electrocatalytic properties, and are very promising for applications in lithium-ion batteries. However, the new colloidal particles, named patchy particles could be very promising for a

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better understanding of the formation of virus capsids and of peptide and protein interactions in the living world. While titanium-oxo clusters, alloys, and metal
oxides hybrid combined with carbon-based nanomaterials could, in the future, present potential applications in sensing, self-assembled inorganic nanorods, core
shell nanoparticles, and oriented mesoporous-silica materials. All of these already exhibited a great number of sensing examples in a virtually endless variety of configurations. Organic hybrid materials are mainly used for electroanalytical performance improvement and present a great opportunity for various chemical modifications and functionalizations by playing the role of platforms for the immobilization of bioelements, such as enzymes, antibodies, nucleic acids (DNA, RNA, microRNA), and biomimetic receptors (cyclodextrins, calixarenes, crown-ethers, etc.). Organic porous polymers and conductive polymers modified with other organic compounds are a few examples of organic nanohybrids that have been used in biosensing area. Most examples of sensing applications are given by organic
inorganic hybrid materials and by the multifunctional ones. Combining the high chemical and thermal stability of inorganic compounds with the synthetic versatility and reactivity of the organic ones, which makes possible the molecular structure modulation, the organic
inorganic hybrid materials represent ideal promoters for the development of the most varied and high-performing configurations of chemical, electrochemical, and optical sensors. The most significant progress concerning organic functional macromolecular compounds combined or modified with inorganic materials such as carbon-based materials (CNTs, graphene), metal oxides, or metal nanoparticles, magnetic micro- or nanobeads, functionalized clays and silica, and ionic liquids recently reported in the literature were summarized in this chapter.

197

A special subchapter (Section 6.5) is dedicated to multifunctional advanced hybrid materials applied in optical biosensor design and development, emphasizing sensing methods based on fluorescence, CL, ECL, PEC materials, luminescent optical labels, SPR, and SERS applications in which hybrid materials are widely used. Currently, more and more advanced hybrid materials are employed for health applications, in some cases, joined together in ingenious “hybrid devices” such as implantable pacemakers, biocompatible catheters and prosthetics, coronary stents, dental materials, and others.

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C H A P T E R

7 Smart, Portable, and Noninvasive Diagnostic Biosensors for Healthcare Srinivasulu Kanaparthi, Patta Supraja and Shiv Govind Singh* Department of Electrical Engineering, Indian Institute of Technology Hyderabad, Kandi, India

7.1 INTRODUCTION Biosensors are widely used as enticing alternatives to the large, expensive, and sophisticated analytical instruments utilized in the healthcare domain. Over the years, many of those devices have been developed for detecting various analytes using optical, piezoelectrical, and chemical transducers. Of these, chemical sensors have gained a dominating role in clinical medicine owing to their high performance, portability, simplicity, and low cost. However, the majority of these sensors require blood samples for diagnoses. Such invasiveness is an obstacle to the patient and hinders the data acquisition required for frequent health monitoring. This is particularly the case for infants and aged patients, from whom blood collection is extremely difficult. Continuous monitoring is of paramount importance in numerous areas of medical applications. For instance, optimal diabetes management requires frequent glucose concentration monitoring. Similarly, athletes need

continual analysis of their fitness levels. Regular and real-time monitoring of pathogens in physiological fluids can be used for early prediction of various diseases. Tracking drug efficiency is another context within which continuous measurements are of obvious significance. In this regard, intrusive devices have constraints as continual availability of the specified fluid samples such as blood and urine is impractical. Sweat and exhaled breath gas sensors have gathered significant attention over the past decade. These nonintrusive sensors offer noteworthy performance for regular or continuous monitoring of a person’s health and fitness. By using the data provided by these sensors at regular intervals, it is possible to change the lifestyles of people. The increasing curiosity in these sensors indicates changes from hospitalbased healthcare systems to home-based personal health monitoring, as the latter is inexpensive. Sweat contains ample data about a patient’s health and, thus it is a superb physiological fluid for noninvasive biosensing. For

* Corresponding authors: To whom correspondence should be addressed

Advanced Biosensors for Health Care Applications DOI: https://doi.org/10.1016/B978-0-12-815743-5.00007-X

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instance, various ions in sweat are indicative of abnormalities such as electrolytic imbalance and stress. Thus continuous monitoring of sweat biomarkers is significant for optimal healthcare. In this chapter, we review the recent developments of sweat-based sensors to detect various diseases like diabetes and cystic fibrosis and for monitoring various ions in sweat. On the other hand, exhaled breath contains a large number of volatile organic gases and a specific set of volatile organic compounds (VOCs) serve as a biomarker for diagnosis of a specific disease. Exhaled breath analysis is a promising technique for early detection of diseases as it is noninvasive, inexpensive, and convenient to use for continuous and regular monitoring. The change in gas profiles in exhaled breath VOCs can be used to discriminate a healthy person from a person suffering from a specific disease. A number of studies reveal that there is strong relation between exhaled VOCs and specific diseases. Particularly, acetone, hydrogen sulfide, nitric oxide, and ammonia have strong correlation with diabetes, halitosis, asthma, and kidney diseases, respectively. Several studies reported that multiple VOCs serve as biomarkers for a single disease and a single VOC can serve as a biomarker for many diseases. Intensive research is ongoing to identify various biomarkers for various diseases to improve the accuracy of detection of specific diseases using gas profiling. Similarly, gas profiles in the gut also give information concerning metabolism and fermentation of food and serve as a biomarker for many diseases. In this chapter, we review the exhaled gas humidity for respiration monitoring to diagnose sleeping disorders and cardiovascular diseases, detection of VOC concentrations to identify a specific disease, and ingestible electronic pills to measure the gas profile of the gut.

7.2 WEARABLE SWEAT SENSORS Detection of diseases at an early stage by using sophisticated biomarkers and continuous

monitoring the concentration of biomarkers is a challenging task. Blood is one of the most predominant body fluids which contain metabolites, long-chain proteins, metal ions, and several other biomarkers which can help us to identify several chronic diseases. The main disadvantage of blood-based disease diagnosis is the risk of collecting blood and that continuous monitoring is critical task. Therefore we need an alternative noninvasive body fluid which reflects the information in blood and be accessible in high quantity without any risk. Among sweat, saliva, urine, and tears, sweat is the most easily accessible as well as very well suited for the wearable sensing platform. Concentration of analytes in sweat compared to blood is low; therefore we need a sophisticated sensing platform which can detect analytes at lower concentrations. In the following subsections we discuss the principles and methods of sweat-based sensing of saccharides, alcohols, acids, and metal ions.

7.2.1 Detection of Saccharides Saccharides are the essential components required for human body. Saccharides are mainly divided into four categories, namely, monosaccharides (glucose, fructose, galactose), disaccharides (lactose, sucrose, maltose), and polysaccharides (starch, glycogen, cellulose). Imbalance of carbohydrates or saccharides in the human body leads to several health problems. Thus detection and continuous monitoring of carbohydrates noninvasively in sweat is highly essential. 7.2.1.1 Glucose Diabetes is the most predominant and persistent disease in the world. Most diabetes patients require a management system to regularly monitor their glucose levels. A wearable sensing platform is the most suitable method for regular noninvasive monitoring of analytes. Among various types of sensing methods, electrochemical

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sensing is the most promising and widely used technology because of its simplicity, low cost, high performance, high sensitivity, and reproducibility. An electrochemical transducing mechanism uses an Ag/AgCl reference electrode to provide constant potential to the system and also for precise measurement of signals, but integrating the Ag/AgCl reference electrode on to the biosensing chip is difficult and remains challenging to researchers. Liu et al. have developed indium oxide (In2O3) based field-effect transistor (FET) biosensors for detection of glucose levels in sweat [1]. In2O3 FET-based biosensing platforms are well acceptable and suitable for wearable biosensor applications because of their high sensitivity, repeatability, reproducibility, wide-detection range, and fast response. Functionalizing on metal oxides and integrating capability of functionalized biosensor onto microfluidic environment is easy [2,3]. In this FET-based biosensor a metal gate works as the reference electrode which supplies stable gate bias to the FET device. The gate, source, and drain are fabricated by two-step shadow masking technique on an ultraflexible Polyethylene terephthalate (PET) substrate. In2O3 nanoribbons act as the semiconducting channel between the source and drain. Fig. 7.1A shows the schematic representation of In2O3based FET fabricated on PET substrate. Most of

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the electrochemical biosensors use the dropcasting technique to immobilize or functionalize analytes on to the electrodes, but it is not an effective method for chip fabrication because of its small dimensions. Inkjet printing is an alternative and promising technique for chip scale fabrication of devices. Fig. 7.1B shows the working principle of a glucose sensor. Usually, single-walled carbon nanotubes (SWCNT), glucose oxidase (GOx), and chitosan are used to prepare inkjet for the source and drain. Chitosan is a biocompatible material which is used as the matrix to immobilize SWCNT and GOx. Carbon nanotubes play a predominant role in increasing the sensitivity of the biosensors, either through good electrocatalytic properties, high ability to immobilize the biomolecules, or both. After forming the chitosan-film matrix consisting of SWCNT and GOx as reinforcement, the GOx enzymes accept electrons after interaction with glucose in the solution. The accepted electrons are transferred to oxygen molecules which results in H2O2 (hydrogen peroxide) The enzymatically produced H2O2 is oxidized and results in H1 ions under the applied bias voltage, as shown in Fig. 7.2B. The generation of H1 depends on the concentration of glucose. Increase in H1 decreases the pH levels and leads to protonation of the hydroxyl (OH2) groups on the indium

FIGURE 7.1 (A) Schematic representation of In2O3-based FET fabricated on PET substrate and (B) working principle of glucose sensor. Reprinted with permission from Q. Liu, et al. Highly sensitive and wearable In2O3 nanoribbon transistor biosensors with integrated on-chip gate for glucose monitoring in body fluids, ACS Nano (2018). doi: 10.1021/acsnano.7b06823. Copyright (2018), American Chemical Society. ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

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FIGURE 7.2 Schematic representation of a lactate sensor. Reprinted with permission from D. Khodagholy, et al., Organic electrochemical transistor incorporating an ionogel as a solid state electrolyte for lactate sensing, J. Mat. Chem. 22 (10) (2012) 4440
4443. Copyright (2012), The Royal Society of Chemistry.

oxide surface, which results in variation of local electric field and eventually causes a change in current and conductance. 7.2.1.2 Lactose Deprotonation of lactic acid results in lactate which is good bioenzymatic chemical for continuous monitoring of the anaerobic metabolism of patients whose circulatory system failed or tends to fail. Concentrations of lactate in sweat vary in proportion to blood levels. Therefore detection of lactate in sweat is an indirect way of measuring eccrine gland metabolism [4]. Generally, lactate concentration in sweat is 9
23 mM, but its concentration increases significantly with exercise. By detecting this trend in variation of concentration, one can monitor individual health

continuously. Thus developing a new miniaturized, reliable, cheap, robust lactate sensor is essential for both sports and healthcare applications. Khodagholy et al. have developed an organic material
based electrochemical transistor for lactate detection by using an ionic liquid (in gel form) which can also act as solid electrolytes at room temperatures [5]. The channel and gate are in ionic contact via the solid electrolyte. Due to variation of applied positive potential at the gate, the cations (which are in solid electrolyte) will try to move into the channel made up of a conducting polymer. The movement of cations into the conducting channel of organic electrochemical transistor results in the variation of doping state and also a reduction in the number of charge carriers, which are holes, consequently decreasing the channel current. The most popular conducting polymer for Organic Electrochemical Transistor (OECT) channel is poly(3,4-ethylenedioxythiophene) doped with poly(styrene sulfonate) because of its high conductivity and biocompatible nature. Fig. 7.2 presents a schematic representation of a lactate sensor. The potential which applied to the gate electrode oxidize the lactic acid to pyruvate and reduces the lactate by the ions in solid electrolyte (Fc/Fc1 couple) over the cycles. During this reduction process, it transfers electrons to the gate electrode and results in the decrease of potential across the interface of solid electrolyte and gate. Consequently, the potential at the interface of electrolyte and channel increases. As a result, more numbers of cations from the solid electrolyte solution enter into the channel and redope the channel, which results in the variation of the drain current in response to a sign and magnitude of voltage applied at the gate. The reactions are shown in Fig. 7.3.

7.2.2 Detection of Organic Compounds Organic compounds are the essential chemicals required by the human body. Some

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FIGURE 7.3 Electrochemical reactions at the gate electrode: (A) at the channel and (B) overall reaction of OECT. Reprinted with permission from D. Khodagholy, et al., Organic electrochemical transistor incorporating an ionogel as a solid state electrolyte for lactate sensing, J. Mat. Chem. 22 (10) (2012) 4440
4443. Copyright (2012), The Royal Society of Chemistry.

compounds are produced internally during metabolism and some must be supplied externally. The human body maintains chemical equilibrium by secreting some of the unwanted compounds. Imbalancing of this chemical equilibrium (due to some external/ internal factors) results in altering the concentration of organic compounds in body fluids. Quantitative detection of organic compounds can be used as biomarkers for the early detection of diseases. Among several organic compounds whose detection has paramount importance in daily life are acids and alcohols. 7.2.2.1 Uric Acid Degradation of proteins in the human body can produce uric acid (UA). The content of UA in blood is an important parameter to understand the functioning of renal system [6]. The importance of UA in strengthening the immune system was reported by Rock et al. [7]. We can find traces of UA in urine as well as in sweat where concentration is correlated with blood UA concentration. So the ultrasensitive detection of UA in body fluids noninvasively with reasonable selectivity plays an important role in clinical diagnosis of cytoplasm problems. Chu et al. developed an electrochemiluminescent (ECL) biosensor which uses polypyrrole reinforced or immobilized uricase as biorecognition species [8]. The

uricase catalyzes the oxidation of UA and produces allantoin, CO2, and H2O2 in the presence of potassium ferricyanide. Enzymatically generated H2O2 is essential for the ECL response because it oxidizes on the electrode surface and results in O22, this oxygen species can act as the oxidizing agent to oxidize UA. The catalytically yielded H2O2 is one of the most reactive forms of oxygen species which enhances the ECL response. Generally, luminol can be used as intensifier, which indirectly demonstrates the electrolysis of potassium ferricyanide. The concentration of UA can be determined by the intensity of luminol. The kinetics of the resulting ECL biosensor is as shown by: Uric acidðC5 H4 N4 O3 Þ1O2 12H2 O -allantoinðC4 H6 N4 O3 Þ1CO2 1H2 O2 This detecting mechanism can further be integrated with flexible sensing mechanisms for real-time applications, which we will discuss in detail in the following section. 7.2.2.2 Detection of Alcohols User-friendly and real-time detection of alcohols with high accuracy has tremendous importance in real life because very high levels of alcohol consumption can lead to accidents and risky behavior. Alcohol-related incidents

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and health issues are rising rapidly in most countries, especially in developing countries. Thus there is an urgent need for a device which can measure alcohol content fast with high accurately. 7.2.2.2.1 ETHANOL

Breathalyzers are the most widely used alcohol sensing devices to estimate blood
alcohol concentration indirectly by measuring the alcohol concentration in breath. The accuracy of these devices is less because the results depend on external factors like temperature and humidity [9,10]. Therefore the development of a new prototype to monitor alcohol level noninvasively in real time is highly desired. Kim et al. described effective noninvasive ethanol monitoring based on integrating of iontophoresis (used for inducing sweat) and amperometric (biosensing of sweat ethanol) onto the tattoo-based wearable system, as shown in Figs. 7.4A and 7.4B. The sensing platform consists of two iontophoretic electrodes, namely, a cathode and anode, and three amperometric electrodes, namely, reference (Ag/AgCl), working (Prussian blue conductive based), and counter electrode (carbon). By applying constant current, the anode compartment of the iontophoretic electrodes introduces pilocarpine (large, positively charged molecule) to induce sweat [11]. The functionalization of enzymatic alcohol oxidase (AOx), BSA stabilizer, and chitosan layer on the working electrode of amperometric system facilitate to detect alcohol in sweat. The principle for sensing of alcohol in sweat is based on the electrocatalytic detection of H2O2. Usually, alcohol oxidase can be used as transducer which converts enzymatic detective biological event to current signal. The working electrode, which is immobilized by alcohol oxidase, needs to be covered with agarose gel which contains potassium ions to maintain stable potential even for different electrolyte concentrations. The schematic representation of enzymatic detection is shown in

Fig. 7.4D. Sweat generation by iontophoresis followed by he detection of ethanol by the threeelectrode
electrochemical system and transmitting the detected information to Android mobile via Bluetooth is shown in Fig. 7.4C.

7.2.3 Detection of Ions Generally the human body is composed of almost all metals available in nature in the form of ions. Metals are essential for the healthy functioning of organs and systems. Metallic ions like sodium, calcium, potassium, magnesium, and sulfur have major functional significance in body. Any imbalance in the levels of these metallic ions in the body leads to several disorders. Therefore detection and monitoring the levels of ions in body fluids can be used as predominating biomarkers for the detection of diseases. In the following subsection, we discuss ammonium ion and metal ion detection. 7.2.3.1 Ammonium Ion Prior to excretion, the liver converts ammonia to urea. Any abnormality in the working of the liver results in changes in the ammonium ion concentration levels in the body. During body metabolism, proteins will break instead of carbohydrates if carbohydrate resources are less in quantity. The breakdown of proteins results in the presence of ammonium in blood or plasma and, thus, there might be a temporary increment of ammonia levels in plasma due to low levels of carbohydrates in the diet. Therefore determination of ammonium ion levels in plasma provides relevant information about dietary conditions and metabolic state of individuals. Czarnowski et al. reported there exists a clear relation between blood ammonium levels and sweat ammonium levels [12]. Hepatitis and some of the hepatic disorders use ammonium ions as biomarkers for detection. Guinovart et al. developed an ion-selective potentiometric cell to monitor

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FIGURE 7.4 (A) Schematic representation of the ethanol sensing device with sensing electrochemical and iontophoretic electrodes. (B) Wearable ethanol sensing device. (C) Schematic representation of ethanol sensor interfaced with android mobile. (D) Schematic representation of iontophoretic system and detection mechanism at electrochemical electrodes. Reprinted with permission from J. Kim, et al., Noninvasive alcohol monitoring using a wearable tattoo-based iontophoretic-biosensing system, ACS Sens. 1 (8) (2016) 1011
1019. Copyright (2016), American Chemical Society.

the ammonium levels in sweat by using a tattoo-transfer platform [13]. Solid-state potentiometric sensing technology along with ammonium-selective polymeric membranes which contain nonactin ammonium ionophore was used for ammonium ion detection. The ammonium-selective membrane contains

nonactin, 2-nitrophenyl octyl ether (o-NPOE), and poly(vinyl chloride). Usually real-time sensors are either solid-state potentiometric or amperometric systems. One of the great features of solid-state potentiometric devices is their constant sensitivity. Solid-state potentiometric sensors are widely preferred because

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they are easy to operate and their power consumption is extremely low. Generally, polyvinyl butyral (PVB) or Ag/AgCl can be used as the solid-state reference membrane in a wearable device. Tattoo-printed ammonium sensors can be placed directly on the skin. A schematic representation of the step-by-step fabrication of the potentiometric sensor is shown in Fig. 7.5. Contact between the electrodes is one of the practical problems raised for this kind of sensor. In order to solve this issue, Kapton (an electrically insulating polyimide film widely used in flexible electronics) strips are placed on both sides of tattoo. These strips create a path where the sweat is able to flow through keeping the electrical contact between both

electrodes through the solution. Additionally, a filter paper at the end of the path is used as a “sink” to retain the sweat. 7.2.3.2 Sodium Ion Sodium ion concentration is an excellent biomarker to provide electrolyte imbalance and valuable information regarding an individual’s physical and mental situation. The levels of sodium ions in the human body can depict pH, water imbalance, and osmotic pressure. Cystic fibrosis is a disease caused due to heavy loss of sodium ions. Sometimes rigorous loss of sodium ions via sweating can cause hyponatremia [14]. Therefore sensing sodium levels in human perspiration during diverse activities

FIGURE 7.5 Schematic representation of the step-by-step fabrication of the potentiometric sensor. Reprinted with permission from T. Guinovart, et al., A potentiometric tattoo sensor for monitoring ammonium in sweat. Analyst 138 (22) (2013) 7031
7038. Copyright (2013), The Royal Society of Chemistry.

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7.3 GAS SENSORS FOR HEALTHCARE

relevant to healthcare, fitness, military, and skin-care domains has significant importance. The detection of sodium ions is quite similar to that of ammonium ions. The difference is that sodium ion detection systems use ion-selective working electrodes which functionalized with sodium ion sensitive ionophore. Bandodkar Amay et al. group has developed potentiometric epidermal tattoo based sensing platform by using sodium ion selective ionophore [15]. The sodium-selective membrane contains sodium ionophoreX, Na-TFPB (sodium tetra kis[3,5-bis (trifluoromethyl)phenyl] borate), PVC (polyvinyl chloride), and DOS (bis(2-ethylhexyl) sebacate) dissolved THF (tetrahydrofuran). The reference electrode membrane contains PVB and NaCl dissolved in methanol. The polymeric PVB membrane, which contains electrolytes, forms a nanoporous structure that allows the exchange of electrolytes with the solution. Reference provides a stable potential, sensitive to changes in the ion concentration over a large concentration range. 7.2.3.3 Calcium Ion Calcium is one of the most essential minerals for human metabolism. The quantity of calcium in human body is approximately 1%
2% of the weight. Usually, the concentration of Ca21 ions in normal body fluids varies from 0.5 to 3 mM [16]. Based on the calcium levels in the physiological fluids, the function of the organs in the body can be predicted. Extreme changes in calcium ion concentration in the human body can cause diseases such as myeloma, cirrhosis, normocalcemic hyperparathyroidism, renal failure, and acid
base balance disorder [16]. Generally, we select ETH 129 as Ca21 ion-selective ionophore because of its ability to translocate calcium ions across the membranes [17]. Ionophore (ETH 129) embedded in a conductive polymer membrane is used to measure the signal selectively for calcium ions. In most of the cases, the detection of Ca21 ions along with pH plays a crucial role in

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diagnoses of diseases. Usually, pH of the skin changes with evaporation of sweat and, thus it is necessary to monitor the pH along with the detection of Ca21 ions. Polyaniline is a promising material for pH measurement in physiological fluids. Polyaniline can be electrochemically deposited onto the electrode (Au, Pt) by the well-known cyclic voltammetry technique. Integration of the Ca21 ion sensor and pH sensor onto the flexible PET substrate can facilitate easy and early detection of diseases [18].

7.3 GAS SENSORS FOR HEALTHCARE 7.3.1 Breath Water Vapor Sensing for Respiration Monitoring Respiratory information along with heart rate and body temperature are vital parameters to know the health status of a person. Irregularity in breathing might be an indication of many diseases such as cardiovascular and pulmonary diseases, sleep apnea, and dehydration [19
21]. Despite the availability of numerous techniques to monitor breath pattern, humidity sensors gain importance due to their ease of use, low cost, portability, and flexibility. Two types of humidity sensing techniques have been implemented to monitor the respiration rate and using humidity sensors, namely, (1) resistive and (2) capacitive [19
21]. In order to implement a resistive sensor, the sensing film should be selective to water vapor and should be insensitive to other gases in exhaled air like ammonia. Exhaled air of a healthy human being contains ammonia levels between 50 ppb and 2 ppm. If the sensing film is sensitive to ammonia, we cannot guarantee that the sensor response is due to humidity. Thus the sensing film of the sensor should be insensitive to other gases and other major VOCs in the exhaled air. In this regard, Kano et al. (Fig. 7.6) and Gu¨der et al. reported resistive humidity

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FIGURE 7.6 Schematic silicon nanocrystal respiration sensor and its response to different breath rates. Reprinted with permission from S. Kano, K. Kwangsoo, F. Minoru, Fast-response and flexible nanocrystal-based humidity sensor for monitoring human respiration and water evaporation on skin, ACS Sens. 2 (6) (2017) 828
833. Copyright (2017), American Chemical Society.

sensors for breath monitoring based on silicon nanocrystals and paper as active elements, respectively [19,20]. Both silicon nanocrystals and paper form hydrogen bonding with water vapor and are insensitive to other gases and VOCs. Although these sensors are selective to humidity, they suffer from strain-induced errors. In order to minimize error due to deformation strain, we reported a capacitive humidity sensor that is insensitive to strain and showed excellent response to respiration [21]. In addition, there are numerous reports on this concept, but authors did not address the selectivity of sensor to breath humidity and strain-induced errors. Even though the humidity sensors are reliable techniques for respiration monitoring, this concept fails with the interference of ambient humidity. For the reliable operation of humidity-based respiration sensing, the ambient humidity should be constant. Thus humidity sensors can be used in a closed room where humidity can be controlled.

7.3.2 Exhaled Volatile Organic Compound Monitoring Exhaled breath contains several lowconcentration VOCs in addition to nitrogen,

oxygen, and carbon dioxide. These VOCs are originated either from endogenous or exogenous sources. Endogenous VOCs are generated by biological processes such as oxidative stress and inflammation. These VOCs are excreted into the blood and then enter the lungs where they will be exhaled. These VOCs can serve as biomarkers for diagnosing various diseases. 7.3.2.1 Acetone Sensing for Diabetes Diagnosis The acetone concentration in healthy people is 300
900 ppb whereas it increases to 1000
1800 ppb in type-1 diabetes patients [22]. Thus sensors with high sensitivity and specific to acetone and sub-ppm concentration detection capability are required to diagnose diabetes. Current methods to monitor and diagnose diabetes involve blood glucose monitoring which poses pain and risk of infection due to needles and, thus continuous or regular monitoring is quite difficult and uncomfortable for the patient. On the other hand, breath-acetone monitoring is a viable alternative and noninvasive technique to diagnose diabetes patients frequently and continuously. Although many sensors based

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7.3 GAS SENSORS FOR HEALTHCARE

on metal oxide semiconductors have been reported for acetone sensing, most of them are cross-sensitive to ethanol [23,24]. Thus a small amount of alcohol traces in the mouth may give false measurements during the diagnosis of diabetes patients. However, ethanol is dehydrogenated with basic oxides and dehydrated with acidic oxides and dehydrated gases show low response compared to dehydrogenated gases [25]. Therefore use of acidic oxides for acidic gases can show lower crosssensitivities to ethanol and presence of small amounts of alcohol in the breath does not affect the diagnosis results. Alternatively, a gas sensing array can be used to increase the precision, accuracy, and selectivity of acetone sensing for diabetes diagnoses. Choi et al. reported a catalyst-loading method using a platinum-loaded copolymer as a template for catalyst transfer to tungsten oxide (WO3), and nanofibers to get uniformly distributed platinum nanoparticles on WO3 nanofibers, as shown in Fig. 7.7 [7.26]. An array of sensors with different loadings of platinum on WO3 nanofibers were used to detect various analytes, including acetone. Pattern recognition

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algorithms were used to successfully discriminate acetone with other interfering gases such as toluene, H2S, and CH3SH.

7.3.2.2 Hydrogen Sulfide Detection for Halitosis Diagnosis Halitosis, also known as bad breath, entails unpleasant odor of exhaled breath which has social and quality-of-life implications, and is categorized into intraoral and extraoral halitosis [27]. Intraoral halitosis originates from the mouth because of gum diseases, bacteria, and dehydration whereas the extraoral halitosis occurs due to diseases like ulcers and malfunctioning of the kidneys and liver [28]. As the extraoral halitosis is due to malfunctioning of different organs in the body, it is very important to diagnose at early stages to cure this disease. In general, bad breath is composed of mainly hydrogen sulfide and methyl mercaptan (CH3SH) whose concentration is less than 0.15 ppm in healthy people whereas this concentration increases to 1 ppm in halitosis-affected patients [29]. There have been quite a few commercial devices based on

FIGURE 7.7 (A) Synthesis of platinum-decorated WO3 nanofibers using template method and (B) classification of different analytes including acetone using an array of sensors and principal component analysis. Reprinted with permission from S.-J. Choi, et al., Novel templating route using Pt infiltrated block copolymer microparticles for catalytic Pt functionalized macroporous WO3 nanofibers and its application in breath pattern recognition. ACS Sens. 1 (9) (2016) 1124
1131. Copyright (2016), American Chemical Society. ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

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gas chromatography to diagnose this disease. However, these are expensive techniques and involve complex analyses of the data and are not portable to use regularly by the patients. In this regard, a highly sensitive and selective detection of portable H2S sensors are required for real-time diagnosis of halitosis on daily basis. Choi et al. reported a highly sensitive and selective H2S sensor with reduced graphene oxide (rGO)-loaded tin oxide (snO2) fibers as shown in Fig. 7.8 [30]. SnO2 fibers were produced with electrospinning and mixed with graphene oxide (GO) followed by annealing in forming gas to reduce GO to rGO. The sensor showed significant selectivity to H2S in humid environment at 0.01% GO loading.

7.3.2.3 Nitric Oxide Gas detection for Asthma Diagnosis Asthma is a respiration-related disease which has symptoms like wheezing and breathing problems. Nitric oxide (NO) gas present in exhaled breath can be used as a biomarker to characterize asthma [31]. The concentration of NO in exhaled air exceeds 30 ppb in asthma patients, whereas it is lower in healthy people [32]. As asthma is a common disease in people, there should be a noninvasive and portable diagnostic device to monitor the disease. Koo et al. synthesized WO3 nanotubes with layer-by-layer assembly of polycations, tungsten precursors, and catalysts on electrospun catalytic templates and subsequent calcination, as shown in Fig. 7.9 [33]. They

FIGURE 7.8 (A) Synthesis of SnO2-rGO fibers. (B) Dynamic response of nanofibers with different rGO loadings to H2S gas. (C) Selectivity of the sensor to H2S with 0.01% rGO loading. Reprinted with permission from S.-J. Choi, et al. Selective detection of acetone and hydrogen sulfide for the diagnosis of diabetes and halitosis using SnO2 nanofibers functionalized with reduced graphene oxide nanosheets. ACS Appl. Mat. Interf. 6 (4) (2014) 2588
2597. Copyright (2014), American Chemical Society.

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(A)

(B)

Pristine WO3 NTs

NO

Resistance (Ω)

1011

Pd- WO3 NTs

5 ppm

4 ppm

1010

3 ppm

2 ppm 1 ppm

9

10

108

PDADMAC

Pt- WO3 NTs

0

20

40

60

80

100

Time (min)

(C) Catalytic precursor

NO

Catalyst

Rair /Rgas = 20.89

H2S

Rgas /Rair = 63.59

NH3 Acetone Pentane Toluene

AMH

WO3

0

10

20

30 40 50 60 Gas response

70

80

FIGURE 7.9 (A) Synthesis of WO3 nanotubes using electrospun polymer templates and tungsten precursor and (B) selectivity of WO3 nanotubes to NO gas. Reprinted with permission from W.-T. Koo, et al., Catalyst-decorated hollow WO3 nanotubes using layer-by-layer self-assembly on polymeric nanofiber templates and their application in exhaled breath sensor. Sens. Actuat. B Chem. 223 (2016) 301
310. Copyright (2016), Elsevier.

demonstrated that these nanotubes exhibit high sensitivity and selectivity to NO gas at 350 degrees. These results indicate the potential of semiconductor metal oxide gas sensors for the diagnosis of asthma. 7.3.2.4 Artificially Intelligent Nanosensors for Multiple Disease Detection As explained, a biomarker related to a particular disease can be detected using highly sensitive and specific gas sensors. However, in general, each disease is characterized by more than one VOC and each VOC is related to more than one disease. Therefore we cannot identify the disease accurately using one VOC as well as a rise in a particular VOC concentration does not give information about the patient’s disease(s) accurately. There is a need to use an array of sensors in which each sensor is sensitive to more than one VOC and

generates a fingerprint based on the composition of VOCs present. Each disease will have a specific pattern so that different diseases can be discriminated more accurately by using pattern recognition algorithms. An advantage of this approach is that there is no need of quantifying the VOC concentrations. By measuring the pattern with the breath samples of the patients having different diseases as well as samples of healthy people, and discriminating with pattern recognition algorithms, we can train the device so that when a new sample is under study it will determine whether subjects have a disease or not, which disease they have, and what stage the disease is. There are quite a few reports on sensors to determine and discriminate diseases based on artificially intelligent nanosensor array. For example, Nakhleh et al. reported a device using an array of 20 sensors based on nanomaterials to diagnose

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and discriminate 17 diseases by analyzing multiple VOCs [34]. They used 2808 breath samples from 1404 subjects (from different parts of the world) having one of the 17 different diseases containing cancerous and noncancerous diseases and healthy people, as shown in Fig. 7.10. As represented in the heat map, each disease has a specific molecular print so that the presence of one disease does not mask the determination of other diseases, which is very important feature required for a device to diagnose multiple diseases simultaneously. The device showed an average of 86% accuracy in determining the disease and discriminated the sick people from healthy people.

7.3.3 Ingestible Sensors for Gut-Gas Monitoring Ingestible sensors are an emerging technology in medical diagnosis. Unlike conformable sensors which are in contact with the skin, ingestible sensors are swallowed and stay inside the body. Ingestible sensors collect data from sensors and transmit it through wireless communication. A qualified medical practitioner can analyze this real-time data and can recommend personalized medication to the patient. Research in this field is relatively new and very few studies such as temperature, pH, pressure, and gas monitoring have been reported. However, a fairly large number of studies reported that certain gut-gas profiles are efficient biomarkers for specific diseases. These exhaled breath gas sensors give the concentration of the gases, but they will not give information about the origin of the gases. A particular gas may be related to various diseases and, thus breath VOC sensing might give false results. Also, the signal-to-noise ratio of breath VOCs are very low, which reduces the limit of detection of VOCs. High response and recovery times are also an issue and a dedicated setup is required. Moreover,

alterations in ambient humidity affect the accuracy of the breath VOC sensing. On the other hand, in case of ingestible sensing, the abdomen gas concentration is large and constant for sufficient time and, thus measurement is quite convenient. The primary intestinal gases are nitrogen, hydrogen, oxygen, and carbon dioxide along with some other gases in small quantities. There are various techniques to measure the gut gases which include flatus analysis, calorimetry, fecal samples analysis, and breath VOC sensing. Despite the advantages of the aforementioned methods, they are invasive or uncomfortable and even unreliable. Thus the use of noninvasive methods like ingestible sensors are required to obtain real-time and reliable information on the health status of a person. Kourish et al. reported a human trial of ingestible electronic pills which can measure H2, CO2, and O2 gases and monitor the gas composition according to changes in diet using ingestible electronic pills which can transmit the collected data through an antenna, as depicted in Fig. 7.11 [7.35]. The device has multiple sensors to measure different gases and temperature modulation was used to improve the accuracy of the sensors. They measured the different gases and core body temperature at different parts of the gut at regular intervals. They also studied the gas profiles obtained with different diet compositions which will help doctors to study how the body is responding to the person’s dietary habits and thus help to give personal medication. This technique is useful to extract gases in both aerobic and anaerobic parts of the gut, which is not possible with available commercial sensing technologies. This study provides a potential diagnostic technique to alternative counterparts such as invasive and noninvasive methods available for diagnosis. These sensors will give information on what food a patient was taken, amount calories he or she consumes, as well as heart and metabolic profiles.

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FIGURE 7.10 (A) Schematic of concept and study of discrimination of multiple diseases from array of sensors. (B) Heat map of 59 sensing features extracted from 20 sensors where each column represents one of 17 diseases under study. (C) Graphical representation of accuracy of the classifiers. Reprinted with permission from M.K. Nakhleh, et al., Diagnosis and classification of 17 diseases from 1404 subjects via pattern analysis of exhaled molecules, ACS Nano 11 (1) (2016) 112
125. Copyright (2016), American Chemical Society.

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FIGURE 7.11 (A) Gas sensing capsule showing its internal components, its photograph, and circuit diagram regarding the operation of gas sensor. (B) Gas and temperature profiles at different locations within the abdomen. Reprinted with permission from K. Kalantar-Zadeh, et al., A human pilot trial of ingestible electronic capsules capable of sensing different gases in the gut. Nat. Electron. 1 (1) (2018) 79. Copyright (2018), Nature Publishing Group.

Moreover, patients taking multiple medications can miss a dosage and can cause further damage to their health. Regular use of these ingestible sensors monitors the dosages and can record what type of medication was taken and how often it was taken.

7.4 FUTURE PERSPECTIVES Diagnosis using sweat biomarkers has a lot of potential as sweat is an easily available fluid in the body. However, such sensors depend on sweat perspiration and the applications are limited to situations that require physical excretion. On-demand extraction of sweat such as iontophoresis is necessary to use these

sensors for versatile applications. Researchers should come up with alternative solutions or develop the existing solutions to extract the sweat on demand and continuously to use these sensors for variety of applications. The sensors that detect VOCs in the breath are cross-sensitive to relative humidity and ambient temperature. To get accurate results, one of the following precautions should be taken in designing and developing the sensors. (1) The sensor operating temperature should be greater than ambient or room temperature and the effect of humidity should be nullified using humidity filters. (2) The sensors working at room temperature and different humidity levels should be programmed so that the gas concentration can be calculated accurately

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REFERENCES

using machine-learning algorithms. Though ingestible electronics-based sensing is an FDA approved technology, many people are not interested in this technique as it will reveal all the health information of a person, which is a privacy concern. Thus developers of these devices should ensure the privacy of data of the patients for the success of this technology. Another reason for hesitation in using this technology is the assumption of side effects of materials used in the development of the device. Although the materials used in this technology are not harmful, nondigestibility of these materials will keep people away from using these sensors. Therefore these sensors should be manufactured with the materials which are digestible as there is a chance that these sensors might get stuck inside the body.

7.5 CONCLUSION Various noninvasive techniques to diagnose diseases using sweat and gas biomarkers using nanosensors have been reviewed. Sweat, exhaled breath, and abdomen gas analysis are very important diagnostic methods because they are noninvasive and facile. There are certain aspects of sweat sensing that should be improved to commercialize them. In addition to the challenges in analyzing the sensor data for early diagnosis of diseases and predictive healthcare, device improvement will contribute to the reliable measurement of these sensors. Sweat-based, wearable sensors suffer from external parameters like bending or stretchinginduced strain. This should be avoided by finding appropriate packaging technologies. Breath VOC sensors suffer from humidity interference, therefore these sensors should be developed with humidity insensitive materials at room temperature. Response and recovery should be improved in order to use breath VOC sensors as handheld devices. Developing

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ingestible sensors based on digestible materials may find its success to monitor patients’ health in the near future.

References [1] Q. Liu, et al., Highly sensitive and wearable In2O3 nanoribbon transistor biosensors with integrated onchip gate for glucose monitoring in body fluids, ACS Nano (2018). Available from: https://doi.org/ 10.1021/acsnano.7b06823. [2] Q. Liu, et al., Highly sensitive and quick detection of acute myocardial infarction biomarkers using In2O3 nanoribbon biosensors fabricated using shadow masks, ACS Nano 10 (11) (2016) 10117
10125. [3] N. Aroonyadet, et al., Highly scalable, uniform, and sensitive biosensors based on top-down indium oxide nanoribbons and electronic enzyme-linked immunosorbent assay, Nano Lett. 15 (3) (2015) 1943
1951. [4] J.M. Green, et al., Sweat lactate response between males with high and low aerobic fitness, Eur. J. Appl. Physiol. 91 (1) (2004) 1
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4443. [6] P. Cirillo, et al., Uric acid, the metabolic syndrome, and renal disease, J. American Soc. Nephrol 17 (12 Suppl. 3) (2006) S165
S168. [7] Y. Shi, E.E. James, L.R. Kenneth, Molecular identification of a danger signal that alerts the immune system to dying cells, Nature 425 (6957) (2003) 516. [8] H. Chu, et al., An electrochemiluminescent biosensor based on polypyrrole immobilized uricase for ultrasensitive uric acid detection, Sens. Actuat. B Chem. 163 (1) (2012) 247
252. [9] M.A. Dumett, et al., Deconvolving an estimate of breath measured blood alcohol concentration from biosensor collected transdermal ethanol data, Appl. Math. Comput. 196 (2) (2008) 724
743. [10] T.M. Worner, J. Prabakaran, The accuracy of breath alcohol analysis using the breathalyzer, Alcohol Alcohol. 20 (3) (1985) 349
350. [11] J. Kim, et al., Noninvasive alcohol monitoring using a wearable tattoo-based iontophoretic-biosensing system”, ACS Sens. 1 (8) (2016) 1011
1019. [12] D. Czarnowski, et al., Plasma ammonia is the principal source of ammonia in sweat, Eur. J. Appl. Physiol. Occup. Physiol. 65 (2) (1992) 135
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[13] T. Guinovart, et al., A potentiometric tattoo sensor for monitoring ammonium in sweat, Analyst 138 (22) (2013) 7031
7038. [14] D.B. Speedy, T.D. Noakes, C. Schneider, “Exerciseassociated hyponatremia: a review, Emerg. Med. Australasia 13 (1) (2001) 17
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7224. [19] S. Kano, K. Kim, M. Fujii, Fast-response and flexible nanocrystal-based humidity sensor for monitoring human respiration and water evaporation on skin, ACS Sens. 2 (6) (2017) 828
833. [20] F. Gu¨der, et al., Paper-based electrical respiration sensor., Angewandte Chemie 128 (19) (2016) 5821
5826. [21] S. Kanaparthi, Pencil-drawn paper-based non-invasive and wearable capacitive respiration sensor, Electroanalysis 29 (12) (2017) 2680
2684. [22] C. Deng, et al., Determination of acetone in human breath by gas chromatography
mass spectrometry and solid-phase microextraction with on-fiber derivatization, J. Chromatograp. B 810 (2) (2004) 269
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5352. [24] F. Zhang, et al., Controlled synthesis and gas-sensing properties of hollow sea urchin-like α-Fe2O3 nanostructures and α-Fe2O3 nanocubes, Sens. Actuat. B Chemical 141 (2) (2009) 381
389.

[25] T. Jinkawa, et al., Relationship between ethanol gas sensitivity and surface catalytic property of tin oxide sensors modified with acidic or basic oxides, J. Mol. Catal. A: Chem. 155 (1
2) (2000) 193
200. [26] S.-J. Choi, et al., Novel templating route using Pt infiltrated block copolymer microparticles for catalytic Pt functionalized macroporous WO3 nanofibers and its application in breath pattern recognition, ACS Sens. 1 (9) (2016) 1124
1131. [27] A. Tangerman, E.G. Winkel, Extra-oral halitosis: an overview, J. Breath Res. 4 (1) (2010) 017003. [28] C.L. Whittle, et al., Human breath odors and their use in diagnosis, Ann. New York Acad. Sci. 1098 (1) (2007) 252
266. [29] L. Ciaffoni, R. Peverall, G.A.D. Ritchie, Laser spectroscopy on volatile sulfur compounds: possibilities for breath analysis, J. Breath Res. 5 (2) (2011) 024002. [30] S.-J. Choi, et al., Selective detection of acetone and hydrogen sulfide for the diagnosis of diabetes and halitosis using SnO2 nanofibers functionalized with reduced graphene oxide nanosheets, ACS Appl. Mat. Interf. 6 (4) (2014) 2588
2597. [31] K. Alving, E. Weitzberg, J.M. Lundberg, Increased amount of nitric oxide in exhaled air of asthmatics, Eur. Resp. J. 6 (9) (1993) 1368
1370. [32] J. Morton, L.H. Richard, S.T. Paul, Exhaled breath condensate nitrite/nitrate and pH in relation to pediatric asthma control and exhaled nitric oxide, Pediat. Pulmonol. 41 (10) (2006) 929
936. [33] W.-T. Koo, et al., Catalyst-decorated hollow WO3 nanotubes using layer-by-layer self-assembly on polymeric nanofiber templates and their application in exhaled breath sensor, Sens. Actuat. B Chem. 223 (2016) 301
310. [34] M.K. Nakhleh, et al., Diagnosis and classification of 17 diseases from 1404 subjects via pattern analysis of exhaled molecules, ACS Nano 11 (1) (2016) 112
125. [35] K. Kalantar-Zadeh, et al., A human pilot trial of ingestible electronic capsules capable of sensing different gases in the gut, Nat. Electron. 1 (1) (2018) 79.

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C H A P T E R

8 Aptamer-Mediated Nanobiosensing for Health Monitoring Madhu Malinee1,*, Alok Kumar2,*, Abhijeet Dhiman3,4 and Tarun Kumar Sharma5,** 1

Department of Anatomy and Developmental Biology, Graduate School of Medicine, Kyoto University, Kyoto, Japan, 2Department of Immunology and Genomic Medicine, Graduate School of Medicine, Kyoto University, Kyoto, Japan, 3Department of Biotechnology, All India Institute of Medical Sciences, New Delhi, India, 4Faculty of Pharmacy, Uttarakhand Technical University (UTU), Dehradun, India, 5Center for Biodesign and Diagnostics, Translational Health Science and Technology Institute, NCR Biotech Science Cluster, Faridabad, India

8.1 INTRODUCTION The World Health Organization (WHO) defined health as “a state of complete physical, mental and social well-being and not merely the absence of disease or infirmity” [1]. Achieving a complete state of good health is almost impossible for the majority of world’s population as infectious diseases and noncommunicable diseases pose global health problems despite significant progress in prevention, diagnosis, and treatment [2]. To remain healthy, one’s health status must be monitored regularly. Early diagnosis and timely treatment are primary requisites for good health. Medical errors are the

third-leading cause of deaths in the United States. This critical situation demands improvements in medical diagnosis and treatments to save human lives (Fig. 8.1) [3,4]. The main aim to be achieved in health care is to assess health status and monitor treatment using noninvasive or less-invasive methods with high specificity and precision. According to the WHO report, the major causes of deaths worldwide are cardiovascular diseases (31%), cancer (13%), respiratory diseases (8%), diabetes (2%), and other communicable and noncommunicable diseases [5]. Regular and precise monitoring of health status is necessary to ensure people’s longer and healthier lives.

*Equal contribution. **Corresponding authors

Advanced Biosensors for Health Care Applications DOI: https://doi.org/10.1016/B978-0-12-815743-5.00008-1

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FIGURE 8.1 Most common causes of death in the United States, 2013. Source: Adapted from BMJ 2016; 353:i2139 http://dx.doi. org/10.1136/bmj.i2139 with permission from BMJ Publishing Group Ltd.

Although significant progress has been observed in diagnostic tests and instrumentation over the past three decades, there is still some lacuna in providing fast, highly specific, and onsite diagnostic instruments for accurate and fast diagnoses [6,7]. The diagnostic setup needs to have a fast response time, user-friendly, costefficient, and with no (or less) dependency on trained professionals for operating the instruments. Conventional in vitro diagnostics tests for diseases or health-related issues are timeconsuming and require centralized laboratories, trained technicians, and costly equipment. Currently, the diagnosis of infectious diseases depends on in vitro diagnostic laboratory-based tests that include microscopy, immune-based

assays like enzyme-linked immunosorbent assay (ELISA), culture-based assays, and nucleic acid amplification assays like polymerase chain reaction (PCR); however, all these methods are timeconsuming and lack sensitivity. In recent years several advancements in the biosensor technology has addressed various limitations associated with the conventional sensing/diagnostic methods. This attribute makes biosensors as a true point-of-care detection systems.

8.2 BIOSENSORS A biosensor is an analytical device which incorporates a biologically active element with

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8.2 BIOSENSORS

an appropriate physicochemical transducer to convert the biological response into a measurable signal. Biosensors have potential applications not only in clinical diagnostic systems as its applications, directly or indirectly, cover all the aspects of human lives (e.g., medical science, security services, transport services, food analysis, and environment surveillance). It is beyond the scope of this chapter to cover all these aspects. Therefore we will focus on how biosensors have improved diagnostics in healthcare services.

8.2.1 Different Parts of Biosensors Biosensors are composed of three parts: (1) a biorecognition or bioreceptor element (BRE) which detects and binds with the target in the sample; (2) a transducer that converts the interaction between the recognition element and its cognate target into a measurable signal; and (3) an electronic system that amplifies, processes, and displays the signal [8,9]. Besides these three components of a biosensor, another essential requirement is the immobilization machinery to immobilize the biorecognition element to make the reaction more feasible and efficient with the analyte of interest in the test sample.

8.2.2 Different Types of Biosensors Biosensors can be classified on the basis of the interaction between the biorecognition element with the target, for example: (1) affinity sensors, if the recognition element binds with the target; (2) metabolism sensors, if the interaction is accompanied by a chemical change in accordance with a corresponding change in the analyte in the sample; and (3) catalytic sensors, where the recognition element causes some catalytic change in the test sample and the equivalent reduction in the test sample can be

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detected, transduced, and amplified by the transduction element [10,11]. Biosensors can also be classified on the basis of transduction principles they use for converting the biological response into measurable signals. According to transduction principles, biosensors are: (1) optical biosensors that convert the biological response into visible light signals; (2) electrochemical biosensors that are based on detecting electrochemical changes (potential, current, or impedance) that happen at the interface of the bioreceptor and target molecule; and (3) massbased biosensors that sense the changes in mass due to biological responses at the analyte and bioreceptor interface.

8.2.3 Different Bioreceptor Elements The key requirement of a bioreceptor is high selectivity for the analyte within a matrix of other chemicals or biological components. The BRE is a biologically derived material that interacts with, or is recognized by, the analyte of interest. BREs can be antibodies, enzymes, nucleic acids, microorganisms, biological tissues, cellular structures/cells/organelles, and biomimetic materials. Antibodies bind with their target antigens very strongly and are good candidates for BREs. The limitations of antibodies like high-molecular weight and low stability have been overcome by using artificial recombinant binding fragments (Fab, Fv, or scFv) or domains (VH, VHH) of antibodies. Likewise, enzymes, cellular structures, and biomimetic materials are also possible candidates for BREs. However, aptamers have become increasingly popular over the past decade as a prime choice for BREs due to their advantages over other equivalents. We have described the different BREs in chapter 2 in detail. In this chapter, we will focus mainly on aptamer-based biosensors for healthcare monitoring.

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FIGURE 8.2 Scheme for the systematic evolution of ligands by exponential (SELEX) enrichment process. A random nucleic acid library is incubated with a target molecule, and unbound molecules are separated from bound molecules. Bound nucleic acids are eluted, amplified by polymerase chain reaction (PCR) and serve as an enriched library for the next cycle. For every target, 6
12 consecutive cycles are performed and the final enriched library is cloned and sequenced. Source: Adapted from Trends in Analytical Chemistry, vol. 27, No. 2, 2008 with permission from Elsevier.

8.2.4 Aptamers: The Bioreceptor Element (BREs)

8.2.5 Merits of Aptamers Over Antibodies

Aptamers are single-stranded DNA or RNA nucleotide sequences that bind with a variety of targets with high affinity and selectivity (having dissociation constants, Kd, in the range of 10212
1026 M) which makes them ideal biorecognition elements for biosensors [12,13]. The word aptamer is derived from the Latin word “aptus” meaning “to fit.” In 1990 the Gold lab and the Szostak lab independently developed a technique for in vitro selection and amplification of aptamers, namely the systematic evolution of ligands by exponential enrichment (SELEX) (Fig. 8.2) [14,15].

Aptamers and antibodies are affinity reagents that have specificity for their target. However, aptamers are superior to antibodies as they are stable in a wide range of temperatures, pH and other storage conditions, lower batch-to-batch variability, smaller in size, not immunogenic as they are very small molecules, can be modified easily, can be produced in vitro within a short period of time unlike antibodies which requires a host for immunization and trained professionals for production [16]. The differences between aptamers and antibodies are summarized in Table 8.1.

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TABLE 8.1 Comparison Between Antibodies and Aptamers. Feature

Antibody

Aptamer

Nature

• Classified as an affinity molecule • Protein molecule (monoclonal or polyclonal)

• Classified as an affinity molecule • Nucleotides sequence (single stranded DNA or RNA oligonucleotide)

Molecular weight

• 100
150 kDa

• 5
25 kDa

Target range

• Can be generated against immunogenic molecules only (proteinaceous large molecules)

• Can be selected against a wide range of targets (small molecules like metal ions to large molecules like proteins and even whole cells)

Production method

• Antibody generation requires host animals and immunization process to develop (in vivo production)

• Can be generated in vitro via the SELEX process.

Target site

• Host’s immune system decides the target site against which the host body will make antibodies

• Investigator decides the target sequences against which aptamers has to be synthesized

Batch-to-batch variation

• Are prone to batch-to-batch variation due to different host and different kinetics

• Less prone no batch-to-batch variation

Stability and storage conditions

• Are stable in a short range of conditions • Very sensitive to degradation due to proteinaceous nature • Temperature sensitive and can undergo irreversible denaturation

• Are stable up to a wide range of storage conditions • Very stable owing to nucleic acid nature • Regain structure and function when temperature become normal

Chemical modification

• Modification may affect the affinity of the antibody so limited modification possible

• Can be modified easily without compromising affinity and selectivity

Immunogenicity nature

• May be immunogenic to patient due to proteinaceous nature

• Nonimmunogenic as they are small molecules

Cost of production

• Production is very costly as it depends upon host animals

• Can be produced very cost-effectively

Production time

• Takes months in production as it depends on the host animal and immunization with the respective antigen

• Can be produced within a week

Affinity (Kd value)

• 1029 M

• 1026
10212 M (with average being 1029 M)

Specificity

• High

• High

SELEX, Systematic evolution of ligands by exponential enrichment.

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Aptamers can be selected from small molecules like ions, amino acids, vitamins, drugs, and peptides to large molecules like dyes, proteins, enzymes, and even whole cells.

8.2.6 Nanobiosensing: Improving Efficacy of Aptasensors in Combination With Nanomaterials Aptamers undergo conformational changes upon binding with the target and this property can be utilized for designing biosensors with high-detection sensitivity and selectivity. Aptamers can incorporate small molecules like nanoparticles in their structural folding or can be integrated into the structure of large molecules such as proteins. Nanomaterials are the particles with its dimension in the range of 1
100 nm [17]. The size constraints of nanomaterials make them very special as they have high surface area per unit of volume ratio that leads to chances of most of their constituent atoms to be located at or near their surface and so their physicochemical properties are significantly different as they follow quantum mechanisms at the nanoscale level. They can play very efficient roles in the sensing mechanism of biosensor technology. Some nanomaterials are widely used in nanobiosensing, for example, nanotubes, nanowires, nanorods, nanoparticles, and thin films which are made up of nanocrystalline matter. Nanomaterials have unique properties like high electrical conductivity, better shock-bearing ability, and sensitive responses such as piezoelectric and color that changes with variations in size and shape of the nanoparticles. These features make them suitable for conjugation with aptamers for biosensing applications. The incorporation of various nanomaterials, for example, metal ions, carbon, gold nanoparticles, and nanospheres, has improved the analytical performance (faster detection and more specific) and commercial application of aptasensors. Their

electromechanical, optical, and mechanical properties are the assets utilized for the development of nanobiosensors (nanoparticle conjugated aptamer-based biosensors). Nanotechnology permits the development of nanoscale biosensors with high sensitivity and versatility and allow the detection of biochemical and biophysical signals associated with health-related issues (communicable as well as noncommunicable diseases including cancer) at the level of a single molecule or cell. Lab-on-a-chip came into reality due to nanotechnology advancements. Their ability to detect disease-associated biomolecules (e.g., disease-specific metabolites, nucleic acids, proteins, pathogens) and cells (e.g., circulating tumor cells) and easy portability due to the reduced instrumentation size makes nanobiosensors ideal candidates for onsite monitoring (as well as for laboratory settings). Nanobiosensing technology, by enabling early diagnosis of chronic illnesses and the precise detection of pathogens, have the potential to transform conventional medical practices. For monitoring health, the glucose biosensor developed by Clark and Lyons in 1962 was the first approach for detecting glucose levels in the blood via potentiometric measurement coupled with the enzyme glucose oxidase [18]. After that, there was a huge surge in the development of biosensors for detecting different metabolites, pathogenic infections, and other health-related issues that led to biosensor technology.

8.3 TYPES OF NANOBIOSENSORS Nanobiosensors can be classified based on their transduction method as well as their conjugated nanoparticles. According to transduction mechanisms, nanobiosensors can be categorized into: (1) optical nanobiosensors; (2) electrochemical nanobiosensors and (3) mass-sensitive nanobiosensors. According to

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FIGURE 8.3 (A) Label-free biosensors versus (B) label-dependent biosensors. Source: Adapted from Sensors 2016, 16(12), 2178.

the nanoparticles conjugated, nanobiosensors can be categorized into: (1) metallic nanoparticles based on, for example, gold nanoparticles (GNP), magnetic nanoparticles, gold nanorods, among others; and (2) carbon and silica nanomaterial-based nanobiosensors. The nanobiosensors can also be classified according to the label as either label-free or labeled (Fig. 8.3). Label-free nanobiosensors can measure the analyte concentration directly through biochemical reactions on the transducer surface, while in the case of labeled nanobiosensors, the analyte is in between the capture and detection reagent where the detection reagent is attached with a specific label (e.g., fluorophore, enzyme, radioisotopes, and quantum dots) [19].

8.3.1 Label-Free Nanobiosensors

mode is basically important for smallmolecular targets that can occupy in the binding pockets of the BRE. Label-free detection methodology depends on the intrinsic properties of the target molecule. Some of their advantages are: (1) allowing quantitative measurements of molecular interaction in real time; and (2) analytes being detected in their natural form without any modification, which is very important for further downstream analysis. Based on the transduction mechanism, they can be classified into three groups: (1) optical biosensors that are used to detect fluorescence, colorimetric changes, or luminescence; (2) electrochemical biosensors that are used to detect enzymatic reactions at the interface of the BRE and the target molecule; and (3) mechanical biosensors which are used to detect changes in mechanical properties at the interface.

Unique features of label-free nanobiosensors is that they require less reaction time and are very sensitive. Label-free biosensors have a simplified assay design as they require only a biorecognition element and so have reduced assay time for the detection of the analyte. Label-free nanobiosensing depends on changes or reactions that occur at the interface of the BRE and the target molecule. This detection

8.3.1.1 Optical Nanobiosensor Due to their high sensitivity, optical transductions are one of the widely used transduction methods for biosensing. Optical biosensors utilize the principle of optical measurements, for example, absorbance, scattering, interferometry, surface plasmon resonance (SPR), fluorescence, chemiluminescence, and

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others. They employ the use of fiber optics and optoelectronic transducers. Most of the wellknown, label-free optical nanobiosensing techniques are based on SPR, waveguide, and interferometric transduction mechanisms, among others. The SPR phenomenon occurs when polarized light hits a metal film at the interface of media with different refractive indices. SPR technology was first demonstrated by Leidberg et al., in 1983 [20]. Among label-free, optical aptamer-based nanobiosensors, SPR has shown great promise as they provide label-free detection and real-time quantitative analysis as well. In another report, Laura et al. investigated the biosensing of RNA with a DNA-triplex affinity capture method. They targeted RNA sequences with secondary structures [21]. The approach was based on selecting DNA tail-clamps as affinity

bioreceptors that have the ability of create a triplex helix with target RNA using the SPR detection method with limit of detection of 50 fmol (Fig. 8.4). Murphy et al. have shown the detection of thyroid transcription factor 1 (TTF-1) using SPR technology-based aptasensors [22]. TTF-1 is a nuclear transcription factor, expressed in normal lungs, thyroids, and their neoplasms. TTF-1 is a specific marker for adenocarcinomas of the lung in body cavity fluids [23]. In this work, the researchers utilized SPR employing a BIAcore X instrument to measure the affinity of the interaction of TTF-1 with aptamers immobilized on a sensor chip. In the work by Park et al., the authors demonstrated the detection of nicotinamide phosphoribosyltransferase (NAMPT) proteins. NAMPT, also known as pre-B-cell colony-enhancing factor 1

FIGURE 8.4 (A) Design of the amino-modified tail-clamp bioreceptor used in this study. The Watson
Crick and Hoogsteen forming strands are highlighted in blue and green, respectively. The vertical spacer (dark blue) and aminoadenines introduced in some tail-clamps receptors are also indicated. (B) Predicted structure with minimum-free energy folding (MFE) for Listeria innocua RNA target obtained from The Vienna RNA Website. The matching region with the tail-clamp receptor is highlighted in purple. Available via license: CC BY-NC 3.0.

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TABLE 8.2 List of Label-Free Optical and Mechanical Nanobiosensors. Method

Target

Limit of Detection

References

SPR

Tubulin

10 μM

[28]

SPR

Thyroid transcription factor 1 (TTF1)

67 nM

[22]

SPR

HIV-1 Tat protein

0.12 ppm

[29]

SPR

C-reactive protein

0.005 ppm

[30]

SPR

Human IgE

141 ng/mL

[31]

SPR

Retinol binding protein 4 (RBP4)

1.58 mg/mL

[32]

SPR

Adenosine

0.21 pM

[33]

SPR

Interferon gamma

10 pM

[34]

SPR

Influenza virus




[35]

BI

Human IgG




[27]

[36]

OPTICAL BIOSENSOR:

MECHANICAL BIOSENSORS: Cantilever based

Escherichia coli




QCM

HIV-1 gp41

2 ng/mL

QCM

Influenza A and B viruses

3

10 cfu/mL

[37] [38]

SPR, Surface plasmon resonance; BI, Backscattering interferometry; QCM, Quartz crystal microbalance.

(PBEF1) or visfatin, is an enzyme belonging to the family of glycosyltransferases, or to be more specific, the pentosyltransferases. This enzyme participates in nicotinate and nicotinamide metabolism. They used graphene oxide (GO) as a label-free platform for screening of the aptamer since GO has the ability to separate free short ssDNA in the heterogeneous solution. [24]. Schofield and Dimmock et al. detected influenza-A virus using SPR technology [25]. Here, kinetic binding constants for virion/IgG interaction could be determined in their SPRbased detection technology. Another approach in optical detection is backscattering interferometry (BI). In BI, a laser is focused onto a microfluidic channel which causes a highly modulated interference pattern and a detector

analyzes the changes in the profile of fringe patterns that shows the quantification of molecular binding events [26]. This technology allows real-time determination and is capable to detect both free-solution or surfaceimmobilized molecular interactions with the limit of detection in the pM range. Utilizing this method, Kussrow et al., have detected human IgG from seropositive syphilis patients using a purified recombinant treponemal antigen r17 [27]. Optical nanobiosensors are listed in Table 8.2. 8.3.1.2 Electrochemical Nanobiosensors Due to high sensitivity and simplicity, electrochemical transducers are also a commonly used approach like optical transducers for biosensing. These nanobiosensors measure the

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electrochemical changes, for example, electric current (amperometric), potential (potentiometric), and ionic or conductance changes (impedimetric) at the interface of the BRE and the target molecule. Amperometric biosensors depend on the flow of electrons as a result of redox reactions. In the enzyme-catalyzed redox reaction, the electron transfer happens from substrate or product to the electrode surface to be oxidized or reduced, and this electron flow (i.e., current) can be measured. The magnitude of the current will be proportional to the substrate concentration in the test sample. The first of its kind, the glucose biosensor developed by Clark et al. in 1962, is the best example of an amperometric biosensor. In this simple, first-generation glucose biosensor, the glucose oxygenase enzyme catalyzes the oxidation of glucose by molecular oxygen into glucuronic acid, and the reduction in oxygen concentration (determined by the electrode) is proportional to the glucose concentration [18]. Further, in the second-generation glucose biosensor, a mediator like ferrocene, ferricyanide, quinines, tetrathialfulvalene (TTF), tetracyanoquinodimethane (TCNQ), thionine, methylene blue, and methyl viologen was used which take up the electron and then transfer it to the electrode to sense performance [39
42]. In potentiometric nanobiosensors, the ionsensitive electrode determines changes in ionic

concentrations. Many enzymatic reactions produce hydrogens ions, ammonia, water molecules, and carbon dioxide as by-products during the reaction. The most common electrode is a pH-sensitive electrode. Similarly, CO2-sensitive or NH3-sensitive electrodes can determine the potential difference caused due to the differences between the reference and the sensor electrode. The potential difference will be proportional to the concentration of the target molecule in the sample. Although potentiometric nanobiosensors are sensitive, the major problem is the sensitivity of the enzyme to the produced hydrogen and ammonium ion concentrations. Ion-selective field effect transistors (ISEFT) are a low-cost device being used for the miniaturization of potentiometric biosensors. For example, during open-heart surgery, intramyocardial pH can be monitored using ISFET nanobiosensors [43,44]. Electrochemical impedimetric spectroscopy (ESI) studies the interfacial properties related to biorecognition events that occur at the modified surfaces. In impedimetric aptasensors, binding of the target sequence, DNA damages, or conformational changes is followed by the impedance change. The advantage of impedimetric biosensing is that it allows the detection of biomolecular recognition events without using enzyme labels. Various label-free electrochemical nanobiosensors are summarized in Table 8.3.

TABLE 8.3 Label-Free Electrochemical Nanobiosensors. Transduction Method

Target

Enzyme/ Aptamer

Immobilization/Process

Limit of Detection

References

Amperometric Glucose

Glucose Plasma-polymerized thin films dehydrogenase




[45]

Amperometric Glucose

Glucose
dehydrogenase




[46]

Amperometric Glucose

Glucose oxidase

Ferrocene mediated electron transfer (competitive immunoassay)




[47]

Amperometric Glucose

Glucose oxidase

Cross-linked redox gels containing enzyme




[48] (Continued)

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TABLE 8.3 (Continued) Transduction Method

Enzyme/ Aptamer

Immobilization/Process

Limit of Detection

References

Amperometric Glucose

Glucose oxidase

Stable charge transfer complex electrode (TTF 2 TCNQ)




[49]

Amperometric Thrombin

Glucose oxidase

Multifunctional polymeric bionanocomposites (PBNCs)

0.1 nM

[50]

Impedimetry

DNA aptamer

Photolithographic gold surface for aptamer

0.1 nM

[51]

Amperometric Dopamine

RNA aptamer

RNA aptamer tethered to cysteamine-modified gold electrodes via the alkanethiol linker.

62 nM

[52]

Amperometric Cytokeratins antigen 21-1




Poly(thionine)
Au composite

4.6 fg/mL [53]

Potentiometric ATP




Polycation-sensitive electrode

[54]

Potentiometric Tryptophan

Apt-MWCNTAuSPE




4.9 pM

[55]

impedimetric

Thromnin




Chemisorption

2 nM

[56]

impedimetric

IgE




Chemisorption

2.5 nM

[51]

impedimetric

Interferon gamma




Chemisorption

0.1 pM

[57]

impedimetric

Adenosine




Chemisorption

0.1 μM

[58]

impedimetric

AMP




Covalence

10 μM

[59]

impedimetric

Neomycin B




Covalence

1 nM

[60]

impedimetric

PDGF




Covalence

40 nM

[61]

ISFET

Thrombin




Covalence

10 nM

[62]

ISFET

IgE




Covalence

0.25 nM

[63]

ISFET

AMP




Covalence

50 μM

[64]

Amperometric Thrombin




Chemisorption

3 nM

[65]

Amperometric ATP




Chemisorption




[66]

Amperometric PDGF




Chemisorption




[67]

Amperometric Adenosine




Covalence

10 nM

[68]

Amperometric Ghrelin




Electrochemical adsorption




[69]

Amperometric 17-beta estradiol




Affinity




[70]

Target

Human IgE

MWCNT, Multiwall carbon nanotube; SPE, Screen-printed electrode; ISFET, Ion-selective field effect transistors; Poly(thionine)-Au is a novel multifunctional substrate.

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FIGURE 8.5 Sensing mechanism of the CMOS DNA cantilever sensor. Source: Adapted from IEEE Transactions on Biomedical Circuits and Systems, 2013, Vol. 7, No. 6, December 2013 with permission from IEEE.

8.3.1.3 Mechanical Transducer
Based Nanobiosensors Among different formats of mechanical biosensors, cantilever and quartz crystal microbalances (QCMs) are the most well-known techniques (Fig. 8.5). In cantilever-based biosensors, there are three methods to transduce the reaction between biorecognition element and target molecule into a micromechanical motion: (1) the frequency change due to additional mass loading where the cantilever is used as a microbalance; (2) the bending of a bimetallic cantilever can be used as a temperature sensor; and (3) measuring the bending due to changes in the surface stress at one side of the cantilever where cantilever can work as stress sensors [71]. Madar et al. demonstrated the application of cantilever-based sensing for the detection of pathogenic Escherichia coli using a cantilever array which was functionalized with carbohydrate molecules as capture agents [36]. Quartz crystals are piezoelectric crystals and they vibrate at a specific frequency when electrical signals are applied to it. In piezoelectric detection (such as a QCM), changes in surface-adsorbed mass due to the interaction of the BRE with the target molecule causes variations in the resonant frequency of the oscillating quartz crystal. Peduru Hewa et al. showed the detection of influenza A and B viruses using a QCM-based sensor with

detection limit 103 pfu/mL [38]. Their results were comparable to culture-based techniques and better than ELISA in terms of sensitivity and specificity. In another work, Lu et al. detected HIV-1 glycoprotein gp41, an indicator of disease progression and therapeutic response, using a biomimetic sensor based on epitope-mediated imprinting with a detection limit of 2 ng/mL [37].

8.3.2 Labeled Nanobiosensors Labeled nanobiosensing is a robust method compared to label-free biosensing. In this methodology, the analyte is sandwiched between the BRE/capture agent and the detection agent. In label-free biosensing, there is no amplification of the signal (see Fig. 8.3). Capture agents are immobilized on the solid surface while detector agents are conjugated with fluorophore, enzymes, nanoparticles, and others. Similarly to label-free nanobiosensors, these biosensors can be classified according to the transduction method they employ like optical, electrochemical, among others. Labeling of the aptamer involves either covalent or noncovalent binding. Nanoparticles (NPs) can be conjugated with an aptamer (the BRE) where the DNA aptamer can tightly bind to the NPs and stabilize them against salt-induced aggregation. Conjugating aptamers on various nanomaterials leads to

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8.3 TYPES OF NANOBIOSENSORS

enhancements in sensitivity and selectivity of aptasensors. Owing to NPs’ unique properties like size- and shape-dependent optical properties and catalytic activities, they are very useful for the signal generation as well as signal amplification in conjugation with aptamers. Numerous nanoparticles were tested in conjugation with aptasensors like metallic nanoparticles (e.g., gold nanoparticles), magnetic

239

nanoparticles (e.g., Fe3O4), quantum dots, silica nanoparticles, carbon nanotubes (e.g., singlewalled carbon nanotubes, graphene), and others. [72
74]. Gold nanoparticles (AuNPs) are one of the most favored nanomaterials for conjugation with aptasensors due to its unique physical and chemical properties. Aptamers have been conjugated on GNP by physical adsorption or chemisorption (Fig. 8.6).

FIGURE 8.6 DNA aptamer could tightly bind to AuNPs and stabilize them against salt-induced aggregation. (A) Upon binding to its target, the aptamer formed 3D structures and led to the aggregation of AuNPs; (B) in the presence of target molecules, purple colored DNA-modified AuNP aggregates were dispersed into red-colored individual AuNPs; and, (C) upon binding the target, AuNPs functionalized with short, complementary ssDNA and aptamers were unstable against salt-induced aggregate, leading to a red-to-purple color change. Source: Adapted from Trends in Analytical Chemistry, Vol. 27, No. 2, 2008 with permission from Elsevier.

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FIGURE 8.7 Colorimetric aptasensors using modified (A
C) and unmodified (D and E) AuNPs. (A) Sandwich-type colorimetric aptasensors using dual aptamers; (B) target-induced displacement of aptamer from AuNPs network; and aggregation of AuNPs via surface condition using (C) DNA folding effect; (D) target-induced conformational change of aptamer; and (E) target-induced assembly of aptamer fragments. Source: Adapted from Biosensors and Bioelectronics 76 (2016) 2
19 with permission from Elsevier.

8.3.2.1 Labeled Optical Nanobiosensors Optical sensors are used to detect fluorescent, luminescent, and colorimetric tags. Out of these, colorimetric detections systems are fascinating due to its cost-effectiveness and simplicity which make them a candidate for being used in point-of-care tests. One research group showed GNP-based aptasensors for cancer cell detection, the CCRF-CEM (human acute lymphoblastic leukemia), using aptamer-conjugated AuNPs which selectively

bound to the surface of target cells, but this could further be useful for imaging instead of detection due to it slow sensitivity (Fig. 8.7). The different strategies for use of GNP in conjugation with colorimetric sensing are well reviewed by another group [75]. In fluorometric sensing, the detection is based on the fluorescence phenomenon as a result of the interaction between the BRE and the target molecule. Many different types of fluorescence and quenching materials can be conjugated with

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241

FIGURE 8.8 Fluorescent aptasensors using AuNPs as fluorescence quenchers. Fluorescent assay using modified AuNPs based on: (A) target-induced displacement fluorescent dye-labeled complementary DNA; (B) target-induced displacement of aptamer from nonlabeled complementary DNA; and (C) target-induced conformational change of aptamer and detached from unmodified AuNPs. Source: Adapted from Biosensors and Bioelectronics 76 (2016) 2
19 with permission from Elsevier.

aptamers (Fig. 8.8). Fluorometric sensing is a more encouraging methodology compared with colorimetric-based sensing in qualitative and quantitative analyses. Different colorimetric and fluorometric sensing systems are summarized in Table 8.4. 8.3.2.2 Nanozyme-Based Turn-off/Turn-on Approach for Aptamer Health Monitoring Sharma and Weerathunge et al. introduced a novel approach based on the peroxidase-like NanoZyme activity of GNPs and molecular recognition ability of aptamers that allows highly sensitive detection of small molecules like kanamycin and pesticide (acetamiprid) [96]. In this assay, when the target is absent, the peroxidase activity of the pristine GNP was inhibited by shielding its surface through the adsorption of acetamiprid-specific aptamer and, hence, no color change is observed on addition of 3,3,5,5-tetramethylbenzidine (TMB) substrate. On the other hand, in the presence

of aptamer-cognate target, the peroxidase activity of the pristine GNP was retained as the aptamer exhibited target-responsive structural changes which further allowed its desorption from the GNP surface, resulting in the color change of colorless TMB to a purplish-blue product in the presence of H2O2 (Fig. 8.9). The output of this NanoZyme activity of GNPs can either be directly visualized in the form of color change of the peroxidase reaction product or can be quantified using UV
visible absorbance spectroscopy [96,97]. As a result of GNPs exhibited peroxidaselike NanoZyme activity, several other researchers have exploited this method in detecting various biological analytes like proteins and small molecules such as thrombin and Pb21 with detection limits of 2.5 nM, 0.02 pM, and 602 pM, respectively [98,99]. By using this approach, researchers are able to achieve highly sensitive, rapid, and cost-effective assays for the detection of various analytes.

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TABLE 8.4 Labeled Aptamer-Based Optical Nanobiosensor. Detection Method

Target

Material (Nanoparticles)

Limit of Detection

OPTICAL: Colorimetric

Thrombin

GNP

2.5 nM

[76]

Colorimetric

PDGF

GNP

10 nM

[77]

Colorimetric

Adenosine

GNP

20 μM

[78]

GNP

250 nM

[79]

Colorimetric

21

Hg

21

References

Colorimetric

Pb

GNP

500 nM

[80]

Colorimetric

Oxytetracycline

GNP

25 nM

[81]

Colorimetric

Glutathione

GNP

17 nM

[82]

Colorimetric

ATP

GNP

0.6 μM

[83]

Colorimetric

Plasmodium vivax lactade dehydrogenase (PvLDH)

GNP

10.3 pM

[84]

Colorimetric

VEGF

GNP

30 nM

[85]

Colorimetric

Carcionoembryonic antigen (CEA)

MNP

1 ng/mL

[86]

Fluorometry

Theophylline

GNP

10 μM

[87]

Fluorometry

Cocaine

GNP

1 μM

[88]

Fluorometry

Lysozyme

Graphene

80 ng/mL

[89]

Fluorometry

Cellular prion protein

Graphene

0.3 μg/mL

[90]

Fluorometry

Prostate-specific antigen (PSA)

Graphene oxide

0.5 ng/mL

[91]

Fluorometry

Hg

Quantum dots

0.5 μM

[92]

Fluorometry

Thrombin

Silica nanoparticle

1 nM

[93]

SERS

Thrombin

Gold nanorods

887 pM

[94]

Fluorometry

Thrombin

SWCNTs

1 μM

[95]

21

GNP, gold nanoparticles; MNP, magnetic nanoparticles; SERS, surface-enhanced Raman Spectroscopy; SWCNT, single walled carbon nanotubes.

8.3.2.3 Labeled Electrochemical Nanobiosensors Nanoparticles (especially GNPs) have the potential to amplify electrochemical signals based on properties of NPs due to their large surface area and high redox activity, which makes them good candidates for use in conjugation with aptamers. The use of nanoparticles for

electrochemical biosensors was reported for the first time in 2002 when Musameh et al. reported the detection of low-potential stable NADH on carbon nanotube-modified electrodes [100]. After this study, there has been huge progress in the field of labeled-electrochemical nanobiosensors. Various electrochemical nanobiosensors are summarized in Table 8.5.

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8.3 TYPES OF NANOBIOSENSORS

FIGURE 8.9 Schematic representation showing inhibition of NanoZyme activity of GNPs using an acetamiprid-specific aptamer (Step A) shows intrinsic peroxidase-like activity of pristine GNP that gets inhibited after conjugating with acetamiprid-specific aptamer (Step B), in the presence of acetamiprid the aptamer undergoes target responsive structural changes and forms a supramolecular complex with acetamiprid resulting in free GNP to resume its peroxidase like activity (Step C). Source: Adapted from Sensors and Actuators B: Chemical 246, (2017), 535
553 with permission from Elsevier.

TABLE 8.5 Labeled-Electrochemical Aptamer-Based Nanobiosensors. Detection Method

Target

Nanoparticles

Limit of Detection

References

Electrochemical

Thrombin

GNP

100 pM

[101]

Electrochemical

Cocaine

GNP

10 μM

[102]

Potentiometric

Salmonella typhi

SWCNT

0.2 cfu/mL

[103]

Potentiometric

Staphylococcus aureus

SWCNT

800 cfu/mL

[104]

Electrochemical

ATP

Graphene oxide

0.1 nM

[105]

Electrochemical

Thrombin

Graphene

0.03 nM

[106]

Electrochemical

K1

Graphene

27 μM

[107]

Electrochemical

PDGF

Graphene

8 pM

[108]

Electrochemical

Thrombin

Magnetic nanoparticles

7.82 aM

[109]

Potentiometric

NADH

CNT




[100]

Electrochemical

PDGF

GNP

0.01 pM

[110]

GNP, gold nanoparticles; SWCNT, single-walled carbon nanotubes.

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FIGURE 8.10 “Sandwich” sensing system for PDGF detection using GNP-labeled electrochemical nanobiosensor. Source: Adapted from Biosensors and Bioelectronics 24 (2008), 1598
1602 with permission from Elsevier.

In another report, by Wang et al., GNPs were used for signal amplification in an electrochemical aptasensor for the detection of platelet-derived growth factor (PDGF) [110]. In this report, a sandwich structure based on selfassembled thiolated aptamers on gold reacting with PDGF and then with aptamer-loaded GNPs labeled signaling aptamer was designed. [Ru(NH3)5Cl]2 was used as the electrochemical probe that could electrostatically interact with numerous strands of aptamers linked to NPs that give large peak currents. This system could detect as low as 0.01 pM of PDGF in the human serum (Fig. 8.10).

8.4 CONCLUSION Over the past few decades, have shown tremendous growth their unique ability to adapt to diagnostic platforms. According

aptasensors by virtue of the various to a recent

market analysis the diagnostics market has grown up to US$16.5 billion. However, the accurate contribution of aptamers to the diagnostics industry has not been estimated to date. The overall aptamer industry is continually growing and is predicted to reach US $244.93 million by 2020, up from US$107.56 million in 2015. More than 50 companies are actively engaged across the globe in aptamerbased research with the primary focus being diagnostics or therapeutics. We expect that the aptamer-market will surge in the near future as they are rivaling the traditional antibodybased assays efficiently. However, to create such disruption, aptamer developers should consider the need of end-users carefully before planning the aptamer development so that they can achieve a target-product profile. Further, a targeting niche market like the food sector where quality control requires extensive food testing to ensure that food is pathogen or endotoxin free can help aptamers penetrate the

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REFERENCES

market. Also, adaption of aptamers to rapid diagnostic/biosensing platforms (such as an electrochemical sensing platform) where antibodies fail or poorly perform can also offer an edge over antibodies. Taken together aptamer technology has an impressive growth trajectory and in the near future aptamers will be able to replace antibodies to be frontline diagnostic tests.

COMPETING FINANCIAL INTEREST The authors declare no competing financial interests.

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C H A P T E R

9 Biosensing
Drug Delivery Systems for In Vivo Applications M. Birgul Akolpoglu1, Ugur Bozuyuk1, Pelin Erkoc2 and Seda Kizilel1,* 1

Chemical and Biological Engineering Department, Koc University, Istanbul, Turkey, 2College of Engineering and Natural Sciences, Bahcesehir University, Besiktas, Turkey

9.1 INTRODUCTION Drug delivery systems have been extensively investigated for the treatment of various chronic diseases, including cardiovascular disease, diabetes mellitus, hypertension, asthma, and cancer [1]. These chronic illnesses mostly require regular drug administration. However, drug delivery systems that can sense biological cues and, subsequently, release their therapeutic payloads could have a transforming impact on the sensitive control and effective management of such diseases. Therefore platforms which combine biosensors and drug delivery can pave the way for more effective treatment strategies for biomedical applications. Diabetes is a disease characterized by metabolic disorder and diabetic people are either insufficient in producing insulin to adjust their blood glucose level (Type I diabetes), or their cells are not responding to insulin (Type

II diabetes) [2]. Biosensor-integrated drug delivery systems are commonly studied for the treatment of diabetes. In addition to diabetes, upregulation of cholesterol is considered to cause cardiovascular diseases. Controlling the level of cholesterol is essential for the prevention of serious damage to the cardiovascular system. A practical biosensor for the treatment of such illnesses is the one that merges fast responsiveness and ease of administration with biocompatibility [3]. Biosensors are frequently used for the management of the chronic diseases. Biosensors detect the amount of specific analyte concentrations and are not originally equipped with therapeutics to treat diseases. In the literature, there are studies that merge biosensing and drug delivery concepts. These unique systems are specific classes of biosensors designed for continuous analysis of biological media and concurrent release of a therapeutic in response to relevant metabolite

*Corresponding authors

Advanced Biosensors for Health Care Applications DOI: https://doi.org/10.1016/B978-0-12-815743-5.00009-3

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levels. Realization of these systems is only possible when they maintain their functionality in complex physiological environments, as in the human body. The ultimate goal of biosensing
drug delivery system design is to combine diagnosis and treatment to treat chronic diseases more efficiently. This chapter begins with the in vivo requirements for biosensing
drug delivery systems. The obstacles to realization of these systems are then addressed in detail. Materials and devices are categorized and their advantages and disadvantages are discussed along with their possible in vivo applications. Finally, case studies of exclusive research findings are explained according to different disease types.

9.2 ADAPTATION OF BIOSENSING
DRUG DELIVERY SYSTEMS TO IN VIVO APPLICATIONS Biosensors that can perform in vivo are powerful tools for healthcare research. Material selection plays an important role as a design parameter due to continuous measurement of the analyte in the surrounding tissue of biosensor. Therefore biosensors should be made from biocompatible and noncytotoxic materials. In addition, the characteristics of biosensors should be comprehensively assessed for a biosensingbased drug delivery system with maximum performance. In this section, we discuss several design criteria such as biocompatibility, selectivity, and sensitivity for an ideal biosensor for in vivo drug delivery applications.

9.2.1 Biocompatibility Biocompatibility of a biosensor is a bilateral term. A biocompatible biosensor causes minimal perturbation in the in vivo environment. Likewise, in vivo environment (i.e host body)

does not interfere with the performance of the biosensor [4]. Whether it is a therapeutic biomaterial or a biosensor, an immune response is often created in the host body when it encounters a foreign material. In vivo biosensors lose their sensitivity by 50% compared to their initial in vitro performance [5,6]. This loss of activity is initially attributed to acute inflammatory response, which takes place immediately after the biosensor enters to the in vivo environment [7]. In this first stage, the biosensor is perceived as a foreign body and plasma proteins and inflammatory cells rapidly migrate to the surface of the material. Plasma proteins are adsorbed on the surface of the embedded biosensor and then phagocytic cells including neutrophils, monocytes, and macrophages start to cover up and strive to destroy this foreign material [8]. Eventually, a foreign body capsule is formed that acts as a barrier separating the implant from in vivo environment (Fig. 9.1). Therefore sensor performance may significantly decrease, since the detection elements of the biosensor will no longer be in contact with the metabolites. The adsorption of proteins on a biosensor surface or infiltration of biomolecules into a biosensor, also known as biofouling, is one of the major causes for foreign body response [9]. However, for a practical in vivo drugrelease application, the biosensor should be stable enough to sense the concentration of metabolite and release its cargo without interruption. Biofouling creates a diffusion barrier for therapeutics in the biosensor, which may ultimately cause unpredictable release behavior. Other than biofouling, biosensor degradation is an important factor and is also driven by inflammatory reactions. Degraded biosensors not only lose their sensitivity or activity, but also become defenseless in response to immunological attacks from the host body [8]. Additionally, for a therapeutic-loaded biosensor, degradation may trigger the early release of the therapeutic, which eventually causes loss of efficiency. To overcome the problems

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Proinflammatory Nonspecific serum protein adsorption

Immune cell infiltration

Antiinflammatory

Macrophage classical activation (M1)

Monocyte

Macrophage alternative activation (M2)

Immune cells

Macrophage fusion

Fibrous encapsulation

Fibroblast

i cru Re t

Neutrophil Ad

IL-10, IL-4, IL-13, TGF-β

re

he

TNF-α, IL-1β, IL-6, MIP-1

Macrophage

After implantation

Hours/days

Biomaterial Week(s)

Month(s)

FIGURE 9.1 Schematics representing host responses to an implant. Source: Reproduced with permission from Kim et al., Biomolecular strategies to modulate the macrophage response to implanted materials, J. Mater. Chem. B Copyright 2018.

related to foreign body response for improved in vivo functionality and durability of a biosensing
drug delivery system, several effective approaches have been reviewed. Biocompatible-material coating is one of the frequently utilized approaches to minimize biofouling. This is achieved by covering the biosensor surface with a biocompatible and hydrophilic material so that a water layer at the interphase can prevent adhesion of proteins and other biomolecules onto the biosensor surface [10]. Numerous natural polymers such as alginate, chitosan, collagen, dextran, and hyaluronan are considered as coating materials due to their nonbiofouling properties [11
14]. Polyethylene glycol (PEG) is widely preferred as a synthetic polymer to improve biocompatibility of a material due to its highwater solubility. For example, PEGylation of biosensors, which release insulin in response to glucose concentration in the blood, increased the durability of devices and improved their function as a sense-and-release mechanism in in vivo experiments [15,16]. Another approach for eliminating foreign body response is to incorporate antiinflammatory drugs into the biosensing systems [10].

For instance, glucocorticoids, a class of steroid hormones, act as an inhibitor that decreases the release of inflammatory cells toward the implantation site. Consequently, the biofouling effect is repressed since capillary permeability is decreased and fibroblast proliferation is inhibited at the interphase between the biosensing material and the surrounding tissue [17]. However, these systems require continuous administration of antiinflammatory drugs, which might cause side effects in the host body. Therefore sustained and localized delivery of drugs at the biosensor implantation site is desired. Drug molecules embedded within hydrogel matrices and drug-loaded biomaterial coatings have been studied as alternative approaches in various studies [18,19]. Moreover, angiogenic drugs can also be considered in biosensor systems, since a vascular network formed around biosensor would improve blood and analyte transport and, thereby, eliminate the possibility of formation of a fibrous capsule without vasculature [10]. This is usually achieved by introducing growth factors to the system, such as the vascular endothelial growth factor (VEGF) [18,20]. However, the use of antiinflammatory drugs

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interferes with endogenous VEGF and hinders angiogenesis [21]. Therefore a concept that employs both antiinflammatory drugs and vasculature formation may be necessary to mitigate biocompatibility issues [10,20]. However, for a biosensing system loaded with a therapeutic agent, introducing different types of drug molecules may interfere with the sensing mechanism and cause a loss in system functionality and, in the literature, such systems are not widely explored.

9.2.2 Sensitivity Sensitivity, or limit of detection, refers to the minimum amount of analyte that can be detected by a biosensor. It is also the ability to respond accurately and in a measurable fashion to the changes in analyte concentration in vivo [22]. Sensitivity is an important criterion for biosensing
drug delivery systems. For a successful in vivo application, the biosensor should be able to detect even trace amounts of analyte concentrations as low as nanogram per milliliter [23]. The release of the relevant amount of therapeutics can only be accomplished if the biosensor’s sensitivity is high enough and within the detection limits. Most biosensors exploit enzyme-based redox reactions and the increase of sensitivity in such systems can be achieved by improving electron transfer between the target redox molecule and the biosensor surface [22]. A substance that augments the rate of electron transfer between the sensor and target redox species can be integrated to a biosensing system to increase sensitivity. These materials function by opening new electron conduction pathways or electron relays [22]. Different types of materials have been studied as electron mediators, such as metal nanoparticles [24], nanofibers [25], conducting polymers [26], and carbon nanotubes [27]. An advanced biosensor sensitivity can also be obtained by

altering the surface properties of the biosensor. For example, platinization of the biosensor surface presents nanoscale roughness. These nanoscale features enhance the surface area and, thereby, sensitivity, since the catalytic activity of biosensor is intensified [28]. In enzyme-based biosensors, analyte react with enzymes to generate a detectable signal. Therefore high sensitivity is associated with high amounts of analyte consumption. This phenomenon may cause depletion of the analyte around the biosensor which is a serious concern if the analyte concentration is important for metabolic activities, that is, if high metabolic activity takes place within the vicinity of the biosensor [29]. Surface poisoning is also a serious problem associated with electrochemical-based biosensors. Products of enzyme reactions may adhere to the biosensor surface and eventually cause loss of sensitivity. To prevent this, electrochemical cleaning is performed, which is achieved by applying high potentials for the desorption of products [30]. However, this process not only weakens biosensor selectivity, but also may harm therapeutic molecules embedded in the biosensor. Biosensor sensitivity and performance are easily altered in physiological conditions. Sensitivity loss may cause inefficient dosing of therapeutics, which eventually lowers the treatment efficacy. Overall, biosensing devices that work through electron transfer mechanisms such as enzyme-based or electrochemicalbased systems are still at the laboratory stage and are yet to be commercialized for in vivo healthcare use [22].

9.2.3 Selectivity Selectivity is the ability of the biosensor to detect a particular analyte among the pool of many metabolites, including contaminants and/or other biomolecules. This property

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9.2 ADAPTATION OF BIOSENSING
DRUG DELIVERY SYSTEMS TO IN VIVO APPLICATIONS

occurs when molecular recognition elements are introduced to the biosensor [8]. Enzyme molecules are commonly preferred for the design of selective biosensors. For instance, oxidases are usually used in selective biosensor designs since their cosubstrate, oxygen, is abundant in physiological environments. Even though the reactions that occur on enzymebased biosensors are selective, the presence of undesirable molecules have a negative impact on the electron conduction pathway and hence on selectivity [22]. One method to enhance selectivity is to integrate membranes that are selective to specific analyte in the body. For example, selectively permeable membranes such as polymer coatings around the enzyme layer of the biosensor can separate molecules based on their size or charge, or both [22]. Electrostatic repulsion systems are used for a charge-based selection. Some electrostatic repulsive materials include chitosan [31], nafion [31], cellulose [32], and 4vinylpyridine-styrene copolymer [33]. These polymeric membranes can be deposited onto the biosensor surface via layer-by-layer technique with electrostatic interactions. In membrane selection, other than analyte interferences, the interaction between therapeutic in biosensor and selective membrane is an essential criterion for biosensor design. Additionally, these membranes are prone to degradation when they are implanted in vivo [22]. Size-exclusion membranes can also be used as selectivity agents. The most important design parameter in these membranes is the uniform membrane pore size. Electropolymerized membranes that consist of poly(phenylenediamine) [34], poly(aminophenol), and polyphenol [35] had a great potential inducing selectivity to glucose biosensors. However, membrane-based selectivity strategies hamper sensitivity [30]. In addition to selective membranes lowering the biosensor detection potential [22], coating the biosensor with an interference preoxidizing layer [36] using mediators such as metal

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complexes [37], organic dyes [38], and conducting organic salts [39] are powerful means for enhanced selectivity. Selectivity and sensitivity are perhaps the most important criteria for a biosensing system to be fully functional. However, the two concepts are interrelated and the strategies to improve sensitivity usually affects selectivity adversely [30]. It should be noted that this adverse interaction influences the drug-release mechanism and must be extensively studied before in vivo adaptation. Research on biosensors mostly focus on sensitivity and selectivity of the sensor, while the stability issue remains unaddressed. For instance, stability is an important factor for enzyme-immobilized biosensors. Long-term stability of enzymes ensures durability for the biosensing unit. Enzyme immobilization to biosensor surface plays a key role in stability. Enzymes can be linked to the biosensor via adsorption, covalent bonding, intermolecular cross-linking, or entrapment in the gel matrix [40]. For a more durable biosensor, cross-linking or entrapment is preferred, since enzyme degradation and leakage can occur in the adsorption approach. Covalent bonding is a better option to eliminate leakage, while intermolecular crosslinking can result in a loss in sensor activity [41]. Since the cross-linking reaction is fast and has unknown mechanisms, it is usually hard to obtain reproducible biosensors. Reproducibility and repeatability should also be considered for the design of biosensing
drug delivery systems. Reproducibility can be determined by testing the sensitivity of each manufactured biosensor and comparing the outputs after analyte is introduced to the system. For example, graphiteTeflon-LOX-POD biosensors were fabricated by Serra et al. and the researchers reported sensorto-sensor reproducibility by assessing the longterm stability and sensitivity [42]. In addition to the most important benchmarks (such as biocompatibility, sensitivity, and selectivity) we should consider many

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other aspects when designing a biosensor with a drug delivery mechanism. Stability of the sensor [43], miniaturization [30], response time of the sensor to the signals [22], and spatial and temporal resolutions [8] are other parameters that affect biosensor function.

9.3 MATERIALS AND DEVICES USED AS BIOSENSING
DRUG DELIVERY SYSTEMS Biosensing
drug delivery applications utilize responsive hydrogels, biomedical microelectromechanical systems (bioMEMS), microdevices, and electrochemical sensors. An ideal biosensor combined with a drug delivery system should be biocompatible, resistant to biofouling, minimally invasive, cost effective, scalable, sensitive, and should respond quickly and operate for long periods. Even though there is no commercial drug delivery combined with a biosensing system, the number of studies on this topic is growing. In this section, we discuss responsive hydrogels, bioMEMS, microdevices, and electrochemical sensors and each system’s pros and cons together with the barriers to commercialization of these systems in detail.

9.3.1 Responsive Hydrogels Hydrogels are three-dimensional (3D) polymeric materials that can absorb water [44
46]. They can mimic natural tissue due to their high-water content, porosity, and flexibility [47]. They have various real-life applications such as contact lenses, wound dressings, and hygiene products [47]. Hydrogels are more advantageous when they are responsive to changes in the environment. They can be responsive to the different environmental stimuli including pH [48
50], temperature [51], light [52], and chemical ligands such as enzymes [53,54]. Particularly, pH-responsive

hydrogels are promising for the treatment of diabetes. These hydrogels use glucose oxidase enzymes which consume glucose and result in a local pH change to trigger insulin release [55,56]. Besides pH stimulus, advances in biomaterial science led to the development of hydrogel systems that can interact with specific molecules and respond accordingly. These hydrogels are also called “intelligent” hydrogels which can sense specific analyte in the body. They sense specific types of molecules such as binding proteins and the interaction between these molecules and polymer network results in the release of therapeutics [56
58]. These “intelligent” hydrogels are widely studied for various applications, including combined biosensing and drug delivery applications. In an ideal design, a hydrogel senses a change of a molecule in the body. Afterwards, a therapeutic is released as a result of a conformational change such as swelling or deswelling in the polymer network (Fig. 9.2) [57]. However, the main problem of responsive hydrogels is delayed response time. Depending on the thickness of the hydrogel, diffusion time of molecules can take several hours [59]. This problem can be solved by fabricating very thin hydrogels. However, thinner hydrogels generally do not meet the mechanical strength requirements, which limit their in vivo applications [56]. Nevertheless, their tissue-like properties, easy fabrication methods, biocompatibility and biodegradability make stimuli-responsive hydrogels promising candidates for biosensing and drug delivery tools. Once the issues related to slow diffusion of therapeutics upon stimulation is solved, these gels can be used for real-life applications.

9.3.2 Biomedical or Biological Microelectromechanical Systems Microelectromechanical systems (MEMS) are devices with electrical and mechanical

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FIGURE 9.2 Representative therapeutic release mechanism of responsive hydrogels depending on conformational change.

components. They are used to sense different variables such as pressure, vibration, strain, sound, flow, and angle. They have wide application areas, such as accelerometers, microphones, and gyroscopes [60]. MEMS designed for biomedical applications are named biomedical or biological microelectromechanical systems (bioMEMS). BioMEMS technology gained tremendous attention in biomedical engineering field mainly for biomolecular analyses and sensing [61]. They convert physical, chemical, or biological signals into electrical signals, which are easier to monitor [62]. BioMEMS offer several advantages in biosensing applications such as lower response time, being highly scalable, reproducible production, and high sensitivity. They can be combined with other MEMS such as drug delivery systems [63]. In an ideal bioMEMS, physical, chemical, or biological signals are converted into electrical signals which trigger the drug release. BioMEMS is implanted into the human body and the drug is released according to sensor

feedback [64]. Thus BioMEMS have a great potential for sensing and controlling release of therapeutics with high precision. Although they offer very attractive properties, there are still number of challenges for bioMEMS applications in vivo. When bioMEMS are implanted into the body, macrophages are accumulated at the implantation site upon protein adsorption on the device and is followed by inflammation. Next, macrophages lead to the formation of a collagen network on the device which causes poor transport of molecules to the device [63]. Moreover, a wrong signal could trigger more or less drug release, which might have fatal consequences. Thus bioMEMS must operate for a long time without any malfunction [63]. After finding a solution to the problems associated with in vivo implantation, bioMEMS will become ideal tools for biosensing
drug delivery applications to treat certain diseases since they have advantageous properties, such as short response time, high scalability, and high sensitivity.

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9.3.3 Microdevices Microdevices are similar to bioMEMS, as they sense different types of signals. However, they do not have electrical and mechanical components [65]. Generally, they have a reservoir filled with the therapeutic agent. The reservoir opens upon stimulation with a biological, chemical, or physical cue. Fig. 9.3 demonstrates the representative working mechanism of microdevices. The drug-release mechanism can be diverse for microdevices and there are no boundaries for a microdevice design, where some systems can be adapted to microdevices with an intelligent design. However, they carry similar concerns as those with bioMEMS concerning host response after implantation. Even though there are limited studies about biosensing
drug delivery applications with microdevices, real-life applications of microdevices could be realized with novel designs. Examples of microdevices are discussed in Section 9.4.

9.3.4 Electrochemical Biosensors Electrochemical biosensors have electrodes which translate the chemical signal into an

electrical signal [66]. Electrochemical sensors are able to detect various biomolecules in the human body such as glucose, cholesterol, uric acid, lactate, DNA, hemoglobin, blood ketones, and others [67,68]. Thus they have great potential to treat diseases related to imbalances of biomolecules. Mostly, they are widely used for biosensing applications, and the number of studies on biosensing
drug delivery applications is limited. Enzyme- and protein-based electrochemical biosensors are widely studied for detecting specific analyte [69,70]. In these biosensors, enzymes or proteins are immobilized on the transducer and specific analyte is measured with the help of measurable and electroactive by-products [69]. Enzyme- or protein-based electrochemical biosensors that have drugrelease capability can be useful for the treatment of various diseases. For example, xanthine oxidase enzyme catalyzes the production of hypoxanthine and xanthine and overproduction of these products cause renal failure [71,72]. Xanthine oxidase is inhibited by allopurinol; a biosensor mechanism that detects overproduction of hypoxanthine and xanthine, and releases allopurinol as a response could be a great example for an electrochemical biosensor for drug delivery.

FIGURE 9.3 Representative working mechanism of a microdevice.

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9.4 CASE STUDIES

As a working principle, an electrochemical biosensor should sense and convert the signal effectively and the same signal must trigger the therapeutic release. However, an additional therapeutic reservoir in the biosensor may make the system complicated. In these systems, an electrical signal should open or close the drug reservoir. Additional mechanisms between the electrical signal and reservoir increases the complexity and challenges the maintenance of the system. Still, since electrochemical-based sensors are widely used and established, combining the existing ones with drug delivery systems is promising for the treatment of various diseases.

9.4 CASE STUDIES In recent years, biosensors have been used as noninvasive devices for the diagnosis of diseases; however, medical intervention remains an unresolved issue [73]. Continuous monitoring and medical intervention are needed for the efficient treatment of chronic diseases. In this chapter we discussed selected in vivo biosensing applications for the detection and treatment of diseases such as diabetes and their present state-of-the-art principles. Diabetes mellitus is a major chronic disease that has serious effects on patients’ quality of life. To date, periodic insulin injection is the most commonly used approach to control the blood glucose levels of diabetic patients. However, it is difficult to maintain blood glucose levels within the normal range with this method. Thus self-regulated delivery systems are required for more efficient treatments [74]. In these recently developing delivery techniques, the main challenge is to reduce the lag time of insulin delivery between interstitial and intravascular areas and to provide improved glucose-lowering action [16]. Glucose-responsive hydrogel-based insulin delivery systems need the synergy of a

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responsive insulin release element and a realtime glucose-sensing mechanism. Glucose oxidase enzymes have been commonly preferred as a glucose-sensing unit. This enzyme converts glucose into gluconic acid that produces an acidic stimulus for a pH-responsive hydrogel. As a result, a shrinking or swelling response of hydrogels releases the loaded insulin [65]. For example, Traitel et al. reported a glucose-responsive hydrogel system based on poly(2-hydroxyethyl methacrylate-co-N, N-dimethylaminoethyl methacrylate) (HEMAco-DMAEMA), with entrapped glucose oxidase, catalase, and insulin. For in vivo experiments, hydrogels were implanted into the peritoneum of rats and the results showed that this platform was effective in reducing blood glucose levels for 6 days, entrapped insulin retained its activity, and no tissue encapsulation was observed around matrices [74]. In another study, a tubular microdevice comprised of an albumin-based bioinorganic membrane with pH-responsive hydrogel nanoparticles was tested on diabetic rats. Similarly, this microdevice provided the conversion of environmental glucose level changes to a pH stimulus. The volume of pH-responsive hydrogel nanoparticles and, thus the permeability of the membrane was regulated, and insulin was released from the reservoir. This system is called a “microdevice” since the responsive membrane is combined with a microfabricated polydimethylsiloxane (PDMS) structure and does not require tethering wires or tubes (see Fig. 9.4). The microdevice demonstrated glucose-responsive insulin release over multiple cycles at clinically relevant glucose concentrations during in vitro testing. Hyperglycemia was reversed in diabetic rats for 1 week in vivo. Moreover, stability of insulin retrieved from an ex vivo microdevice at day 15 of implantation suggested that the insulin function was retained [65]. Matsumoto et al. suggested a synthetic material
based approach as an alternative

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FIGURE 9.4 (a) Schematic of the PDMS grid
gel biosensor with bioinorganic membrane (with inset for (c)). (b) Picture of completed PDMS grid
bioinorganic gel membrane biosensor. (c) Schematic diagram presenting responsive insulin release in a glucose-rich environment to generate nanopores. Source: Reproduced with permission from Chu et al., In vitro and in vivo testing of glucose-responsive insulin-delivery microdevices in diabetic rats, Copyright 2018, Lab on a Chip.

strategy to protein-based insulin delivery systems. The glucose-responsive system was made of boronic acid containing a polymer confined within a single catheter. Subcutaneous implantation of the device to healthy and diabetic mice controlled glucose metabolism by sensing interstitial glucose level fluctuations under both insulin-deficient and insulin-resistant conditions for 3 weeks [16]. Lee et al. developed another synthetic material
based glucose sensor. They built an electrochemical biosensor consisting of graphenedoped with gold, which has an interface for the successful transfer of an electrical signal. Incorporation of gold enhanced the electrochemical activity of the patch compared to that obtained with plain graphene. These wearable patches were suitable for sweat-based diabetes

monitoring and response therapy. The stretchable patch is composed of a heater, temperature, moisture, glucose and pH sensors, and polymeric microneedles. They showed that the patch can deliver Metformin transcutaneously via thermal activation and decrease blood glucose levels in diabetic mice (Fig. 9.5) [75]. BioMEMS technology is adapted to miniaturize glucose biosensors and to provide improved response time in vivo. Recently, Senseonicsr developed a fluorescent-labeled glucose sensor which was implanted into the arms of 24 human patients and continuously monitored for 90 days [63]. In this system, a fluorescent tag was copolymerized with poly (2-hydroxyethyl methacrylate) (pHEMA) which is covered with a semipermeable membrane. Glucose interacts with the fluorescent tag to

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FIGURE 9.5 Image of the wearable diabetes monitoring and therapy system. The heater combined with the microneedles, which is adhered on the abdomen skin of the diabetic mouse (top-left), optical image of mouse skin stained with trypan blue to observe the microsized holes resulted from the penetration of the microneedles (top-right). Optical and infrared camera images of the patch with the thermal actuation (bottom-left). Blood glucose concentrations of mice for the treated group (drug and patch) and control groups. Source: Reproduced with permission from Lee et al., A graphene-based electrochemical device with thermoresponsive microneedles for diabetes monitoring and therapy, Nat. Nanotechnol, Copyright 2018.

create a signal by diffusing through the membrane. The results of the clinical trials were promising when compared with blood glucose levels obtained intravenously [76].

9.5 CONCLUSION In this chapter, we introduced a novel class of biosensors, namely biosensing
drug delivery systems. We provided an overview on the requirements of biosensors for in vivo applications, and the current status of materials and devices used as biosensing
drug delivery systems. Finally, we reviewed in vivo research findings of biosensing
drug delivery systems. A complex structure that senses biological

cues in vivo and delivers its therapeutic cargo has a tremendous amount of specifications. A quintessential biosensing mechanism that releases the drug in response to analyte measurements in vivo requires all these qualities. Translation from benchtop to healthcare necessitates rigorous in vitro and in vivo tests. With the discovery of novel analyte
sensor interaction species, miniaturization of biosensors and better communication between researchers and life scientists, the future of biosensors with drug-release capability is bright. Nevertheless, development of a technology that considers all analytical requirements for in vivo use and can adapt biosensing drug delivery systems for in vivo applications still remains a big challenge.

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956.

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C H A P T E R

10 Nanobodies and Their In Vivo Applications Prashant Singh1,*, Fanding Gao2 and Andrea Bernat1 1

Department of Nutrition, Food and Exercise Sciences, Florida State University, Tallahassee, FL, United States, 2Food Science Program, Division of Food Systems and Bioengineering, University of Missouri, Columbia, MO, United States

10.1 INTRODUCTION Mammalian antibodies are mostly composed of two identical heavy chains and two identical light chains. Both types of chains contribute to the antigen-binding site, which is usually flat or concave. However in 1993, Hamers-Casterman et al. reported immunoglobulin class G (IgG) antibodies from the Camelidae family consisting of heavy-chainonly antibodies (HCAbs) [1]. In addition to camels, HCAbs have also been found in llamas and sharks. These antibodies are completely devoid of light chains and consist of three globular domains. HCAbs are comprised of two constant domains (CH2 and CH3, which are homologous to the Fc domain of classical antibodies), namely, an antigen-binding variable heavy chain (VHH) domain and a hinge region [2]. The antigen-binding VHH region of the HCAbs can be cloned and expressed in a bacterial expression system (Fig. 10.1). These

expression products have been found to be strictly monomeric and single-domain binding entities [3]. Due to the small size of the VHH (diameter B2.5 nm and height B4 nm) and its shape, like a rugby ball, the company Ablynx (a company focused on developing therapeutic application using camelid HCAbs) renamed VHH as nanobodies [2]. The organization of the VHH region of HCAbs is very similar to the variable heavy (VH) domain of the conventional antibody. However, the VHH region of the HCAbs consists of important notable differences that explain their antigen-binding capacity and their properties [4]. The VHH is composed of four conserved sequence stretches called the framework regions (Fig. 10.1). These four framework regions surround the three hypervariable regions, which are also known as complimentary determining regions (CDR). These three hypervariable regions are located in loops where they participate in antigen

*Corresponding author.

Advanced Biosensors for Health Care Applications DOI: https://doi.org/10.1016/B978-0-12-815743-5.00010-X

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FR1

CDR1

FR2

VHH or Nanobody

CDR2

FR3

CDR3

FR3

(A)

F c (D) Heavy-chain IgG

Nanobody

(B)

(C)

FIGURE 10.1 (A) Structural organization of heavy-chain-only antibody with representation of framework region (FR) and complementarity determining regions (CDR) (B) The camelid heavy-chain antibodies completely lack light chains (constant light and variable light chains) and the first heavy-chain constant domain (CH1). (C) The antigen-binding domain of heavy-chain antibodies is formed in the heavy-chain variable domains (VHH) which are also known as nanobodies. The VHH sequence can be cloned and expressed in bacteria, yeast, and plant expression systems. (D) Nanobodies can be either monomeric or they can be joined to form dimeric or polymeric structures. Adapted from P. Vanlandschoot, C. Stortelers, E. Beirnaert, L.I. Iban˜ez, B. Schepens, E. Depla, et al., Nanobodies: new ammunition to battle viruses. Antiviral Res. 92 (2011) 389
407.

recognition. The CDR1 and CDR3 amino acid sequence of the VHH are much longer and form fingerlike extensions which enable them to extend into cavities and bind to hard-toreach antigenic sites (e.g., active sites of enzymes) [5]. The CRD1 and CRD3 also contain cysteine, an amino acid that participates in disulfide bond formation providing higher stability [4]. Additionally, the framework-2 region of the VHH contains more hydrophilic amino acids, which help to increase the solubility of HCAb [6]. Moreover, these longer CDRs of VHH (CDR1 and CDR3) allow formation of longer loops and result in increasing the size of the VHH paratope (antigen-binding site) [7], allowing the VHH to leave a small footprint of its paratope on its antigen, which is only half the size when compared to the classical antibody format (VH-VL).

10.2 HEAVY-CHAIN-ONLY ANTIBODIES GENERATION Antigen-specific nanobodies can be generated by subcutaneous immunization of camels or llamas with a target antigen (e.g., purified protein, bacterial cells). Subcutaneous immunization of camels or llamas is repeated several times in the presence of an adjuvant. The lymphocytes from either a lymph node biopsy or from peripheral blood samples are collected after 6
7 weeks of immunization, and peripheral blood mononuclear cells (PBMC) are then collected. The PBMC are lysed and used for isolating mRNA. Reverse transcriptasepolymerase chain reaction (RT-PCR) is used to convert isolated mRNA into cDNA. The obtained cDNA is, therefore, used to amplify the entire VHH region using conserved PCR primers of VHH. The amplified VHH gene

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10.4 NANOBODY CHARACTERISTICS MAKING THEM SUITABLE FOR THERAPEUTIC APPLICATION

Once the titer level rises peripheral blood mono nuclear cells(PBMCs) are collected.

After immunization a period blood sample is collected and tested. for antisera titer. Immunization of llama with target antigen. Immunization is performed several times (4-6 rounds) in the presence of an adjuvant. Multiple rounds of affinity panning is performed and positive clones are sequenced. These clones can be expressed in different expression systems (bacteria, yeast, and plant) to obtain nanobodies.

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Total RNA is isolated

mRNA is reverse transcribed to produce cDNA

Using cDNA as a template and conserved VHH primers all VHH regions are amplified. Thus obtained VHH sequence are cloned in a phage display vector creating a VHH library.

Using cDNA as a template, conserved VHH primers are used to amplify all VHH regions.

Using cDNA as a template, conserved VHH primers are used to amplify all VHH regions.

FIGURE 10.2 A schematic diagram of nanobody generation.

products are cloned using a phagemid vector to obtain a VHH library. M13-based phage display vectors are commonly used for cloning the amplified products. The libraries express the cloned gene product on their surface and are screened through multiple rounds of affinity panning to identify antigen-specific VHH. Positive clones are identified using enzyme-linked immunosorbent assay and are subsequently sequenced. Finally, positive clones are expressed and purified to generate nanobodies [2] (Fig. 10.2).

10.3 DRAWBACKS OF MONOCLONAL ANTIBODIES Monoclonal antibodies (mAbs) are extensively used in the areas of diagnostics and healthcare. Despite their extensive use, mAbs present several disadvantages, namely, they are very expensive and difficult to produce. Additionally, their large molecular size

(150 kDa) limits their tissue and tumor penetration, thus limiting their biodistribution and efficacy. Moreover, they can also induce immune reactions, which further limits their long-term application. Lastly, mAbs have a half-life of several days, which leads to high background signals when used for molecular imaging.

10.4 NANOBODY CHARACTERISTICS MAKING THEM SUITABLE FOR THERAPEUTIC APPLICATION Nanobodies have clear advantages over mAbs, which make them more suitable for in vivo or therapeutic applications. Compared to mAbs, nanobodies are very small, which enable them to target hard-to-reach epitopes on the antigens [8]. Additionally, nanobodies have been shown to have higher stability and solubility [7,9,10]. Furthermore, nanobodies

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can be easily cloned in bacterial, yeast, and other expression systems, enabling their mass production at comparatively lower cost. Additionally, they have very low immunogenic potential elicited by the Fc domain. In a phage I clinical trial conducted by Ablynx (Gent, Belgium), the application of a nanobody-based therapeutic showed very low immunogenicity [11]. All these advantages allow for the development of nanobody-based therapeutics that can be marketed in a readyto-use format. Antigen-specific nanobodies are selected using the phage display method, which, in principle, enables selection of nanobody for any protein of interest. As nanobodies cannot freely cross the cell membrane, they can be used to target extracellular antigens. Proteins that are solely overexpressed in disease cells are suitable targets for developing nanobodybased therapeutics. Nanobodies can be used to target transmembrane proteins, cancer cells, cell surface receptors, and tumor vasculatures.

10.5 NANOBODIES BINDING TO TRANSMEMBRANE PROTEINS Transmembrane proteins that are overexpressed in cancerous cells can be effectively targeted using specific nanobodies. Nanobodies were initially developed to target extracellular proteins, which are either unique expression proteins or are overexpressed in cancerous cells, for example, human epidermal growth factor receptor 2 (HER2). The HER2 protein is overexpressed in 15%
30% of breast cancer and 10%
30% of gastric cancer cases [12]. Therefore antagonistic nanobodies targeting specific cancer proteins can be used for controlling tumor-cell growth, tumor proliferation, and can also induce their apoptosis by blocking signal transduction molecules produced by cancerous cells. Epidermal growth factor receptor (HER1)-specific nanobodies and biparatopic

nanobodies were developed through the phage display selection method to select antagonistic anti-EGFR nanobodies [13,14]. Application of antagonistic anti-EGFR nanobodies resulted in effective inhibition of EGFR signaling and tumor growth [13,14]. The VEGFR2 receptor has an important role in angiogenesis and embryogenesis. This particular receptor is overexpressed in lung and colon cancers [15]. Nanobodies against VEGFR2 were selected through the phage display selection method and in vitro tests were conducted using VEGFR2-specific nanobodies which have shown inhibition of capillary tube formation [16]. Similarly, the c-MET receptor is involved in motility, morphogenesis, and regulation of cell proliferation [17]. The c-MET receptor is activated by hepatocyte growth factor (HGF) binding and has been implicated in several types of cancers such as colon, breast, ovarian, and hematological [18]. Nanobodies specific to c-MET receptors and nanobodies that compete with the HGF for binding on the c-MET receptor were identified through selection on immobilized c-MET receptors [19]. In this study, a drug delivery system consisting of anti-c-MET nanobodies was cross-linked with albumin nanoparticles. This anti-c-MET nanobody complex was specifically taken up by the c-MET expressing cells and transported to lysosomes for degradation. Treatment of tumor cells expressing c-MET with the anti-C-MET nanobody complex resulted in downregulation of the total c-MET proteins [19]. Chemokines and chemokine receptors also play an important role in several cancers. CXCR7 belongs to the membrane-bound G protein
coupled receptor superfamily that binds to chemokine CXCL11 and CXCL12, regulating breast and lung cancer malignancies [20]. Overexpression of CXCR7 promotes the growth of tumors on breast and lung cancer cells. In a recent study, the specific nanobody against CXCR7 was identified [21]. The study demonstrated the inhibition of tumor growth using this nanobody

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10.6 APPLICATION OF NANOBODIES IN MEDICAL IMAGING

specific for CXCR7. The application of the selected nanobody also reduced secretion of angiogenic chemokine CXCL1 from cancer cells [21].

10.6 APPLICATION OF NANOBODIES IN MEDICAL IMAGING Early detection and diagnoses of diseases are the most critical steps for treatment. In the diseased state, some antigens are expressed at an elevated level in blood samples, which are easy to detect. However, some antigens are produced on internal organs and are hard to reach. Radiolabeled mAbs have been used for localization of antigens, but the use of these radiolabeled mAbs has major drawbacks. mAbs have a slow diffusion rate and long half-life time, therefore, it takes 2
4 days to reach an optimal contrast level, thus necessitating the use of a radioisotope with a long halflife period. Additionally, due to the large size of a conventional antibody, a typical antibody is characterized by low-tissue penetration, while their clearance from the body is also slow. Therefore the use of a radioisotope with a long half-life period and slow clearance from the body results in patient exposure to radiation for longer periods, which may cause serious side effects [22]. Nanobodies are good candidates for imaging and localization of disease-related antigens. Due to their small-molecular size they can be used to target antigens located in hardto-reach areas and access cryptic antigens. Additionally, due to the small size of the nanobody-imaging dye complex, they can quickly reach their target site (within a few hours) and facilitate rapid detection of the targeted antigen. The small size of nanobodies also facilitates speedy clearance of unbound nanobodies. Further, the high affinity and specificity of nanobodies toward their target make them ideal candidates for molecular

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imaging and facilitate effective localization of targeted antigens [23]. At present, there are several imaging techniques which have been incorporated into daily clinical use, for example: (1) nuclear imaging entailing single photon emission computed tomography, and positron emission tomography (PET) using radionuclides such as 99mTc, 89Zr, 68Ga; and (2) optical imaging via the near-infrared (NIR) fluorescence imaging [24,25]. However, there are only a few reports on radiolabelednanobodies for imaging applications [26,27]. Short-lived PET-radioisotopes fluorine-18 (18F, t1/2: 110 minutes) or gallium-68 (68Ga, t1/2: 68 minutes) coupled with nanobodies are considered to be more suitable for imaging applications as they have significantly lower radiation burdens on patients [22]. As noted in Section 10.5, specific nanobodies have been selected for the detection of various extracellular cancer biomarkers. In one study, an anti-EGFR nanobody [28] and an anti-HER2 nanobody [29] radiolabeled with 68Ga was evaluated for localizing target proteins. The 68 Ga-labeled nanobody 7D12 specific to a A431 tumor showed high uptake by the tumor cells (tumor-to-blood ratio), which further increased with time (1
3 hours). The targeted tumors were clearly visualized using PET imaging [28]. In another study, anti-HER2 nanobody (2Rs15dHis6) labeled with 68Ga was evaluated using HER2-positive and HER2-negative tumor xenografts. Biodistribution studies performed with 2Rs15dHis6 nanobodies conjugated with 68 Ga showed specific and fast uptake by the tumor cells (HER2) within 1 hours of receiving the injection. This high tumor-to-blood ratio of a nanobody conjugated with 68Ga obtained after 1 hours, showed a high contrast in PET images and facilitated rapid detection and localization of the tumor [29]. In another in vivo study [30], the efficacy of nanobodies and mAbs were compared. An ART2-specific monovalent nanobody (s 1 16a), a conventional antibody (Nika102), and a bivalent

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Fc-fusion protein of s 1 16a (s 1 16-mFc) were labeled with AlexaFluor680. In vivo experiments were conducted using two mouse models (ART2 gene–knock out and ART2 overexpressing) and AlexaFluor680 conjugates were intravenously injected in mouse models. NIR fluorescence imaging reveled specific labeling of target lymph nodes. AlexaFluor680labeled nanobody s 1 16a and s 1 16mFc inhibited ATR2 expressing T cell lymph nodes within 10 minutes, whereas the Nika102 (conventional antibody) labeled with AlexaFluor680 required 2 hours for labeling the target protein. The nanobodies labeled with short-life dye or radionuclides are best suited for noninvasive, short-term use. However, rapid clearance of unbound nanobodies are retained in the kidneys, which makes targeting any antigen close to the kidneys difficult.

10.7 INFLAMMATORY DISEASES Immune-mediated inflammatory diseases (IMID) are a group of conditions characterized by inflammation caused by inflammatory pathways, which are activated by abnormal immune responses [31]. Two examples of IMID are rheumatoid arthritis and Crohn’s disease, which are seemingly unrelated to the naked eye but share some common inflammatory pathways [31]. Ion channels are proteins that regulate ion transfer across biological membranes [32]. Ion channels have been linked to many diseases, including IMID [32]. mAbs have been explored for treating IMID. Even though antibodies are highly selective toward their target, they are not able to halt ion-channel functions [33]. Nanobody technology has been explored for developing therapeutic applications for treating ion-channel
related IMID. Injured cells release a large amount of ATP. Ion channel P2X7, upon sensing ATP molecules, initiate a

proinflammatory signaling cascade resulting in the release of cytokines (e.g., interleukin-1β), causing inflammation. Therefore blocking this interaction is a good candidate for treating numerous diseases [33]. Danquah et al. [33] reported the development of a nanobody 13A7 that effectively blocked mouse P2X7 ion channel and another nanobody 14D5 which potentiated gating of the channel. Using 1-fluoro-2,4-dinitrobenze (DNFB) as the contact allergen, they found that the animals treated with nanobody 12A7-HLE did not gain as much ear weight as the control animals, and displayed lower levels of IL-1β in the ear tissues [33]. Another nanobody, Dano1, characterized in the same study was specifically designed to inhibit human P2X7. In endotoxintreated human blood, application of Dano1 was 1000-times more effective in preventing the release of IL-1β. The study showed application of nanobody technology for improving inflammatory disorders. Acquired thrombotic thrombocytopenic purpura (TTP) is a life-threatening, autoimmune disorder. Acquired TTP is caused by severe deficiency of ADAMTS13 due to the presence of inhibitory autoantibodies [34]. A decrease in ADAMTS13 activity results in the formation and accumulation of large multimers of von Willebrand factor, which, in turn, bind to the platelets and induce aggregation and formation of microthrombi [35]. These microthrombi prevent blood from flowing to the organs and are responsible for causing tissue ischemia and organ failure (e.g., brain, heart, and kidney). Thus the mortality rate is very high [36]. Treatment regimens used for treating acquired TTP include frequent plasma exchange for removing autoantibodies and ultralarge von Willebrand factor multimers, replenish ADAMTS13, and immunosuppressive therapy. Caplacizumab is the first nanobody that is expected to be licensed for clinical applications in near future [37]. Caplacizumab is an

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10.9 APPLICATION OF NANOBODIES AGAINST VIRUSES

anti-von Willebrand factor humanized nanobody that targets the A1 domain of von Willebrand factor and blocks interaction between large multimers of von Willebrand factor and glycoprotein Ib-IX-V receptor of platelets. Caplacizumab is currently in clinical trial for treating acquired TTP. In a phase 2, controlled randomized trial, Caplacizumab (10 mg/daily) was subcutaneously administered to patients. Compared to placebo, Caplacizumab induced faster resolution of symptoms during the treatment period [36].

10.8 CHRONIC RESPIRATORY DISEASES Chronic respiratory diseases affect airways and other lung structures [38]. Each year, more than three million people die from chronic obstructive pulmonary diseases (e.g., asthma, occupational lung diseases, and pulmonary hypertension) [39]. Death caused by chronic obstructive pulmonary diseases is estimated to be responsible for 6% of all deaths worldwide [39]. During an allergic reaction, production of IgE antibodies are induced when exposed to an allergen. IgE interacts with FcεRI, a high affinity receptor present on mast cells and basophiles. This interaction is very strong and is essential for inducing allergenic responses. IgE neutralization is a modern way to treat allergic asthma. In 2001, researchers found an anti-IgE antibody, omalizumab, for treating allergies. Omalizumab is a monoclonal antibody that targets the high affinity receptor binding site of IgE and competitively inhibits levels of circulating IgE. Application of omalizumab can reduce exacerbation of symptoms and the steroid requirement in patients with allergic asthma [40]. This reduction was also observed in patients with moderate-to-severe allergic asthma [40]. After a 30-week doubleblind study, omalizumab improved asthma symptoms in patients. These patients also showed a reduction in inhaled corticosteroids

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dosage requirements [40]. Omalizumab is the only anti-IgE antibody that has been approved to date [40]. Using a similar strategy, Rinaldi et al. [41] reported development of ALX-0962, a bispecific anti-IgE nanobody, targeting IgE as well as human serum albumin [41]. In initial in vitro trials this nanobody showed dual functionality. First, it neutralized soluble IgE with a 5-times higher potency than omalizumab. Second, it resulted in binding and displacement of preformed IgE-FcεRI complexes on basophils. Initial work with ALX-0962 has also shown a reduction in allergen sensitivity [41]. Further research in this area can result in the development of a nanobody-based therapy with faster action to relieve allergic symptoms.

10.9 APPLICATION OF NANOBODIES AGAINST VIRUSES Viruses, due to their ability to rapidly mutate and transform into a new strain, can cause pandemic outbreaks. This ability to rapidly mutate helps the virus escape from the host’s immune system. These infections are often hard to treat. Several experimental nanobody-based therapeutics against viruses have been studied, for example, human immune deficiency virus-1 (HIV-1), hepatitis B/C virus, influenza virus, respiratory syncytial virus (RSV), rabies virus, foot-and-mouth disease virus (FMDV), poliovirus, rotavirus, and porcine endogenous retrovirus (PERVs) [42]. Viruses use specific epitopes on their capsid or envelope to attach to receptors on host cells. These epitopes are usually concealed and, therefore it is difficult to target these concealed epitopes using conventional antibodies; however, they can be accessed by nanobodies [5]. Thus many studies have explored the inhibitory effect of nanobodies to viruses. The influenza virus is a single-stranded RNA virus. The diversity among this group of viruses is based on their antigenic differences and are classified into four types: A, B, C,

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and D [43]. Influenza virus cases peak during the winter months. Highly pathogenic influenza A and B viruses can spread between animals and humans, especially H5 and H7 subtypes, and cause serious seasonal epidemics and pandemics with high morbidity and mortality [44]. Global influenza virus monitoring programs help to predict antigenic composition of influenza in the winter months and this information is used for formulating vaccines before the peak season. However, RNA viruses can mutate rapidly, resulting new strains that can cause unpredictable outbreaks (e.g., influenza A H1N1 virus, influenza A H5N1) which makes them a major public health concern. Oseltamivir is an antiviral medication and the drug-of-choice for treating influenza A and B infections. Iban˜ez et al. [45] assessed the in vivo protective potential of monovalent and bivalent nanobodies against H5N1 hemagglutinin (HA). In a mouse model, administration of a selected nanobody before H5N1 infection showed a protective effect and strongly reduced the viral replication in mice and significantly delayed time of death. The study further evaluated the efficacy of intranasal delivery of nanobodies to prevent or fight H5N1 infection. Intranasal delivery of nanobodies 24 hours before virus infection provided complete protection to the animal from death and morbidity. Additionally, the bivalent nanobody complex was found to be more effective than monovalent nanobody complex [45]. Respiratory syncytial virus (RSV) is a RNA virus that causes cold-like symptoms. Most people recover from RSV infection within a week or two; however, infants, senior citizens, and immunocompromised patients are highly susceptible to this virus. This virus is a major cause of acute lower respiratory tract disease in children, which results in inflammation of the small airways in the lung and eventually results in hospitalization [46]. Due to the lack of effective antiviral agents against RSV and

antiinflammatory therapies, treatment after hospitalization is mainly focused on rehydration and oxygen supply. Further, as RSV infections do not induce any long-term immunity against the virus, RSV infections occur throughout life, causing serious morbidity and even mortality among elderly individuals [47]. RSV possess two classes of transmembrane glycoproteins on the viral surface, namely, the fusion protein (F) and attachment protein (G). The G-protein of RSV mediates the receptor binding on the host cell, which is followed by F protein
mediated fusion of the virus with the host-cell membrane and viral entry [43,48]. Therefore F protein is considered to be essential for virus infection. Both proteins (F and G) contain epitopes for developing neutralizing antibodies. Detalle et al. [49] reported the development of a trimeric nanobody called ALX-0171, which binds with subnanomolar affinity to the antigenic site II epitope of RSV F-protein and prevents viral fusion with the host-cell membrane. The ALX-0171 nanobody showed superior RSV neutralization, which was much higher than palivizumab (a FDAapproved, humanized monoclonal antibody used prophylactically in high-risk infants). The study further reported complete RSV replication blockage in 87% of the samples tested. Moreover, the ALX-0171 nanobody was also found to be highly effectively when directly administered to the lungs of rats and reduced nasal and lung RSV titer levels. In 2014, a multicenter clinical trial was initiated to evaluate the safety, tolerability, and clinical activity of ALX-0171 nanobodies. This nanobody with extraordinarily high RSV-neutralizing activity holds great promise to treat RSV infections. The HIV pandemic started around 1981, which resulted in millions of deaths from acquired immunodeficiency syndrome (AIDS). The HIV virus replication can be effectively suppressed using antiviral drugs, but the virus cannot be completely eliminated from the infected person’s body. Further, applications

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of these drugs have considerable side effects, are costly, and can also result in the development of drug resistance. These factors are providing a push for the development of new methods for treating HIV. Gray et al. [50] reported a nanobody that has high sensitivity and specificity to global HIV-1. The nanobody developed in the study specifically captured ligands to detect p24, which is one of the earliest markers of HIV infection [50]. This nanobody was a 10th of the size of a conventional nanobody and showed higher sensitivity, broader specificity to global HIV-1 subtypes, and performed better than mAbs for the same target. The HIV enter the target cells using their trimeric envelope glycoprotein (gp), which consists of gp120 and gp41. The gp120 first binds to the CD4 of the target cells and results in a conformational change, which facilitates binding of gp120 to either CXCR4 or CCR5 (-transmembrane-spanning receptors). Following this interaction, gp41 induce a fusion of the target cell plasma membrane with the viral particles and eventually results in virus entry [51]. As these glycoproteins play important roles in virus entry, they are considered the main targets for developing neutralizing antibodies or nanobodies. In the past, multiple nanobodies each targeting an important component of the HIV protein have been developed: nanobodies A12, D7, and C8 targeting envelope protein gp120 [52]; nanobody Nb190 targeting Rev protein [53]; sdAb19 targeting viral Nef protein [54]; nanobodies 238D2 and 238D4 targeting the CXCR4 receptor [55]; nanobody JM4 exhibit potent HIV neutralizing activity by binding to CD4i epitopes [56]; J3, another potent HIV-1neutralizing nanobody that targets CD4 binding site and blocked the spread of HIV from primary macrophages to CD4 T cells [57]; and nanobody 2H10 which is specific to an membrane-proximal external region (MPER) [58]. All these nanobodies varied in their specificity and neutralized HIV with varying

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efficiency. Although, the nanobody-based approach has shown promising results in the laboratory, none have been approved for clinical use against viral infections. Rotavirus is one of the most frequent causes of child gastroenteritis. The rotavirus infection is more common in developing countries. In the United States, prior to introduction of vaccine all children below age 5 were infected by rotavirus. In the United States, prior to introduction of rotavirus vaccine, this pathogen was responsible for 400,000 doctor visits each year [39]. Rotavirus is a double-stranded RNA virus and possess an outer and inner capsid. The outer capsid is composed of VP7 and VP4 proteins, which are targeted for developing antibodies and nanobodies [42]. Van der Vaart et al. [59] developed a rotavirus-specific nanobody (2B10). In vivo trials conducted on mouse pups by daily administration either prevented occurrence of diarrhea or reduced the number of sick days. Gualtero et al. [60] developed a nanobody (3B2) targeting VP6 proteins of rotavirus. Oral administration of the 3B2 nanobody to mouse pups provided protection against bovine and murine rotavirus. Rabies, a RNA virus, is carried in the saliva of carrier or infected animals. Rabies viruses are transmitted through animal bites and the virus travels from the wound to the peripheral nerves and then to the central nervous system, causing brain infection. If postexposure prophylaxis is administered promptly, it provides a protective effect. In order to isolate the rabies virus-specific nanobody, Hultberg et al. [61] immunized llamas with a rabies vaccine meriux strain. In vitro tests conducted using five nanobodies (Rab-F8, Rab-E8, Rab-E6, Rab-H7, and Rab-C12) selected in the study showed neutralization potential against rabies prototype strains. Hepatitis C virus (HCV) is a common disease among intravenous drug users and causes serious liver infection, which often progresses to hepatocellular carcinoma or fatal cirrhosis.

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At present, there is no vaccine for Hepatitis C. The viral glycoprotein E2 is conserved among genotypes, thus making it a good candidate for developing a HCV-neutralizing antibody or nanobody. Tarr et al. [62] identified an E2specific nanobody (D03) that neutralizes HCV pseudotyped particles from six genotypes. Further, nanobody D03 efficiently inhibited cell-to-cell transmission of HCV.

10.10 APPLICATION OF NANOBODIES AGAINST BACTERIA Application of antibiotics is one of the most important medical interventions to treat bacterial infections. However, extensive use of antibiotics has dramatically increased antibiotic resistance among pathogenic bacterial strains and reduced the effectiveness of treatment, which has resulted in more severe infections and increased fatalities [63]. Emergence of antibiotic-resistant pathogenic bacterial strains, the high cost of alternative treatments, and the lack of new antibiotics against resistant strains has prompted the urgent need for development of new alternative therapeutics. Shiga toxin-producing Escherichia coli (STEC) consists of multiple E. coli serogroups, which possess ability to produce Shiga toxin. STEC are one of the most important foodborne pathogens and cause hemolytic uremic syndrome in humans. Infection caused by STEC are often hard to treat as administration of antibiotics can result in toxic shock syndrome in patients. Currently, no effective therapy or licensed vaccine is available for treatment of STEC infection. Mejı´as et al. [64] reported development of a family of Shiga toxin-producing gene Stx2-specific nanobodies, which neutralized Shiga toxin 2 at subnanomolar concentrations. One of the nanobodies (2vb27) showing superior results was further engineered into a trivalent nanobody complex (two copies of

anti-Stx2B VHH and one antiseroalbumin VHH). This engineered trivalent nanobody complex had extended life under in vivo conditions and also showed higher Stx2 neutralization capability in mouse models [64]. Enteric pathogens are responsible for a large number of infections and deaths in humans and domesticated animals. Oral passive immunization is a promising strategy to enhance the immunity of animals against infection. Enterotoxigenic E. coli are responsible for a large number of infections among postweaning piglets and cause severe diarrhea. Virdi et al. [65] designed an anti-ETEC nanobody by fusing variable domains of llama VHH to the Fc part of a porcine immunoglobulin (IgG or IgA) and this recombinant antibody was expressed in Arabidopsis thaliana seeds. These recombinant antibodies were produced at different levels (3% and 0.2% of seed weight) in A. thaliana seeds. In vitro analysis of A. thaliana seed extracts containing expressed antibodies showed inhibition of bacterial binding to the pig gut villous enterocytes. In an in vivo study (feed-challenge experiment) piglets receiving feed that expressed VHH-IgA at a dose of 20 mg/d/pig showed protection against infection. Piglets receiving this treatment showed a dramatic decrease in fecal pathogenic bacteria load and significantly higher weight gain [65]. Nanobodies can be used to target specific epitopes on bacterial surface proteins and prevent their binding with the host cells. Further, due to high specificity of nanobodies they can also be used for more effective delivery of prodrugs. Based on this principle, Moonens et al. [66] generated a nanobody that can specifically bind to the FedF lectin domain of F18 fimbrial adhesin of enterotoxigenic E. coli (ETEC) and STEC. In an in vitro assay four nanobodies (NbFedF6, NbFedF7, NbFedF9, and NbFedF12) generated in the study specifically inhibited attachment of E. coli cells with F18 fimbrial adhesin to piglet enterocytes. These selected nanobodies either competed for binding on the

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10.10 APPLICATION OF NANOBODIES AGAINST BACTERIA

blood-group antigen or they induced conformational change in FedF and prevented binding [66]. In future these nanobodies can be expressed in plants and consumption of these plants could provide protection against ETEC and STEC. Camelidae nanobodies have shown potential for development of oral immunotherapeutics. Positive VHH clones can be further stabilized using error-prone PCR methods and selection of resulting variants in a low-pH environment. Orally delivered nanobodies have to pass through the harsh gastric environment. Presence of proteolytic enzymes in the gastrointestinal tract can initiate degradation of orally administered nanobodies. Harmsen et al. [67] selected nanobodies that could withstand proteolytical degradation. These nanobodies can maintain their activity in the gastrointestinal tract and reduce ETECinduced diarrhea by inhibiting their adhesion in the gut [67]. However, orally delivered VHH still requires considerable stabilization as significant level of degradation occurs in the pigs’ stomachs [67]. Campylobacter jejuni is the leading cause of foodborne infection in the United States. C. jejuni infection results in acute diarrhea and can also result in a neurological disorder known as Guillain-Barre´ syndrome. According to CDC, each year Campylobacter is responsible for approximately 845,000 illnesses in the United States. Poultry is the major reservoir and source of infection for the Campylobacter. A previously characterized nanobody (FlagV1M) targeting C. jejuni flagella was engineered for greater thermal and proteolytic stability. A mutated version of the FlagV1M nanobody was obtained using error-prone PCR followed by panning under protease pressure. Using this combined approach, a hyperstabilized nanobody (F23-DSB) was created and this hyperstabilized nanobody was capable of potentially inhibiting C. jejuni [68]. Further, in vivo potential of the nanobody in reduction of C. jejuni colonization

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in the chicken GI tract was also evaluated [69]. Two C. jejuni flagella-specific nanobodies were pentamerized resulting in a pentabody complex capable of binding the target with high affinity and restricting the motility of the pathogen. Administration of this pentabody complex in two-day-old chickens significantly reduced C. jejuni colonization levels in the ceca of chickens [69]. Clostridium botulinum is anaerobic sporeformer and causative organism for botulism. C. botulinum produces a very potent neurotoxin known as botulinum neurotoxins (BoNTs). This toxin based on serological specificity is divided into seven categories (A, B, C, D, E, F, and G), out of which humans are affected by toxin A, B, and E. Administration of antitoxins and supportive care is provided for treatment. Baghban Roghayyeh et al. [70] expressed a BoNT/E-specific nanobody in a eukaryotic expression system using Pichia pastoris. The BoNT/E toxin neutralization potential of the purified nanobody was evaluated using a mouse model. Upon injection, the P. pastoris nanobody 20 and 40 μmol/g prolonged animal survival twofold. Flagellum plays an important role in bacterial pathogenesis and is also important for the first stage of biofilm formation. Pathogenic bacterial cells can form a biofilm on the biotic and abiotic surfaces, which allows them to resist sanitation treatment and bacterial cells in biofilm are more resistant to antibiotic treatments. Flagella of Pseudomonas aeruginosa play an important role in biofilm formation. Therefore a nanobody targeting P. aeruginosa flagella (antiflagellin) could inhibit bacterial motility and biofilm formation, thereby preventing infection. Adams et al. [71] reported an antiflagellin nanobody that has nanomolar affinity towards P. aeruginosa flagella. In vitro tests of this antiflagellin nanobody showed capability to prevent biofilm formation and inhibit its motility. Using similar approach, Payandeh et al. [72] developed a nanobody

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(DE3) to target the conserved region of the biofilm associated protein (Bap) of Acinetobacter baumannii, which causes bacteremia, meningitis, pneumonia, and other infections in humans. An in vivo neutralization assay using male BALB/C mice models demonstrated 100% survival of mice at the sixfold LD50 level. Surviving mice showed no signs of infection [72].

10.11 CONCLUSION Research in the field of nanobodies and their in vivo applications is gaining more attention. However, there are still some limitations associated with nanobodies. Production of nanobodies in a bacterial- or a yeast-expression system is an economical process. However, developing a specific nanobody starting from injecting a llama to affinity panning of positive clones is an expensive process, which restricts the development of new nanobodies. Another major limitation of unmodified nanobodies is their inability to cross the cell membrane [73]. This limits targeting of intercellular epitopes with a nanobody. Another potential challenge associated with nanobodies is their degradation during passage through gastrointestinal tract environment, which limits their application as orally administered therapeutics. Nanobodies are very similar to VH domain of conventional antibody. Their smaller size allows them to bind hard-to-reach antigenic sites. They also contain cysteine, allowing them to form disulfide bonds for additional stability. Nanobodies are versatile molecules with unique properties, which make them suitable for diagnostic and therapeutic applications. Next generation nanobodies will overcome few of these limitations and will be more compact, stable, and effective for in vivo applications in future.

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C H A P T E R

11 New Micro- and Nanotechnologies for Electrochemical Biosensor Development Nan Hao1,2, Jinwen Lu1, Rong Hua1, Wei-Wei Zhao2,* and Kun Wang1,* 1

Key Laboratory of Modern Agriculture Equipment and Technology, School of Chemistry and Chemical Engineering, Jiangsu University, Zhenjiang, P.R. China, 2Key Laboratory of Analytical Chemistry for Life Science and Collaborative Innovation Center of Chemistry for Life Science, School of Chemistry and Chemical Engineering, Nanjing University, Nanjing, P.R. China

11.1 INTRODUCTION Over the past few decades electrochemical biosensors have attracted much attention and made enormous progress. Various electrochemical analytical methods such as amperometric sensors, impedance sensors, electrochemiluminescence (ECL) sensors, and photoelectrochemical (PEC) sensors have been developed and can provide low-cost, easyoperation, sensitive, and selective detection performance [1]. Electrochemical biosensors are mainly composed of two elements, the biological recognition and the transducer elements. The biological recognition elements— including enzymes, proteins, antibodies, nucleic acids, cells, and others—can selectively react with the target. And the transducer elements can produce a potential or current signal that is

related to the concentration of the targets [2]. Because biomaterials are used as the sensitive element of the sensors, electrochemical biosensors have the characteristics of good selectivity, high sensitivity, fast analysis speed, as well as low cost, and can perform online continuous monitoring in complex systems. The working principle of the electrochemical biosensor is that the analyte passes through the diffusion into the biosensitive-membrane layer. The biochemical reaction occurs after the molecular recognition, the generated information is converted into an electrical signal related to the concentration of the analyte by the corresponding transducer. Based on these principles and advantages, some research results have been applied in the food industry [3,4], clinical testing [5
15], pharmaceutical industry [16
19], environmental analysis [20,21], and other fields.

*Corresponding authors

Advanced Biosensors for Health Care Applications DOI: https://doi.org/10.1016/B978-0-12-815743-5.00011-1

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The rapidly developing micro- and nanotechnologies have made a significant contribution to the improvement of electrochemical biosensors. Nanomaterials with sizes ranging in 1
100 nm have unique properties because of their dimensions and nanostructure. They may be classified as nanoparticles (NPs), nanorods, nanotubes, nanosheets, nanowires, and others according to their shape. Depending on their different compositions, there are noble-metal nanomaterials, semiconductor nanomaterials, ferromagnet nanomaterials, carbon-based nanomaterials, rare-earth nanomaterials, among others. With good conductivity, catalytic ability, large specific
surface area, and unique optical properties, nanomaterials have been widely applied in electrochemical biosensors. The large specific
surface area is beneficial to load more biological recognition probes onto the electrode. Good conductivity and catalytic ability can effectively amplify the signal. There is no doubt that more probes and stronger signal change allow bring better sensitivity and selectivity. For example, applications of gold NPs and carbon nanotubes (CNT) in electrochemical biosensors have been extensively studied since the beginning of the 21st century [22]. Wang et al. found that singlewall and multiwall carbon nanotubes (SWCNT and MWCNT)-modified glassy carbon electrodes can promote electron-transfer reactions and decrease the overvoltage of NADH oxidation [23]. Limoges et al. developed a sensitive electrochemical immunoassay using a colloidal gold label. Gold NPs were dissolved in an acidic solution and indirectly determined by anodic stripping voltammetry at a single-use, carbon-based, screen-printed electrode for immunoglobulin G (IgG) detection [24]. The performance of various nanomaterials in electrochemical biosensors has been investigated. In this chapter, we mainly summarize the progress of nanomaterials in the field of biosensor made

over the past decade, such as graphene, graphitic carbon nitride (g-C3N4), and quantum dots. Besides nanomaterials, the microand nanofabrication technologies also greatly improved traditional electrochemical biosensors. Next we briefly introduce the applications of microfluidics chips in electrochemical biosensing.

11.2 MICROFLUIDICS CHIPS Microfluidics technology focuses on the research and applications of devices manipulating fluids in the range of μL
aL (1026 to 10218 L) [25]. Through the miniaturization and integration of analysis equipment, various microfluidics chips have been developed to integrate the whole chemical processes, including sample pretreatment, reaction, separation, analysis into a chip of a few centimeters scale or smaller, which are also called “lab on a chip.” Biological processes such as cell cultures, cell lyses, and protein crystallizations also can be achieved on a chip using a microchannel network. As an interdisciplinary field of chemistry, physics, nanotechnology, engineering, and biotechnology, microfluidics has attracted much attention [26]. This technology can provide a powerful tool for researchers to manipulate and regulate the microenvironment, which is difficult to achieve with other conventional methods. Different functional units may be flexibly combined according to the demand and multiple independent parallel channels can be easily integrated onto one chip for high-throughput analysis. Because of their small size and high integration, microfluidic chips have a number of advantages such as the ability to use very small quantities of samples and reagents, low cost, short time for analysis, integration for analytical devices, and small size for the control of concentrations of molecules in space and time. By miniaturizing the analytical

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system, this technology has been widely applied in electrochemical biosensors and has made great progresses in food safety, environment monitoring, clinical diagnostics, and others. Rackus et al. discussed how electrochemistry, biosensors, and microfluidics can be combined to form four new application areas: (1) electrochemistry and microfluidics; (2) electrochemical biosensors; (3) microfluidic biosensors; and (4) microfluidic electrochemical biosensors [27]. It is worth noting that microfluidic paper-based analytical devices (μPADs) was proposed by Whiteside et al. in 2007 [28,29]. Compared with microfluidic analytical devices fabricated with silicon, glass, or polydimethylsiloxane (PDMS) the μPADs are affordable, user friendly, do not require external instruments, and complex fabrication processes, which provide a disposable and low-cost platform for prototyping new pointof-care testing (POCT) devices [30,31].

11.2.1 Microfluidics Chips in Biomarkers Detection Sensitive disease-relevant biomarkers detection is very important for clinical diagnostics and therapy [32]. Although current advanced diagnostic technologies have good determination performance, their drawbacks include sophisticated instruments and relatively high costs, which limits their application in lessdeveloped areas or for POCT. μPADs comprised of single or multiple layers of paper substrates can transport fluids autonomously for performing multistep analytical assays through hydrophilic paper channels on the substrate. Various analytes in human fluids (e.g., urine, serum, blood) can be quantified with the advantages of low cost, portability, and ease of operation. Zhao et al. reported a microfluidic paper-based electrochemical biosensor array for multiplexed detection of physiologically relevant metabolic biomarkers [33]. Most existing μPADs need a benchtop

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potentiostat to read the output signal, which is expensive and unfavorable for POCT or other resource-poor environments. Meanwhile, commercially available portable potentiostats usually have a relatively high price. Another choice is to use a glucose meter, which is a highly mature and inexpensive commercial electrochemical reader, and can also be applied for the detection of different types of analytes other than glucose. But it only accommodates one paper sensor at a time and, therefore, cannot meet the demand of massive determinations. To improve detection efficiency and further reduce the instrumental demand, they developed a handheld custom-made electrochemical reader for signal readout. Integrating with an array of eight electrochemical sensors, the obtained electrochemical devices can achieve the detection of multiple analytes in a single run. To demonstrate feasibility, they used the device to detect glucose, lactate, and uric acid in urine simultaneously. The analytical performance was comparable to that of other commercial platforms. Rusling et al. demonstrated the great promise of microfluidic electrochemical devices for high-throughput multiplexed detection (Fig. 11.1) [34]. First they developed a low-cost, high-throughput electrochemical array featuring 32 individually addressable microelectrodes. Submicroliter hydrophobic wells surrounding each sensor were prepared to prevent potential cross contaminations during antibodies immobilizations. Then 8 prepared electrochemical arrays were connected with a miniaturized 8-port manifold to obtain a 256-sensor system. This system was used to determine prostate cancer biomarker proteins, namely prostate-specific antigen (PSA), prostate--specific membrane antigen, interleukin-6 (IL-6), and platelet factor-4 (PF-4) in serum. With the precapture of enzymes and antibodies labeled magnetic nanoparticles, 256 measurements could be finished in less than 1 h with a good sensitivity (0.05 pg/mL) and a wide dynamic range (pg/mL to ng/mL).

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FIGURE 11.1

A schematic of high-throughput electrochemical microfluidic immunoarray for multiplexed detection of cancer biomarker proteins. Reprinted with permission from C. K. Tang, A. Vaze, M. Shen, J.F. Rusling, ACS Sens. 1(8) (2016) 1036
1043. Copyright 2016 American Chemical Society.

To evaluate the cardiotoxicity and other cardiac safety issues of newly developed drugs, it is important to continually monitor secreted biomarkers during long-term cultures in the biomimetic human organoid models that simulate both the biology and the physiological microenvironment of the human system to understand their responses to drug exposure in a noninvasive manner. Traditional biosensing techniques, such as the current gold standard enzyme-linked immunosorbent assays (ELISA), have limitations in sensitivity, selectivity, stability, and require large working volumes of reagents. The complex biological environments of cell culture media usually contain various interfering compounds and trace amounts of biomarkers of interest. The low abundance targets bring a critical challenge. To solve this problem, Shin et al. reported a novel, label-free, microfluidic, aptamer-based, electrochemical biosensing platform for monitoring damage to cardiac organoids [35]. This platform was composed of an in-house designed gold (Au) microelectrode and a microfluidic system. The performance was assessed by the trace changes (,1 ng/mL) in secreted creatine kinase (CK)-MB levels in a

heart-on-a-chip system upon drug insults in a dose-dependent manner measured by the aptamer-based electrochemical immunosensors. The detection was in agreement with the beating behavior and cell viability analyses, which demonstrates that the developed system is versatile and can be used to detect various cell-secreted biomarkers. Shin et al. developed a label-free, electrochemical, microfluidic biosensor for continual monitoring of cell-secreted soluble biomarkers human albumin and glutathione-S-transferase-alpha [36]. This platform has the ability of regenerating the sensing interface through the treatment of H2SO4 and K3Fe(CN)6 with the application of electrical sweep. Integrated with microfluidic valves and a human liver-on-a-chip platform, this system was capable of continual monitoring for up to 7 days with few manual operations. The bipolar electrode (BPE), a conductor immersed into the solution, is a recently developed tool in ECL biosensors [37,38]. With an applied external voltage across the solution, the faradaic reactions at the anode and cathode of the BPE may be induced respectively without the need of direct connections between the electrode and power source [39]. This improvement

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simplified the fabrication processes of the microfluidics ECL chips and is beneficial for portable ECL devices. Wu et al. extensively studied the application of BPE-ECL biosensors in the detection of cancer biomarkers. For example, they designed a novel wireless ECL biosensor for the visual detection of PSA using a sandwich-type strategy [40]. The ECL device consists of two ITO electrodes embedded in a two-channel microfluidic chip. The electronic conductivity of the gap between the two ITO electrodes could be tuned by PSA-guided silver particles deposition. The target-guided electrical switch made two neighboring BPEs behave like a continuous electrode, which resulted in a decreased activating voltage for the oxidation reactions of Ru(bpy)321. The experiment demonstrated that the ECL signal was totally activated by the target, while the background ECL signal was very weak. On this basis, they further developed a novel ECL array platform for simultaneous detection of multiple cancer biomarkers [41]. A detection channel array and sensing channel array were connected by a group of parallel BPEs on a glass substrate, and the cathodes of the BPE array were modified with electrochemically deposited Au films and various biorecognition elements. Adenosine triphosphate (ATP), α-fetoprotein (AFP), PSA, and thrombin were detected through faradaic reaction
mediated ECL signals. They developed a multicolor ECL sensor via modulating the resistance of a closed BPE device, [Ru (bpy)3]21 and [Ir(ppy)3] as luminophores at the anode was selectively excited by the changing interfacial potential and the emission color was finely tuned in the range of green
yellow
red (Fig. 11.2) [42]. A sensitive ECL biosensor for PSA in serum samples was built based on the closed BPE system with multicolor ECL emissions. The green color of ECL emission means the concentration of PSA is at normal level while the red color indicates a high risk.

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FIGURE 11.2 The illustration of the closed BPE system with multicolor ECL emissions. Reprinted with permission from Y.Z. Wang, C.H. Xu, W. Zhao, Q.Y. Guan, H.Y. Chen, J. J. Xu, Anal. Chem. 89 (15) (2017) 8050
8056. Copyright 2017 American Chemical Society.

11.2.2 Microfluidics Chips in Nucleic Acid Detection Micro total analysis systems for use in onsite rapid detection of DNA or RNA are increasingly being developed [43]. Crooks et al. developed a strategy for the determination of oligonucleotides using a paper analytical device (PAD) [44]. The device consisted of two parts, a double-sided printing of a wax pattern onto chromatography paper as the base layer and a slip layer with a hole to expose the electrode. Three carbon electrodes were printed onto the double wax-patterned paper. Then a layer of Ag/AgCl paste was painted on top of the reference electrode and Au was electroplated onto the working electrode. The detection was based on the principle proposed by Plaxco et al. that the target would cause the conformation change of oligonucleotide probes. The probe was modified on the gold electrode through the thiol group and

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linked an electrochemical label (such as methylene blue or ferrocene) at the distal end. The target-induced folding or unfolding of probes altered the location of the electrochemical label relative to the electrode surface and resulted in a change in faradaic current. This design provided a simple and robust method for the mass production of paper electrochemical biosensors. Ghodssi et al. presented a microfluidics device with arrayed electrochemical sensors for the analysis of DNA hybridization events [45]. DNA hybridization detection with microfluidic chips usually have limitations in sample preparation and mixing due to the low sample volume and low Reynolds number. Some physical and chemical effects such as capillary forces, surface roughness, and reduced surface area and volume restricted diffusion, and mass transport decreased probe surface density, affected the interactions between probes and analytes, and brought low electrochemical signal-to-noise ratio. To overcome these limitations, they developed a new dual-layer microfluidic manipulation system for high-throughput analysis with automated control. The chip was comprised of two PDMS layers. Assay microchannels were on the bottom layer and valves were on the top layer. Single-stranded DNA (ssDNA) probes were modified onto Au electrodes in the individually addressable reaction chambers array. The hybridization between the probe and ssDNA target was detected with EIS, which allowed to evaluate charge transfer and diffusion resistance effect. This design can provide accurate programmable and automated capability for high-throughput ssDNA analysis with a theoretical limit of 1 nM. MicroRNAs (miRNA) are valuable biomarkers for clinical diagnosis. Sensitive detection of multiplex miRNAs is of great theoretical significance and practical value. Yu et al. reported a microfluidic paper-based electrochemical analytical device for sensitive

detection of microRNA [46]. μPADs were modified with Au nanorods for the immobilization of hairpin probe DNA. The probe hybridized with target miR-21 and opened the hairpin structure. The retained single-strand fragment of hairpin probe further hybridized with electrochemical probe labelled capture DNA. Ru (NH3)631 as electron mediator was adsorped into the DNA duplex to improve electron transport for enhancing electrochemical signals. The paper-based electrochemical biosensor showed a good detection performance with a wide linear range (1.0
1000 fM) and a low detection limit (0.434 fM) for miR-21 detection. Shamsi et al. integrated ECL with digital microfluidics, which was a powerful liquid handling platform for controlling discrete droplets on an array of insulated electrodes [47]. Combined with magnetic particle
based nucleic acid hybridization assays, the detection of miRNA can be achieved with sample volumes of 1.8 μL. With such a small sample volume, this system had good sensitivity with a detection limit of 1.5 fM. Moreover, the system has the ability of detecting miRNA-143 in breast cancer cell lysates to discriminate different cell types. To increase the detection sensitivity, a series of nucleic acid amplification technologies have been developed. Integrating these amplification processes into a chip has the advantage of fewer manual steps, preventing contamination, and greatly reduced reagents consumption. Kim et al. developed a microfluidics chip for performing cell lysis, polymerase chain reaction (PCR), and quantitative analysis of DNA amplicons in a single step [48]. This chip included three parts, an electrochemical cell lysis zone, a continuous-flow PCR module and capillary electrophoresis amperometric detection system built with glass substrate using an indium-tin-oxide (ITO). Genomic DNA in the reagentless electrochemical cell lysis product was successfully analyzed after PCR reactions.

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11.2.3 Microfluidics Chips in Bacteria Detection With more and more concern on food safety, the contamination of pathogen bacteria in food and water has attracted considerable attention. The traditional culturing method is time-consuming and labor-intensive, making it necessary to develop portable biosensors for bacteria detection [49
51]. Jiang et al. developed a low-cost and miniaturized microfluidics device for sensitive bacteria detection [52]. The detection was based on the electrical impedance spectroscopy (EIS) method conducted via a smartphone. Smartphones have a high popularizing rate and are powerful data collectors. The device was composed of sensing electrodes, microhole array silicon substrate and a sensing microfluidic chamber bounded by a nanoporous filter paper. Escherichia coli was proconcentrated on the chip and analyzed with a specifically designed impedance network analyzer. The real-time EIS data was transmitted to a smartphone via a Bluetooth circuit module, and the smartphone controlled the detection and showed the results. Altintas et al. designed a fully automated microfluidic-based electrochemical biosensor for real-time pathogen detection [53]. Detection was based on the direct proportion between the HRP-labeled secondary antibody and the target E. coli. The measured HRP
TMB enzymatic reaction electrochemical response increased with the bacterium concentration increase. With antibodies adsorbed AuNP amplification, the sensitivity was greatly enhanced. A detection limit of 50 cfu/mL for E. coli detection was achieved while that of a standard sandwich assay was 1.99 3 104 cfu/mL. Traditional microfluidics fabrication methods are expensive and need complex processes, which were not suitable for commercially available chips. The chip surface could be regenerated multiple times using 0.1 M HCI (1 min, 120 μL/min), which significantly reduced the cost.

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Pathogens can also be detected through the genetic analysis. Zourob et al. designed a microfluidic electrochemical biosensor for performing loop-mediated isothermal amplification (LAMP) to achieve sensitive detection of E. coli [54]. LAMP is a fast and cost-effective isothermal DNA amplification technique. Because of the precise control of small amount of liquid, microfluidics can effectively reduce the consumption of expensive biochemical reagents. DNA extracted from E. coli was amplified through LAMP reactions and Hoechst 33258 redox molecules were bound to the DNA minor groove, which caused a significant drop in the anodic oxidation of linear sweep voltammetry (LSV). The detection limit was as low as 24 cfu/mL of bacteria. Hsieh et al. also designed a microfluidic electrochemical quantitative loop-mediated isothermal amplification (MEQ-LAMP) system for POCT [55]. The chip contained only a chamber with three electrodes (around 20 μL) that served both as the LAMP reaction vessel and electrochemical measurement cell. The real-time monitoring of the LAMP reaction was achieved via the redox electron transfer and interaction of methylene blue (MB). The interaction of MB into double-stranded regions of LAMP products reduced free-MB concentration in the solution and caused the decrease of the redox current. This system has the capability of detecting 16 copies of Salmonella typhimurium genomic DNA in less than an hour.

11.2.4 Microfluidics Chips for Small Molecule Detection There are many reports on applications of microfluidics biosensors for small molecule detection. An important application field is environment monitoring. Adam et al. provided a detailed summary about microfluidic electrochemical devices for pollution analysis [56]. Typical examples included phenolic

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compounds, pesticides, and herbicides [57
60]. Caetano et al. developed an electrochemical microfluidic biosensor combined with commercial textile threads for phenol detection [61]. This design can fully take advantage of capillary phenomena and gravity forces to promote the solution transportation without the need of external forces or an injection pump. Tyrosinase (Tyr), a polyphenol oxidase (PPO), can catalyze the oxidation reactions of phenolic compounds and was modified onto the screen-printed electrodes. The concentration of the target may be evaluated from the cyclic voltammetry signal. Mycotoxins, as the toxic secondary metabolite of filamentous fungi, exist widely in agriculture products and are a great threat to human health [62,63]. Traditional mycotoxin detection methods such as gas chromatography (GC), mass spectrometry (MS), liquid chromatography (LC) or ELISA require expensive equipment and time-consuming procedures. To reduce these costs, Olcer et al. developed a new sensing platform called realtime electrochemical profiling (REP) for mycotoxin detection [64]. This technology is mainly based on electrochemical immunoassay detection. This system is composed of an electrode array, microfluidics-based assay and real-time amperometric measurements, which combined the enzyme immunoassay and reaction solution detection on one electrode surface. The whole assay was conducted during the reagents flow without the need of incubation time. Therefore this system can provide fast detection. The performance of deoxynivalenol (DON) detection in a wheat sample showed satisfactory sensitivity and recovery rate. Pharmaceutical detection is essential in clinical drug therapy and drug residue monitoring. Droplet-based microfluidic systems in which small droplets encapsulated by an immiscible phase are generated and manipulated have unique advantages. The volume of droplets can be precisely controlled with a

high generation rate. Each droplet is well isolated and can be regarded as an individual micro/nanoreactor. So droplet technology can effectively avoid surface adsorption and crosscontamination and provide high efficiency mixing [65], which is an ideal platform for high-throughput analysis. Coupled with chipbased carbon paste electrodes (CPE), dropletbased microfluidics was developed for the determination of dopamine (DA) and ascorbic acid (AA) in intravenous drugs with chronoamperometric detection [66]. But this platform was not as sensitive as previous reports. The modification of graphene-polyaniline (GPANI) on CPE is beneficial for the current response enhancement. A high-throughput and sensitive droplet-based microfluidic sensor using a G-PANI modified electrode was developed for the 4-aminophenol (4-AP) detection in pharmaceutical paracetamol (PA) formulations [67]. Liu et al. established a sensitive electrochemical microfluidics platform for therapeutic drug monitoring [68]. The working electrode was fabricated with an Au
Ag alloy microwire and electropolymerized molecularly imprinted polymer (MIP) film. Combined with a molecularly imprinting technique, they reported a novel strategy based on the gate effect. Briefly, without the presence of target warfarin sodium (WFS), cavities in the MIP film were in an open state and probe ions could reach the electrode surface to generate electrochemical signals. Once targets were bound by MIP, the signal was greatly affected due to increased steric hinder. The strategy showed a good performance with a liner range from 2 3 10211 to 4 3 1029 M. The abuse of antibiotics in human and veterinary medicine caused the generation and spread of multidrug-resistant bacteria, which has become a serious public health problem. Traditional microbiological antibiotic tests are based on the inhibition effect of antibiotics on bacterial growth in the laboratory, are timeconsuming, and cannot meet the demand of

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POCT. It is necessary to develop a low-cost, rapid, and sensitive sensing platform for multiple antibiotics detection for medical analyses and environmental and food control. A electrochemical microfluidic platform based on eight enzyme-linked assays for simultaneous multiple antibiotics detection was reported [69]. To demonstrate the feasibility, two commonly employed antibiotics (tetracycline and streptogramin) in a human plasma sample were detected. In the presence of the target, the antibiotic sensitive
repressor protein (TetR or PIP) would have a conformational change and cannot bind to its designated operator DNA. These two repressor proteins were biotinylated to facilitate a subsequent binding to an avidin-glucose oxidase (GOx) conjugate. GOx catalyzes the reaction from a glucose substrate to hydrogen peroxide (H2O2) and generates an electrochemical signal. Thus the more antibiotics presented, the fewer proteins bind to their operator DNA and less GOx immobilized on the electrode, resulting in a decreased signal. The total assay time was around 2 h 30 min and the detection limit was around ng/mL.

11.3 QUANTUM DOTS Quantum dots (QDs) are an important lowdimensional semiconductor material and their sizes are typically between 1 and 10 nm. The initial research began in the early 1980s. Brus et al. [70,71] found that different sizes of cadmium sulfide particles can produce different colors [70,71]. This work was very helpful to understand the quantum confinement effect, which explains the relationship between QD size and color, and also paved the way for the application of QDs. In addition, QDs have unique physical and optical properties such as a broad absorption band, a narrow emission band, tunable emission from the visible to the near infrared (NIR) range, excellent brightness, and long fluorescence lifetime. They are highly

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resistant to photobleaching and have a relatively large surface area for biofunctionalization. Based on these advantages, QDs had been widely studied and applied over the past 30 years, and with the continuous improvement of preparation technology, QDs have become more and more likely to be applied to biological research. There are some reviews on the application of QDs in biological analysis [72
76]. As shown in Scheme 11.1, we summarize the applications of QDs in electrochemical biosensing over the past five years to provide readers with a faster and clearer understanding of this field and discuss them using three aspects: (1) electrochemical enzyme biosensing based on QDs; (2) electrochemical gene biosensing based on QDs; and (3) Electrochemical immunosensing based on QDs.

11.3.1 Electrochemical Enzyme Biosensing Based on Quantum Dots The interaction between the active substance and the enzyme-catalyzed reaction can be converted to photoelectric signal changes under certain conditions. Based on this principle, the construction of PEC enzyme biosensors can be divided into two methods. One approach is to use an enzyme-catalyzed reaction to produce insoluble substances on the surface of the electrode and form physical effects such as steric hindrance, which can hinder the electron transfer process between the PEC active substance and the electron donor or acceptor in the solution, thereby causing changes in the photoelectric signal. The second approach is to use the soluble products of the enzymecatalyzed reactions directly as electron donors or acceptors, which will cause the system to generate corresponding electrical signals. Therefore this method can be used to detect enzyme substrate consumption [77,78] and monitor enzyme activity [79,80].

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SCHEME 11.1 Schematic illustration of electrochemical biosensors based on quantum dots, which mainly displayed enzyme-based biosensors, gene biosensors, and immunosensors.

Zhang et al. proposed a photoelectrochemical enzyme
sensing method for glucose based on low toxicity Ag2S QDs. Oxygen is sensitive to PEC biosensors and can act as an electron acceptor and generate a cathodic photocurrent during photoelectrochemical processes [81]. In addition, experimental comparisons found that O2 is more effective than H2O2 in PEC detection. Therefore when the electrode

was exposed to the solution containing glucose, the covalently linked glucose oxidase (GOD) could catalyze the conversion of glucose to gluconic acid and H2O2. The dissolved oxygen is consumed in this process, resulting in a decrease of photocurrent, thereby enabling efficient detection of glucose. Similarly, Wang et al. described a novel PEC enzyme
sensing method to detect glucose based on a CdS QDs

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sensitized p-type NiO photocathode [82]. Oxygen was used as an effective electron acceptor for irradiation of CdS QDs to further increase the charge separation efficiency of CdS QDs. Oxygen served as an effective electron acceptor for irradiation of CdS QDs to further increase the charge separation efficiency of CdS QDs. When glucose is present in the solution, GOD consumed O2 to catalyze the oxidation of glucose causing a change in the photocurrent. Besides these, some researchers have used the coupling of three enzymatic reactions (guanylate kinase, pyruvate kinase, lactate dehydrogenase) to separate and transfer electron
hole pairs under light excitation of CdS/ZnS-based QDs. It worked to detect guanosine monophosphate [83]. By combining CdSeTe@CdS@ZnS QDs with TiO2 thin films or CdS QDs, the alkaline phosphatase catalyzes the production of an excellent electron donor from the substrate trisodium 2-phosphate-1-ascorbate, and two sensors for detecting caspase-3 activity [84] and 5-hydroxymethylcytosine [85] were constructed, respectively. Hou et al. proposed a sensor based on the in situ formation of homogeneous solution CdS QDs and the enzymatic reaction of acetylcholinesterase (AChE), and the AChE activity and inhibitory effect in the optimal state were determined [86]. ECL enzyme biosensing involves the emission of a chemical reaction between the reaction products at the electrode surface or between the reaction products and the substances in solution under an applied potential. When the target was present, under the action of the enzymatic reaction, the substances that blocked chemiluminescence was consumed or generated, thereby changing the intensity of the luminescent signal to achieve the purpose of detection. The construction of glucose biosensors has attracted the attention of the majority of researchers in this field. The development of third-generation glucose biosensors in the enzyme-catalyzed reaction using direct

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electron transfer (DET) was developed since 2000s. In recent years, many research groups have proposed ECL enzyme biosensors based on graphene QDs to detect glucose [87,88]. Zhou et al. proposed GQDs=SO22 3
based ECL detection of H2O2 and a glucose sensing system [89]. In the presence of SO22 3 , GQDs can produce strong cathode ECL light. Then, H2O2 can significantly inhibit the ECL of the GQDs=SO22 system because of the oxidation 3 of coreactant (SO22 3 ) with H2O2 to a noncoreactant (SO22 ). Therefore glucose can be detected 4 by using glucose oxidase (GOx) to convert glucose and O2 to gluconic acid and H2O2, respectively. Liang et al. reported the use of a biological function AChE-GNs-CdTe QD sensing platform for OPs ECL analysis [90]. Without target substances, the ECL signal was reduced due to the consumption of coreactant dissolved oxygen by acetylcholinesterase. And the target can inhibite the activity of acetylcholinesterase, resulting in a gradual recovery of the ECL signal. Stewart et al. designed an electrochemical enzyme biosensor that generated ECL signals through CdSeTe/ZnS QDs interaction with H2O2 produced in ChOx-catalyzed cholesterol oxidation and can be used for cholesterol monitoring [91]. In addition, a scheme for detecting Pb(II) by using a deoxyribozymebased biosensor has also been developed. Target Pb(II) may combine with T30695 to form G-quadruplex that would bond to the hemoglobin to form DNase with peroxidaselike activity. Because DNase can consume H2O2, the coreactant of P-GO @CdS QDs, the ECL signal was reduced [92].

11.3.2 Electrochemical Gene Biosensing Based on Quantum Dots A PEC gene biosensor is a device that integrates the probe DNA as a biorecognition element, light is used as the excitation source, and photocurrent is used as the detection

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signal. A novel PEC DNA biosensing strategy based on PbS QDs was developed [93]. The cathodic photocurrent of PbS QDs was sensitive to dissolved oxygen and was not affected by H2O2. Moreover, the Pt NPs/G-quadruplexes/hemoglobin complexes encoded by biobarcodes have an effective catalase (CAT)-like catalytic activity and, therefore, can be used to decompose H2O2 to release O2 for signal amplification. When the target DNA was present, one end hybridized to the capture probe immobilized on the PbS QDs and then the other end of the target DNA hybridized to the biobarcode Pt/G-quadruplex/hemoglobin to form a sandwich structure. Pt/G-quadruplex/hemoglobin catalyzed the decomposition of H2O2 to release O2, which acts as an electron acceptor to promote the generation of a stronger cathodic photocurrent by PbS QDs. The sensor had good stability, selectivity, and reproducibility. Yan et al. constructed a new PEC biosensing platform by assembling CdSe QDs and DNA on liquid-phase deposited TiO2 thin film electrodes for o-aminophenol (OAP) detection [94]. Briefly, CdSe can effectively increase the visible light absorption of the TiO2 film. Due to the interaction between the DNA and OAP, the OAP response can be increased by immobilizing DNA on the sensing membrane. Recently, there have been numerous reports on the construction of DNA biosensing strategies based on CdS QDs, such as a strategy for direct detection of miRNAs [95], the detection of carcinoembryonic antigen cascade secondary signal amplification strategy [96], the endonuclease-assisted loop amplification strategy for DNA detection [97], and others. Zhang et al. designed a new dual-channel self-reference biosensor [98]. In this biosensor, CdTe and CdTe-graphene oxide (GO) were selected to form two adjacent working regions, and they were functionalized with Aflatoxin B1 (AFB1) aptamers. The cathode current of

CdTe-GO and the anode current of CdTe can be well distinguished by adjusting the bias voltage during the experiment. Finally, by applying the “signal on” and “signal off” models at the same time, double-concentration information can be obtained in one detection and used as a reference for each other. This DNA biosensor could eliminate irrelevant interference factors, improve the accuracy of detection and make the monitoring results more reliable. Therefore it is promising for the future development of PEC biosensing. Chen et al. reported a novel electrochemical bioanalysis platform based on semiconducting organic
inorganic nanodot heterojunctions [99]. This work provides a direction for future PEC bioanalysis. Another related work is a DNA biosensor based on the energy transfer (ET) between semiconducting polymer dots (Pdots) and gold nanoparticles (Au NPs) in photoelectrochemical systems (Fig. 11.3). On this basis, various PEC systems based on ET effect of Pdots can be studied, which opens up new directions for the application of electrochemical biosensors to biological analysis. [100] It is well-known that miRNAs play an important role in the process of cell differentiation, biological development, and the development of diseases, and are increasingly attracting the attention of researchers. In recent years, Yuan et al. has developed the following series of QD ECL biosensing strategies to detect miRNAs: (1) a biosensor for ECL detection of miRNAs based on the intercalation of doxorubicin-conjugated CdTe QDs (Dox-QDs) into the DNA/RNA hybrids as a new signal acquisition and amplification platform [101]; (2) a novel dual amplification for miRNA assay based on graphene quantum dots as a sensing platform [102]; (3) an ECL resonance energy transfer (ERET) strategy based on small molecule dye (Alexa Flour 488) as a donor and CdSe@ZnS QDs as acceptors, followed by

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FIGURE 11.3 Schematic mechanism of the operating Pdots-Au NPs based PEC system. Reprinted with permission from X.-M. Shi, L.-P. Mei, Q. Wang, W.-W. Zhao, J.-J. Xu, H.-Y. Chen, Anal. Chem.90 (7) (2018) 4277
4281. Copyright 2018 American Chemical Society.

construction of a DNA nanomachine-based regeneration biosensor for detection of miRNA from cancer cells without any enzymes [103]; (4) a QDs-Ru(dcbpy)321 nanostructure containing Ru(dcbpy)321 as a donor and CdSe@ZnS QDs as ECL emitters was combined with the target cyclic amplification and dual output conversion strategy (based on this proposal, an ECL biosensor was constructed and applied to the detection of miRNAs) [104]; and (5) an ECL biosensor established with a 3D DNA walking machine based on target transformation as well as quenching and enhancement mechanisms for distancecontrolled signals, which enabled the detection of miRNAs [105]. As shown in Fig. 11.4A, the Ru(dcbpy)321 was loaded onto the surface of CdSe@ZnS QDs (QDs) through an amide reaction to form a nanostructure with the realization of ECL-RET. Then the capture probe 1 (capture1) was bound to the nanostructure surface by an amidation reaction to form a novel ECL signal tag of QDs-Ru(dcbpy)321/ capture1. Next, the hairpin DNA1 (H1) was

opened by the target miRNA-141, thereby obtaining dsDNA with some exposed bases, which hybridized with H2 to substitute the miRNA in the presence of hairpin DNA2 (H2), achieving the target miRNA recovery. Under the action of Bst 2.0 DNA polymerase, H1 began to polymerize with the H2 template, yielding fully complementary dsDNA with two endonuclease recognition sites. After the nicking endonuclease Nt.BbvCI was added, two specific cleavage sites were identified and an amount of reporter DNA was obtained (Fig. 11.4B). In Fig. 11.4C, one end of the obtained reporter DNA hybridized to capture DNA1 on the surface of the electrode, while the other end hybridized to QDs-Ru (dcbpy)321/capture1. Thus quantitative detection of miRNA-141 from human prostate cancer cells was achieved. Zhu et al. reported a sensitive ECL miRNA based on N-CQDs [106]. Hairpin probes 2-NCQDs with helper probes and miRNA formed a Y-linked structure that released miRNAs and helper probes by adding a nicking enzyme Nb.

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(A)

ECL-RET HOOC in one nanostructure N

C

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template strand reporter DNA

Output Target Recycling Nt.BbvCl

d c b a

polymerase

H1 target microRNA-141 Input

a: GCE b: AuNPs

c: HT d: capture 2

FIGURE 11.4 The construction of the miRNA biosensor based on the novel highly efficient ECL-RET in one nanostructure. (A) The preparation of QDs-Ru(dcbpy)321/capture 1 bioconjugate and the mechanism of ECL-RET in one nanostructure; (B) the dual amplification assay including target recycling and double-output conversion strategies; and (C) the establishment of the proposed biosensor. Reprinted with permission from Z. Li, Z. Lin, X. Wu, H. Chen, Y. Chai, R. Yuan, Anal. Chem. 89 (11) (2017) 6029. Copyright 2017 American Chemical Society.

BbvCI released miRNAs and assistant probes can initiate the next recovery process. The generated intermediate sequence nitrogen-doped carbon QDs-DNA was further hybridized with the hairpin probe 1 immobilized on the surface of the electrode, and the initial ECL intensity was enhanced. The ECL intensity increased with increasing concentration of the target miRNA. Zhang et al. successfully synthesized

a new luminescent material, boron-doped graphene QDs, by the electrolytic stripping method and constructed an ECL sensor for detecting miRNAs-20a based on the novel QD material [107]. In addition to these, there have been reports of the detection of DNA [108
110], hepatitis B and C viruses [111], EGFR genes [112], and others by constructing an ECL biosensor.

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11.3.3 Electrochemical Immunosensing Based on Quantum Dots The electrochemical immunosensor combined the high sensitivity of sensing technology with the specificity of the immune response, and converted the signal generated during the antigen
antibody-specific reaction into an electrical signal through a transducer, thereby quantitatively detecting the antigen or antibody. With the development of immunoassay technology and chemical sensing technology, more and more researchers are studying the construction of immunosensors. Sun et al. used 3-mercaptopropionic acid (MPA) as a modifier and bridged the immobilized antibody. The MPA-CdS/RGO (reduced graphene) oxide nanocomposites were synthesized by the solvothermal method through in situ growth of CdS nanoparticles on RGO sheets. Therefore a novel ultrasensitive PEC immunosensor for detecting indole-3-acetic acid was developed [113]. In detail, when CdS quantum dots absorbed energy higher than their bandgap photons, electrons were excited from the valence band (VB) to the conduction band (CB), forming electron
hole pairs. Immediately due to the CB of ITO had an energy level lower than that of CdS, then electrons were transferred to the ITO electrode and a photocurrent was generated because of the lower energy level of ITO. Holes can be trapped by the electron donor triethanolamine (TEA) in solution.

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When the insulated protein molecule Ab1 was immobilized on the CdS QDs-modified electrode surface, it inhibited the diffusion of TEA to the CdS surface and reacted with the photogenerated holes, resulting in a decrease in photocurrent. Biobarcode platinum nanoparticles-G-quadruplex/heme served as probes and bond to Ab2. Since the biobarcode had enzyme-like activity it can catalyze the oxidation of hydroquinone with H2O2 as an oxidant, while its oxidation product can act as an effective electron acceptor for irradiated CdS QDs, thereby increasing photocurrent to achieve signal amplification [114]. Moreover, Lan et al. constructed a PEC immunosensor for the detection of carcinoembryonic antigens based on multichannel signal amplification by multibranch hybridization chain reaction and PdAu enzyme mimetics in combination with a microfluidic paper analysis device [115]. Some groups also have reported new PEC immunosensors that detect prostate-specific antigens [116,117]. In recent years, there have been many reports on ECL immunosensing for the detection of carcinoembryonic antigens (CEA) [118
120], alpha fetoproteins (AFP) [121], and the simultaneous detection of CEA and AFP [122]. As shown in Fig. 11.5, the CEA antibody (Ab1(CEA)) and the AFP antibody (Ab1(AFP)) are simultaneously covalently linked to the modified electrode surface. CdSe nanocrystals (CdSe NCs) and CdTe nanocrystals (CdTe FIGURE 11.5 Schematic illustration of spectrumresolved, dual-color ECL immunoassay with nanocrystals as tags. Reprinted with permission from G. Zou, X. Tan, X. Long, Y. He, W. Miao, Anal. Chem. (2017). Copyright 2017 American Chemical Society.

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NCs) were used as labels for the CEA probe antibody (Ab2(CEA)|CdSe) and the AFP probe antibody (Ab2(AFP)|CdTe), respectively, to form a sandwich structure in the presence of the target CEA and AFP. In the presence of the target, it formed a sandwich structure with Ab1 and Ab2. In addition, (NH4)2S2O8 can be used as a coreactant for CdSe and for CdTe and it was, therefore, used as a coreactant for the sensing system. CdTe and CdSe NCs in NC-immune complexes can be electrochemically reduced while emitting monochromatic ECL emission in the near infrared and green regions, respectively. Because the ECL spectra of the two surface-encapsulated NCs can be well separated and there was no cross-energy transfer interaction, a spectrally resolved dual-color immunosensor can be successfully constructed to simultaneously detect CEA and AFP [123]. There were other groups that developed ECL immunosensors for the detection of lead ions [124], brombuterol [125], insulin [126], and laminin (LN) [127]. The size of the Ag2S: Mn QD was much smaller than the average size of the antibody protein, which made the ECL intensity produced by the QD as a signal marker in the antibody very small. Bovine serum albumin (BSA) was a good linking agent or stabilizer for the construction of luminescent nanomaterials or signaling probes. Combining the above two points, BSA was attached to the surface of Ag2S:Mn QDs to form BSA-Ag2S:Mn bioconjugate as well as being used as a new signal tracer and immobilizing the LN antigen on its surface. When LN was present, LN antibodies on the surface of the electrode captured the antisource LN via a sandwich immunoreaction, resulting in a stronger ECL signal [127]. It is worth noting that the above QD-based electrochemical biosensors utilized the excellent physical and chemical properties of QDs, while the QDs did not participate in chemical reactions during the entire detection process. Therefore some teams have designed a new

type of electrochemical biosensing strategy that involves dissolving QDs into ions during the experiment and then performing related detections. For example, the QDs were dissolved by acid treatment [128
132] and enzymatic etching [133,134] to achieve detection.

11.4 GRAPHENE Graphene is a single layer of carbon atoms arranged in a hexagonal lattice [135]. Each carbon atom is connected to three other carbon atoms in addition to the σ bond, and the remaining π electrons form π bonds with the π electrons of other carbon atoms [136]. The electrons can move freely in this region, which brings excellent electrical conductivity. Graphene is a basic building block of fullerenes, carbon nanotubes, and graphite and can be regarded as an infinite aromatic molecule [137]. In addition to its excellent electrical properties, graphene also has unique and excellent mechanical, thermal, and optical properties [138] such as high strength [139], good thermal conductivity, and a large specific
surface area of 2630 m2/g [140]. Due to having oxygen-containing functional groups, graphene oxide (GO) is hydrophilic and can be dispersed in aqueous solutions [141]. Because oxygen-containing functional groups can bind to large surface areas, GO can be an ideal platform for covalently immobilizing biorecognition molecules [142]. Microsensors made of graphene can sense single atoms or molecules. When molecules attach or detach from the surface of graphene, the adsorbed molecules change the locality of graphene [143,144]. The carrier concentration causes a step change in resistance. This feature can be used to make various types of sensors and graphene-based composites have been widely used as electrode materials to improve functional electrochemical sensors and biosensors [145].

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11.4 GRAPHENE

11.4.1 Syntheses of Graphene There are several reported methods for the syntheses of graphene, including mechanical exfoliation of graphite and epitaxial growth on electrically insulating surfaces such as SiC, chemical vapor deposition (CVD), opening CNT, and the thermal expansion reduction of GO [146]. In 2004, Geim and Novoselov used mechanical exfoliation to repeatedly peel highly oriented pyrolytic graphite tapes [147]. The graphene-tagged tape was then sonicated in acetone and the graphene dispersed in acetone was removed using a silicon wafer. Large sheets of high-quality graphene can be produced, but the yield and the monolayer ratio of graphene in the product are extremely low and, thus it is not possible to use this method to produce graphene on a large scale. Selecting a solvent that matches the surface energy of graphene, such as 1-methyl-2pyrrolidone, N,N-dimethylformamide, dichlorobenzene, and others, as the medium, the mechanical force generated by the ultrasonic wave can be used to make the graphene separate from the graphite matrix, and then the separated graphene can be stably suspended in the solvent due to the interaction between the graphene and the solvent molecules [148
151]. The graphene obtained by this method is mostly monolayered or has few layers. Raman spectroscopy and XPS spectroscopy show that the prepared graphene is not oxidized and has fewer defects. However, this method also faces the problem of low yield. The concentration of graphene solution usually obtained is in the range of 0.01
0.03 mg/mL. Coleman et al. extended the sonication time to 460 hours in order to increase the yield of exfoliated graphene in the liquid phase. A graphene solution with a concentration as high as 1.2 mg/mL was obtained, but such long-time ultrasound was not suitable for practical applications [152]. For epitaxial growth, the silicon atom is removed from the surface of a single-crystal

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silicon carbide (SiC) wafer at a high temperature (1200 C
1500 C) to obtain epitaxially grown graphene is the first method that has been proven to prepare graphene thin films on a large scale [153]. With this method, a large area of single-layer graphene can be obtained with high quality. However, since singlecrystal SiC is expensive, growth conditions are harsh, and the grown graphene is difficult to transfer, the graphene prepared by this method is mainly used for the research of graphene devices using SiC as a substrate. Sutter et al. [154] used metal germanium as a substrate, used carbon atoms to infiltrate germanium at a high temperature of 1150 C, and then cooled to 850 C, and the carbon atoms floating on the crucible surface formed a monolayered carbon atom with a sheet structure and eventually grew a complete layer of graphite thin. Kim et al. [155] used carboncontaining gas to catalytically grow a large area of graphene on the surface. Ruoff et al. also obtained high-quality graphite refining films on the foil surface [156]. This growth of graphite by a metal-catalyzed process is a polycrystalline process and it is a new direction for the growth and transfer of graphitic refining. However, the graphite flakes produced by this method often have uneven thickness, a large number of layers, and high metal costs, making it difficult to produce on a large scale. In 2009, Kosynkin reported a method for the preparation of graphite stencils for MWCNTs [157]. First, the MWCNTs are oxidized and then cut in the radial direction. The carbon nanotubes are disassembled into single-layer or few-layered water-soluble graphite nanoribbons and their conductive ability is restored through chemical treatment. This method has a high yield and can obtain intrinsic graphite thin ribbons with excellent electrical properties, but the product has many structural defects and the number of layers is not controllable. Nakamura and Natori reported the

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method of inserting metal atoms between the tubes of MWCNTs and cutting the MWCNTs along the diameter of the tubes, and then using ammonia treatment and acid neutralization, followed by rapid annealing [158]. Graphite sheets and nanoribbons were obtained. Cut carbon nanotubes are a new method for preparing graphite stencils, but the homogeneity of this process and product is not controllable. Thermal expansion and reduction mean that the GO is rapidly heated to over 1000 C in a short period of time. The high temperature rapidly decomposes the O2-containing group in the GO and releases CO2 and other gases. The pressure generated when the gas is released enables effective separation of the GO sheets [159,160]. A significant effect of thermal expansion reduction is the destruction of the graphene structure due to the release of CO2. In the original process, the weight loss of GO was about 30%. Therefore voids and some structural defects remain on the reduced graphene base surface. However, despite these defects, the electrical conductivity of the reduced graphene can still reach 1000
2300 S/m, indicating that thermal reduction effectively restores the conjugated structure of graphene [160]. Currently biosensors are gaining much attention especially in the fields of healthcare, food, and environmental quality control. As a novel carbon nanomaterial with unique properties, graphene has been widely used as an electrode matrix material or a modification material in the field of biosensing. Scientists believe that graphene and its functionalized derivatized nanocomposites are excellent links between electrode surfaces and the redox centers of enzymes and proteins [161]. Applications of graphene and functionalized derivative materials in electrochemical biosensing include bioactive small molecule biosensors, enzymatic biosensors, DNA biosensors, immunosensors, and cell biosensors.

11.4.2 Electrochemical Bioactive Small Molecule Biosensing Based on Graphene Biologically active small molecule substances such as hydrogen peroxide, glucose, dopamine, ascorbic acid, uric acid, and folic acid, which are closely related to human health and play very important roles in the body. Functionalized graphene nanocomposites have fast electron transfer kinetics and excellent electrochemical catalytic activity. The constructed electrochemical sensors provide a good detection platform for the determination of these active substances. Komathi et al. prepared a new type of G/Ti(G)3DNS/CS nanocomposite and fabricated a DETbased electrochemical-PEC dual-mode biosensor for detecting cholesterol [162]. This material combines the excellent photocatalytic properties of TiO2 nanostructures, the biocompatibility and electron transport properties of CS, and the large surface area of G, thus promoting the DET between the active centers of ChOx and modified electrodes. It is well-known that effective separation of h1 -electron pairs is a challenge for the PEC-based biosensors. The application of this material solved this problem. Compared to PEC biosensors that did not contain G and/or ChOx, the G/Ti(G)-3DNS/CS/ChOx biosensor achieved the highest photocurrent response, which indicated that the PEC sensing was enhanced due to the charge separation at the interface between TiO2 and G and enhanced by the biocatalytic reaction involving ChOx and cholesterol. Monitoring the photocurrent at a low applied potential ( 6 0.05 V) enabled high selectivity of the PEC-sensing mode. This electrochemical-PEC biosensor can be considered as a model for the future design of other semiconductor novel multimode biosensors based on carbon nanostructure-based nanocomposites. Hou et al. used a graphene nanocomposite for electrochemical analysis of dopamine [163]. Studies have shown that the electrochemical biosensor has good electrochemical catalytic activity for dopamine with a detection limit of

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FIGURE 11.6

Mechanism of the graphene-based ECL dopamine biosensor. Reprinted with permission from S. Hou, M.L. Kasner, S. Su, K. Patel, R. Cuellari, J. Phys. Chem. C 114 (35) (2010) 14915
14921. Copyright 2018 American Chemical Society.

0.01 μM. The combination of carboxyl groups introduced on the edge of graphene by chemical modification process with Nafion not only provided a suitable environment for the oxidation of dopamine, but also made the oxidation behavior of dopamine exhibit excellent reversibility in cyclic voltammetry. The mechanism of the sensor is shown in Fig. 11.6. Lian et al. constructed electrochemical biosensors based on tryptophan-functionalized graphene nanocomposites and achieved simultaneous detection of ascorbic acid, dopamine, and uric acid with linear ranges of 0.2
12.9 mM, 0.5
110 μM, and 10
1000 μM, respectively, while the detection limits were 10. 09 μM, 0. 29 μM, and 1.24 μM [164], respectively. In previous works, Wang et al. investigated the enhanced performance of metal-graphene composites [165,166]. Copper NPs decorated with nitrogen-doped graphene (Cu
NG) was prepared by a facile thermal treatment, and further employed as a novel sensing material for fabricating a sensitive nonenzymatic glucose sensor. The Cu
NG nanocomposites showed enhanced electrocatalytic activity to glucose oxidation due to the integration of NG, which exhibited the oxidation peak current of glucose B23-fold higher than that of pure Cu NPs [167]. AuNRs were applied to modify the α-Fe2O3 nanocrystals anchored on the NG substances to form ternary α-Fe2O3-NG-AuNRs hybrids. The absorption of α-Fe2O3-NG-AuNRs

was successfully strengthened due to the SPR of AuNRs. After immobilizing the 17β-estradiol (E2) aptamer, this PEC active material can serve as an ideal sensing interface for PEC E2 detection [168].

11.4.3 Electrochemical Enzyme Biosensing Based on Graphene Enzymes are immobilized on the electrode surface by covalent and noncovalent interactions. The key step to successfully making a biosensor is how to effectively immobilize the enzyme on the electrode surface. Currently applied fixation strategies include adsorption, covalent binding, embedding, cross-linking, or affinity of complementary biomolecules. Zhang et al. developed a novel ECL biosensor for the determination of cholesterol through the direct immobilization of CeO2
graphene composites on a glassy carbon electrode (GCE), which exhibited outstanding reproducibility, long-term stability, and selectivity [169]. The enzyme sensor was used to detect glucose with a linear range from 12 μM to 7.2 mM and detection limit of 4.0 μM (signal/noise 5 3). The research results demonstrated that the use of CeO2
graphene not only could amplify the luminol ECL signal, but also promises excellent stability and a long lifetime of the ECL biosensor by providing a better

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biocompatible microenvironment for the immobilized enzyme. CeO2 nanocrystallines (CeO2 NCs) ensemble-on-nitrogen-doped graphene (CeO2
NG) nanocomposites were prepared through a one-step heat-treatment to improve the poor electron conductivity and the severe aggregation [170]. Compared with pristine CeO2 NCs, the nanocomposites can enhance the ECL intensity by 3.3-fold and decrease the onset ECL potential. A facile ECL method for cholesterol detection was developed with the CeO2
NG nanocomposites as the matrix to immobilize enzyme ChOx. Zeng et al. designed an innovative photoelectrochemical (PEC) biosensor platform based on the in situ generation of CdS QDs on GO using an enzymatic reaction [171]. The design principle of the sensor was that under the catalytic conditions of horseradish peroxidase, hydrogen peroxide reduces sodium thiosulfate to generate H2S and H, H2S reacts with

Cd21 to form CdS QDs, and then CdS QDs are excited by light. An increased photocurrent is generated as a readout signal. This strategy was successfully applied to PEC detection of CEA and showed a wide linear range from 2.5 ng/mL to 50 μg/mL with a detection limit of 0.72 ng/mL. A new direction has been provided for the development of numerous rapid and convenient analytical techniques using the PEC method. The fabrication process and detection principle of the biosensor is shown in Fig. 11.7. Similarly, Li et al. designed a subtle signal-photoelectrochemical sensing method for concanavalin A detection based on graphene supported TiO2 mesocrystal with an efficient enzyme-free labeling strategy [172]. In this paper, the research and development of NiCo2O4 labeling can effectively accept the photoinduced electrons from the conduction band of TiO2. Under optimized conditions, the photocurrent increases linearly with the increase of concanavalin A concentration from

FIGURE 11.7 Schematic illustration of the assembly process of the PEC Biosensor. Reprinted with permission from X. Zeng, W. Tu, J. Li, J. Bao, Z. Dai, ACS Appl. Mater. Interf. 6 (18) (2014) 16197. Copyright 2018 American Chemical Society.

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0.05 ng/mL to 0.5 μg/mL, further broadening the scope of application of excellent photosensitive materials photoelectric sensing platform. Ravenna et al. prepared a highly efficient glucose biosensor based on a GO film reduced with adsorbed phenothiazine to act on the electron transfer between flavin adenine dinucleotide (FAD)-dependent glucose dehydrogenase and the electrode [173].

11.4.4 Electrochemical DNA Biosensing Based on Graphene DNA sensors are constructed by immobilizing ssDNA molecules of a known sequence as a molecular recognition element on a GRmodified surface by physical adsorption or chemical binding. According to the DNA base pairing principle, the probe can recognize and bind the target DNA in the sample. Compared with enzyme electrodes, microbial sensors, and immunoelectrodes, DNA sensor research started relatively late and was not reported until the mid-1990s. Gupta et al. developed a highly sensitive method for detection of DNA hybridization which was based on the modification of GCE with AuNPs involving p-aminothiophenol (ATP) functionalized GO [174]. The film exhibited excellent properties for immobilizing DNA probes and sensing DNA hybridization. The linear detection range was from 1.0 3 10213 M to 1.0 3 1027 M and the detection limit was 1.10 3 10214 M (n 5 6). Qi et al. constructed a new electrochemical DNA biosensor by using a novel nanocomposite material prepared by Co3O4 nanorods (nano-Co3O4), graphene (GR), and chitosan [175]. The ssDNA probe was immobilized on the CTS
Co3O4
GR/CILE surface by electrostatic attraction, which could hybridize with the target ssDNA sequence under the selected conditions. The linear detection range was from 1.0 3 10212 to 1.0 3 1026 M and the detection limit was 4.3 3 10213 M. Lou et al.

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proposed a novel strategy for highly sensitive ECL detection of DNA based on site-specific cleavage of BamHI endonuclease combined with the excellent ECL activity of GQDs and bidentate chelation of the dithiocarbamate DNA (DTC-DNA) probe assembly [176]. The principle of this novel signal transduction ECL DNA biosensor was that the formed DTCDNA was directly attached to the gold surface through the S
Au
S bond, making the DNA more firmly immobilized on the gold surface. The BamHI cut the double-stranded symmetrical sequence, making the double-stranded DNA fragment shed from the electrode surface and inducing a decrease in the ECL signal. Congur et al. designed an EIS biosensor for microRNA-34a (miRNA-34a) detection [177]. GO was modified onto pencil graphite electrodes (PGEs) through passive adsorption. The hybridization between miRNA-34a target and its complementary DNA probe would increase the impedimetric.

11.4.5 Electrochemical Immunobiosensing Based on Graphene Immunosensors are biosensors that use immunological substances such as antigens or antibodies as molecular recognition elements. According to the determination of the need for markers, labeling and nonlabeling methods can be used. An obvious example of a NP-based electrochemical immunosensor is shown in Scheme. 11.2 [178]. For example, Pan et al. reported a novel dual-modality electrochemical biosensor that can simultaneously detect vascular endothelial growth factor (VEGF) and PSA in human serum with GO modified gold electrodes [179]. Dual-antibody modified poly-L-lactide nanoparticles (PLLA NP) could recognize the VEGF for signal amplification and further were used to capture and detect PSA.

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SCHEME 11.2 Mechanism of a nanoparticle-based immunochemical biosensor. Reprinted with permission from Farka, Z., Juˇrı´k, T., Kova´ˇr, D., Trnkova´, L.; Skla´dal, P., Chem. Rev. 2017;117(15):9973. Copyright 2018 American Chemical Society.

Yan et al. explored the fabrication of a novel, disposable, and highly sensitive electroanalytical immunosensor using graphene nanosheets (GS) and horseradish peroxidase (HRP)-labeled antibody functionalized with gold nanoparticles (HRP-Ab2/Au NPs) on a novel screen-printed electrode (SPE) [180]. Yang et al. fabricated a novel magnetic field controllable and disposable electrochemical immunosensor for rapid determination of clenbuterol (CLB) [181]. Graphene sheets -Nafion film was dropped on the screen-printed carbon electrode. Then Fe3O4-Au nanoparticles coated bovine serum albumin-CLB (BSA-CLB) conjugates were absorbed on the electrode with the aid of external magnetic field. The linear detection range was from 500 to 2 3 105 ng/mL and the detection limit was 220 pg/mL. Yang et al. used Au
Ag nanocomposite functionalized graphene as the sensing interface to detect tumor markers with GQDs coated porous PtPd nanochains as signal labels [182]. Due to the good conductivity and large surface area, porous PtPd nanochains can bind more GQDs and secondary antibodies to increase the sensitivity.

11.4.6 Electrochemical Cell Biosensing Based on Graphene Wu et al. reported a dual signal amplified electrochemical cell sensing [183]. Hep3B liver cancer cells were selectively recognized and captured by epithelial cell adhesion molecule antibodies immobilized on the interface of chitosan/reduced graphene oxides. Then the captured cells were labeled with SiO2 NPs modified with QDs. The electrochemical cell sensor has a linear detection range of 10
106 cells/mL and a detection limit of 10 cells/mL. Yan et al. immobilized MUC-1 aptamers on the porous GO/AuNP modified interface to selectively recognize and capture MCF-7 breast cancer cells, which were then labeled with thioneine and MUC-1 [184]. The detection range of this sensor for MCF-7 cells was 100
5.0 3 107 cells/mL and the detection limit was 38 cells/mL. Liu at al., developed a microfluidic paper-based analytical device and achieved the detection of CA153 antigens at the cell surface [185]. Au nanoflowers grew on the surface of paper fiber in the paper working electrode (PWE), which could greatly

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enhance the ECL intensity due to the good conductivity, biocompatibility, and large specific surface. Surface villous Au nanocage (SVAu nanocage) was applied to carry GQDs for higher ECL intensity due to its large loading amount and accelerated electronic transmissions.

11.5 GRAPHITIC CARBON NITRIDE (g-C3N4) BASED NANOMATERIALS Graphitic carbon nitride (g-C3N4) is a 2D metal-free polymeric semiconductor composed of C and N. Similar to graphene, the stacking layers of g-C3N4 were bound by van der Waals forces and the basic tectonic units are tri-striazine rings connected with planar amino groups [186]. In recent years, this material has attracted much attention due to its unique properties such as the facile synthesis process, good stability, nontoxicity, superior electronic band structure, and “earth-abundant” element composition [187]. The first report about carbon nitride (C3N4) can be traced back to 1834, when Berzelius and Liebig synthesized C3N4 polymers [188]. However, the wide research and utilization of g-C3N4 started around a decade ago as a promising photocatalysis for visible-light-driven water splitting due to the moderate band gap of 2.7 2 2.8 eV [189]. Compared to bulk compounds, exfoliated ultrathin 2D nanosheets exhibit extraordinary electronic properties besides the large specific
surface area [190]. The developed methods to obtain exfoliated g-C3N4 include the chemical vapor deposition, ultrasonication assisted liquid exfoliation, liquid phase, and surfactant-assisted exfoliation, the postthermal oxidation etching exfoliation, and others. [191
195] Because of the polymeric feature, the properties of g-C3N4 can be easily modulated by means of surface engineering. And various g-C3N4-based nanocomposites have been developed by the incorporation of noble

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metals, semiconductors, carbonaceous nanomaterials, and many other approaches to further improve their physical and chemical properties, such as synergistic effects enhanced catalytic activity and conductivity [196]. With these appealing features, the application of gC3N4-based nanomaterials in electrochemical sensing also has been well studied.

11.5.1 Amperometric Sensors Graphitic carbon nitride (g-C3N4) nanosheets can serve as a low-cost, green, and highly efficient electrocatalyst. With the good catalytic ability, g-C3N4 was applied in the direct detection of hydrogen peroxide, glucose, ascorbic acid, dopamine, uric acid, mercuric ions, nitrobenzene, and nicotinamide adenine dinucleotide [197
201]. For example, Zhang et al. prepared g-C3N4 through thermal polymerization and the exfoliation was achieved through ultrasonicating in water [202]. The obtained g-C3N4 nanosheets had strong affinity with mercuric ions because of amino groups. Compared with bulk g-C3N4, g-C3N4 nanosheets modified glass carbon electrode showed enhanced electrochemical performance. Under optimized experimental parameters, the detection limit for trace Hg21 detection was as low as 0.023 μg/L. Zhao et al. studied g-C3N4 as the electrochemical sensing material for the detection of tetrabromobisphenol-A (TBBPA), a kind of brominated flame retardant [203]. They investigated the electrochemical oxidation responses of TBBPA on g-C3N4-modified electrodes at different pH values and discovered the interaction between g-C3N4 and TBBPA was significantly affected by the pH value. As reported, g-C3N4 could combine H1 through the direct protonation and the medium-contacted surface was converted from negatively charged to positively charged, which could bring more active absorbing sites and make TBBPA concentrated

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on the modified electrode surface. The accumulated TBBPA effectively improved the electrochemical signal and this sensor showed good performance in environmental sample determinations. Besides direct detection, Zhou et al. reported an electrochemical immunosensor based on g-C3N4 [204]. Compared with bulk g-C3N4, mesoporous graphitic carbon nitride (mpg-C3N4) possesses large surface areas, tunable pore diameters, faster electron transfer rate, better biocompatibility, and more sensing sites. Thionine (Th), a widely applied electron transfer mediator, was bonded onto mpg-C3N4 through π
π stacking. The obtained Th-mpg-C3N4 nanocompound served as the electroactive probe and further carried secondary antibodies to build a sandwich-type structure for the sensitive detection of subgroup J of avian leukosis viruses (ALVs-J). To improve the detection performance, g-C3N4 was incorporated with many other nanomaterials. Gu et al. explored the combination of graphene and g-C3N4 to enhance the poor conductivity of g-C3N4 [205]. Cyanamide was loaded on the surface of GO and g-C3N4 was in suit synthesized during the reduction of GO to prepare G-g-C3N4. The unique electron structure effectively promoted the electron transfer ability and enhanced the electrocatalytic activities for redox reaction. The biosensing capacity of G-gC3N4 was tested with some electroactive biomolecules, such as rutin, uric acid, norepinephrine, tyrosine, tryptophan, and acetaminophen using the CV method. Another noble metal
free approach is the combination of metal oxide and g-C3N4 [206
208]. Tian synthesized g-C3N4/ ZnO nanosheets via a simple microwaveassisted hydrothermal treatment [206]. The exfoliated g-C3N4 nanosheets were beneficial for the growth of ZnO. The unique structure of g-C3N4/ ZnO nanosheets can improve the adsorption of H2O2 molecules on the electrode. So the fabricated H2O2 sensor had a good sensitivity with a detection limit of 0.05 mM. Sodium nitrite (NaNO2) is a common food additive and

preservative in the modern food industry; however, the excess of NaNO2also affects human health. Fe2O3 is a narrow n-type metal oxide semiconductor material and is regarded as an efficient electroactive material toward NaNO2. So an electrochemical nitrite sensor was fabricated based on the nanocomposites of Fe2O3, protonated carbon nitride (H
C3N4) and rGO [207]. G-C3N4 was treated with nitric acid to obtain H
C3N4 with smaller sizes, which was in favor of the uniform distribution on the surface of rGO. As active sites, H
C3N4 can effectively adsorb Fe31 and facilitate the in suit growth of Fe2O3. With synergistic effects among Fe2O3, H
C3N4, and rGO, the nanocomposites had significantly enhanced electrocatalytic ability toward NaNO2 and the sensor showed good sensitivity. Polymers also could be combined with g-C3N4. Cylindrical spongy shaped polypyrrole was doped with g-C3N4 using the chemical polymerization method. The as-prepared nanohybrid was utilized to adsorb negatively charged cholesterol oxidase (ChOx) through an electrostatic force because of the large specific
surface area and positive electricity. So a stable nanocomposite with high-enzyme loading was obtained for the construction of for cholesterol electrochemical biosensing [209]. Several works reported the incorporation of g-C3N4 with noble metal nanomaterials, such as gold, platinum, and palladium with different shapes and structures [210
212]. Tian et al. prepared well-aligned ZnO nanorods on the substrate of exfoliated g-C3N4 nanosheets through the microwave-assisted hydrothermal method [211]. Then Pt NPs were deposited to obtain Pt/ZnO/g-C3N4 nanostructures. This sensing interface exhibited remarkably enhanced gas response for ethanol and NO2 in electrochemical impedance spectroscopy, which could serve as a dual-functional gas sensor by changing the corresponding working temperature. Yuan et al. prepared Pd wormlike nanochains/graphitic carbon nitride (Pd WLNCs/g-C3N4) nanocomposites. The

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Pd WLNCs/g-C3N4-modified electrode was further functionalized with acetylcholinesterase (AChE) for organophosphorus pesticides (OPs) and huperzine-A (hupA) detection [212]. Ding et al. reported a sandwich-type immunoassay for PSA detection [213]. AuNPsdecorated g-C3N4 nanosheets (AuNP/g-C3N4) were synthesized by the wet-chemistry method and utilized for the labeling of polyclonal anti
PSA antibody and horseradish peroxidase. This complex catalyzes 4-choloro1-naphthol into an insoluble product and is coated on the electrode surface. So the increased impedance was proportional to the target PSA concentration. This electrochemical biosensor exhibited a good sensitivity and the detection limit was as low as 5.2 pg/mL.

11.5.2 Electrochemical Sensors The electrochemiluminescence of g-C3N4 was firstly reported by Xiao et al. The ECL measurements were conducted with potassium peroxydisulfate (K2S2O8) as the coreactant by cyclic voltammetry [214]. The possible cathodic ECL mechanism was proposed. Electrons were injected to the conduction band of g-C3N4 once the potential was negative enough. The electro-reduced g-C3N4 can readily react with a coreactant in a powerful oxidant state to produce an excited state. The subsequent decays from excited state to ground state emitted strong luminescence with a maximum peak around 470 nm. The cathodic ECL of g-C3N4 may be quenched by trace amount copper ions. Based on this mechanism, a sensitive and selective ECL sensor for Cu21 detection was fabricated with g-C3N4. They also studied the anodic ECL behavior of g-C3N4 with triethanolamine (TEA) as the coreactant [215]. The proposed possible anodic ECL response mechanism was similar to the anodic ECL pathways of other semiconductors. The positively charged g-C3N4 was produced from

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the electrooxidation. Electrooxidized TEA subsequently decomposes to generate the radical with powerful reductive properties, which reacted with the oxidized state of g-C3N4 to produce the excited state. And the anodic ECL of g-C3N4 was successfully applied in the determination of rutin. Liu et al. reported ECL from g-C3N4 nanosheets after the liquid exfoliation was 40-times stronger than bulk g-C3N4 with triethylamine (Et3N) as a coreactant [216]. They also discovered the oxidation product of dopamine would annihilate the Et3N radical and resulted in the decrease of ECL emission. Thus a sensitive ECL method for dopamine detection was developed with a subpicomole detection limit. To enhance ECL intensity, various g-C3N4 nanocomposites have been prepared, such as gold, silver, and graphene [217
223]. Chen et al. prepared and studied AuNPfunctionalized g-C3N4 NS nanohybrids (Au-gC3N4 NHs) [224]. AuNPs can trap and store the electrons from the conduction band of g-C3N4 besides preventing high energy electron-induced passivation. Compared with pristine g-C3N4, the nanohybrids exhibited strong and stable cathodic ECL emission and a novel ECL immunosensor was developed. The function of AuNPs can be replaced by other metals. For example, Fan et al. prepared an Ag-doped g-C3N4 to fabricate ECL biosensor for ultrasensitive detection of concanavalin A [225]. Wu et al. prepared amino-coated Fe3O4 nanoparticles decorated g-C3N4 [226]. With enhanced ECL performance, the developed ECL immunosensor exhibited a wide dynamic linear range for tumor marker carbohydrate antigen 125. Another popular approach is the combination of g-C3N4 and graphene [227,228]. With the excellent electron-transfer ability of graphene, the ECL intensity of g-C3N4 can be greatly enhanced. Li et al. prepared g-C3N4-graphene based on the electrostatic adsorption between positively charged poly(diallyldimethylammonium chloride) (PDDA) and negatively

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charged carboxylated g-C3N4 and graphene [216]. Carboxylated g-C3N4 exhibited higher water-dispersibility, which was beneficial to the stability of immunosensor. Carboxyl groups can be applied in the immobilization of antibodies through the covalently bond and g-C3N4 can be functionalized with antibodies through the covalently bond between carboxyl groups and amino groups. Xu et al. proposed a novel strategy utilizing C3N4 QDs to modify g-C3N4 nanosheets by simple oxidation of bulk C3N4 with H2O2 and UV light irradiation [229]. The ECL enhancement of obtained composites may be attributed to the rich surface defects that supply a greater number of hanging-bond atoms, which were more active to take part in the ECL process. Ferrocene (Fc)-labeled aptamers were further modified onto the nanocomposite via strong π
π interaction and constructed a signal-on ECL biosensor for platelet derived growth factor BB (PDGF-BB). Wang et al. discovered that the introduction of an O2 vacancy in Eu-doped g-C3N4 nanostructures could greatly enhance the ECL intensity (Fig. 11.8) [230]. The doping of Eu2O3 caused a remarkable positive shift of onset potential and signal amplification for cathodic ECL emission because the O2 vacancy can promote its catalytic activity. On this basis, a novel ECL aptasensor for 17β-estradiol (E2) was developed.

False positive or negative errors are inevitable when the detection was based on one signal change because of the instrumental or some environmental factors. To eliminate the possible interferences and make the detection more convincing, ECL ratiometric technology was developed [231]. Several works about ECL ratiometric sensing based on g-C3N4 were reported. Xu et al. developed a dual-signaling ECL ratiometric biosensor for HL-60 cancer cells detection with g-C3N4 and Ag
polyamidoamine (PAMAM)
luminol nanocomposites [232]. The ECL emission from g-C3N4 could be quenched by AgNPs due to the spectra overlap. The concentration of HL-60 could be quantified by both the quenching of cathodic ECL from g-C3N4 nanosheets and the enhancement of anodic ECL from luminol. Sensitive and reliable detection could be achieved through the ratio of ECL intensities at two excitation potentials. They further developed a dual-wavelength ratiometric ECL microRNA biosensor based on resonance energy transfer (RET) between g-C3N4 and Ru (bpy)321 [12]. Au-g-C3N4 NHs could exhibit ECL emissions with peak around 460 nm, which matched well with the absorption peak of Ru (bpy)321 and the emission at 620 nm was stimulated. So ECL-RET with high efficiency occurred that the ECL signals at 460 nm were quenched and signals at 620 nm were enhanced. Through FIGURE 11.8 Illustration of the fabrication and mechanism of the ECL aptasensor based on oxygen vacancy in Eu-doped g-C3N4 nanostructures. Reprinted with permission from X. Du, D. Jiang, L. Dai, W. Zhu, X. Yang, N. Hao, et al., Anal. Chem. 90 (5) (2018) 3615
3620. Copyright 2018 American Chemical Society.

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measuring the ratio of ECL intensities at two wavelengths, the concentration of miRNA-21 could be accurately quantified. Zhang et al. simultaneously applied the anodic and cathodic ECL reactions of g-C3N4 to fabricated potential-modulated dual-signal sensing of metal ions [233]. The selectivity of g-C3N4based sensors to target metal ions wasn’t good enough and may be interfered by other metal ions with much higher concentrations. So it’s difficult to distinguish the concentration of specific metal ion relying on a singleECL signal. But different metal ions have various energy-level matches toward g-C3N4 and cause different catalytic interactions in ECL reactions, which result in distinct quenching/ enhancement effects on ECL signals at different potentials. The accuracy and reliability of g-C3N4-based ECL sensors for metal ions may be greatly improved with modulating multiple-ECL signals strategy. The potentialresolved mode also can be used in the dual target detection. Guo et al. realized simultaneous detection of CA125 and SCCA based on the Ru-NH2@GO-COOH and AuNPs/g-C3N4 as the luminophore [234]. The two reactants K2S2O8 and DBAE did not interfere with each other.

11.5.3 PEC Sensors G-C3N4 has drawn plenty of scientific interest in PEC sensing [235]. Several intrinsic drawbacks, such as high recombination rate of charge carriers, low carrier mobility, and insufficient visible light absorption limit the applications of g-C3N4 in photo-to-electric conversion [236]. To overcome these problems, a hybrid with other nanostructures to form heterojunction nanocomposites is a promising approach for g-C3N4-based PEC sensors, including traditional semiconductors, carbonbased materials, QDs, metals, and organic dyes [237,238]. A typical method is the

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combination of other semiconductors with appropriate valence band (VB) and conduction band (CB) potentials [239
241]. TiO2, as a most popular photoactive semiconductor, has many advantages such as high resistance to photocorrosion, good physical and chemical stability, and low toxicity. But the wide band gap still limits its direct applications in PEC analysis. The combination of g-C3N4 and TiO2 can retard the photogenerated charge recombination of g-C3N4 p and improve the poor visible light excitation of TiO2, which significantly improves the photocatalytic activity [242,243]. For example, Kang et al. developed a simple double-channel PEC sensor based on g-C3N4/ TiO2 nanotube array hybrid film [244]. This method exhibited signal-enhancement response to ascorbic acid and showed a good performance in the determination the activity of alkaline phosphatase. Meanwhile, TiO2 can specifically bind with phosphate groups on substrate peptide. Li et al. synthesized TiO2/ g-C3N4 for immobilizing the phosphorylated peptide (P-peptide) by conjugating phosphate groups and providing PEC signals (Fig. 11.9) [11,245]. Then PAMAM dendrimer and alkaline phosphatase (ALP) can be captured on the modified ITO electrode via reactions between the
COOH groups of PAMAM dendrimer and the
NH2 groups of peptide and ALP for signal amplification. Based on the signal amplification and protein kinase A (PKA)-catalyzed phosphorylation reaction in solution, a novel PEC assay was developed for sensitive detection of PKA activity and the evaluation of PKA inhibition. CdS, widely applied QDs in PEC analysis, have broad visible light absorption and ideal band gap (around 2.4 eV). However, the selfoxidation of CdS may occur during the PEC process due to the photocorrosion. The formation of a CdS/g-C3N4 heterojunction could transfer the photogenerated holes of CdS to g-C3N4, resulting in an efficient separation and transfer of charges. This could effectively

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FIGURE 11.9 Enhanced PEC method for sensitive detection of protein kinase A activity using TiO2/g-C3N4, PAMAM dendrimer, and alkaline phosphatase. Reprinted with permission from X. Li, L. Zhu, Y. Zhou, H. Yin, S. Ai, Anal. Chem. 89 (4) (2017) 2369
2376. Copyright 2017 American Chemical Society.

improve the instability of CdS, limit the recombination rate of g-C3N4, and extend the visible light absorption of g-C3N4 [130,246,247]. Liu et al. prepared g-C3N4
CdS nanocomposites by directly mixing the aqueous solutions of g-C3N4 and CdS QDs with ultrasonic treatment [248]. Sensitized with CdS QDs, this composite exhibited dramatically enhanced photocurrent response. Tetracycline (TET)-binding aptamer was facilely anchored onto the electrode surface through π
π stacking interaction. TET was captured by the aptamer and oxidized by photogenerated holes, which inhibited the recombination of photogenerated electron
hole pairs. Photogenerated electrons would be more easily driven to the counter electrode by the bias potential and the photocurrent was enhanced. A sensitive “signal-on” PEC aptasensor for TET detection was successfully

developed. Liu et al. synthesized CdS@g-C3N4 p
n heterojunction and immobilized this high PEC efficiency material on two separated electrodes (WE1 and WE2) to build a spatialresolved ratiometric PEC platform [249]. WE1 was the working electrode while WE2 incubated with a fixed concentration of PSA served as the reference electrode. The target PSA detection was conducted through a sandwichtype PEC assay with Ab2
CuS conjugates as PEC signal quencher. The concentration of PSA was quantified by the ratio of photocurrents generated from WE1 and WE2. Dong et al. investigated the exciton
plasmon interactions (EPI) between CdS@g-C3N4 heterojunction and Au@Ag NPs [250]. Core
Shell CdS@g-C3N4 nanowires were prepared by immobilizing g-C3N4 on the surface of a CdS nanowire via the solvothermal approach, and

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the CdS@g-C3N4-modified electrode showed a great photocurrent due to the formation of the p
n heterojunction. With the presence of target microRNA-21 and duplex-specific nuclease (DSN), the hairpin probes on the electrode were circularly cleaved and affect the capture of Au@Ag NPs. Because of the overlap of the absorption spectrum of Au@Ag NPs with the emission spectrum of CdS@g-C3N4, the EPI can greatly quench the photocurrent. A highly sensitive PEC biosensor was successfully constructed. Other semiconductors with different nanostructures such as Zn0.1Cd0.9S nanocrystals and ZnO@CdTe core
shell nanocable were also applied in the fabrication of g-C3N4-based PEC biosensors [251,252]. Concerning the applications of metals, Yin et al. fabricated a “signal-on” PEC biosensor for microRNA detection based on nano-gC3N4-AuNPs composite [253]. Li et al. developed a novel PEC immunoassay for protein based on the enhancement of Cu21-TiO2 on gC3N4 nanosheets. G-C3N4 was firstly sensitized by a novel water-soluble cationic conjugated polymer denoted as PT-Cl through electrostatic adsorption. The obtained photoactive PT-Cl/g-C3N4 film was employed as the immobilization matrix for sandwich-type immunoassay. Copper(II) (Cu21) was doped on the surface of nanoporous TiO2 as labels, which could trigger synergy effects in the PEC platform for multi-signal amplification [254]. For carbon-based materials, the doping of GO can improve the dispersion, film-forming ability, and the PEC performance of g- C3N4 [255]. And the large specific
surface area and π-conjugated structure of GO/g-C3N4 was suitable for immobilizing the DNA aptamer. Lv et al. designed 0-dimensional/2-dimensional nanoheterostructures based on CQDs and g-C3N4 [256]. This nanocomposite was prepared by the solvothermal treatment to separated synthesized g- C3N4 and aminated CQDs. This combination was beneficial for the photoexcited electron
hole separation and

largely increased the photocurrents. After the immunoreaction, CuNCs as the tracer was dissolved into copper ions and captured through the coordination of amino groups and copper ions. The photocurrent was quenched by copper ions because of the photoinduced electrons would be transferred from the CB of the nanoheterostructures to copper ion.

11.6 CONCLUSION In summary, this chapter presents recent advances of micro- and nanotechnologies for electrochemical biosensor development. The introduction of these technologies can significantly improve sensing performance and has exceptional potential in the design of electrochemical biosensors. The large specific
surface area of nanomaterials is beneficial to load more biological recognition probes on the electrode. The good conductivity and catalytic ability can effectively amplify the signal. Microfluidics chips have the advantages of portability, reduced reagent consumption, and less analysis time due to their small size and high integration. Great efforts have been made on the preparation, functionalization, and application of micro- and nanostructures in the field of biosensing. We believe that the unique features of micro- and nanotechnologies are very suitable for low-cost and portable devices. But these works are mainly studied in labs and few products have entered the market. One of the final goals of analytical chemistry is to improve the daily testing technology, which means to develop simple, fast, lowcost, and sensitive detection methods. Therefore besides fundamental research, future works should focus on the miniaturization and large-scale manufacturing of these novel electrochemical biosensors to obtain commercially available devices for POCT testing, food safety, and environment monitoring.

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C H A P T E R

12 Cholesterol-Based Enzymatic and Nonenzymatic Sensors Rajasekhar Chokkareddy1, Niranjan Thondavada1, Surendra Thakur2 and Suvardhan Kanchi1,* 1

Department of Chemistry, Durban University of Technology, Durban, South Africa, 2 eSkills CoLab, Durban University of Technology, Durban, South Africa

12.1 INTRODUCTION Current improvements in the field of biochemistry, physics, chemistry, and molecular biology initiated the use of biosensors, which now have a huge range of biological recognition elements with enhanced analyses of sensitivity and selectively. Low-cost biosensors have assisted the developments in the engineering field. Furthermore, improvements within the fields of microelectronics and fiber optics have provided more durable and smaller signal transducers, which, in turn, provide better signal-to-noise ratios and less processing costs [1]. Biosensor is a combined electronic device that utilizes biological substances including enzymes, cells, antibodies, nucleic acids, receptor proteins, or tissue sections used as parring of biological component and analyte sensor toward the transducer for

signal recognition that is totally exhibited on a section after amplification and processing (Fig. 12.1). The selectivity possessed by biosensors in biological particle and the dispensation power of optoelectronics and modern microelectronics offers improved analytical techniques with many uses in the field of medical diagnostics. This mixture allows one to determine the analyte of interest without using components. Hence, physiochemical modification formed by definite exchanges among the analyte of interest as well as a biorecognition component is identified, which is determined by a transducer. In this context, the biochemical signal is translated by the transducer into an electronic signal, which is converted into a digital format or analogue. The absorption of the targeted analyte of interest increases as the signal increases, permitting both qualitative and

*Corresponding author.

Advanced Biosensors for Health Care Applications DOI: https://doi.org/10.1016/B978-0-12-815743-5.00012-3

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FIGURE 12.1 Diagram representing a biosensor.

quantitative measurements. The fast and precise detection of cholesterol has stimulated much interest due to its close related analysis of hypertension, coronary heart diseases, arteriosclerosis, myocardial infarction, and lipid absorption dysfunction. In spite of the research determinations, a description of its purpose has been limited [2]. Cholesterol expressively contributes in a number of significant biochemical, physiological, and cellular procedures, including the immune response [3]. In cellular membranes, cholesterol is essential as it assists as a required site for significant protein aspects including medium metalloprotease7. Moreover, this complex is utilized as a next messenger, that is, as an activator of the protein kinase C and a controller of combinations of steroid hormones [4]. It acts as a chief mechanical component of a plasma membrane and is an originator of natural components, such as vitamin D, steroid hormones, and bile acid [5,6]. For consistent approximation, biosensor advancements that permit suitable and quick detection are under intensive investigation and improvements. Cholesterol oxidize, which is mechanically significant for use in bioconversions, is suitable for the clinical examination of cholesterol. Cholesterol originates from endogenous biosynthesis and diet. The structure of cholesterol

consists of four member rings (A, B, C, D) and two of the angular methyl groups are located at C13 and C10. These structures are characteristic properties of all-natural steroids. A 3β-OH group found in cholesterol is located at ring A, with a double bond at C6 and C5, and side chain of isooctyl at C17 [2]. Fig. 12.2 displays some of cholesterol’s structures. The examination of enzymatic of cholesterol utilizing cholesterol oxidise is the topic of interest in this chapter. Brevibacterium sterolicum (BCO) and Sterptomyces hygroscopicus (SCO) are the best mutual sources of cholesterol oxidise. The determination of BCO and SCO for cholesterol by the enzymatic kinetic technique was reported by Srisawasdi et al. According to the obtained results, it was found that SCO possess more sensitivity compared to BCO. Moreover, the linearity for BCO and SCO was 2.6 mmol/L and 20.7 mmol/L, respectively. It was also described that the SCO is specifically new and precise, and does not show any interference from hemoglobin as high as 7.5 g/L [7]. Biosensors also have the potential to frequently determine the absence, presence, or concentration of definite inorganic or organic materials in the samples of interest, and the ability to observe cholesterol in blood is motivated by the expansion of effective cholesterol-based biosensors.

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FIGURE 12.2 Structures of cholesterol.

12.2 CHOLESTEROL ABSORPTION Accurately 0.35 g/day of cholesterol consumption according to nutritional sources in Western societies and has been reported that it depends on food habits. In addition, from a usual Western diet, approximately 55% of expended cholesterol results from fish and meat, although dairy products (20%) and egg consumption (25%) also significantly contribute toward cholesterol intake. It was also estimated that a proper dietary routine can lower the density of lipoprotein cholesterol (LDL-c) up to 25%
30% [8,9]. According to existing estimations, about 50% of dietary cholesterol is absorbed. A number of procedures have been engaged in determining cholesterol absorption [10,11]. Moreover, these measurement provides small absorption as anticipated, and then later can be determined by abdominal perfusion procedures in which

evaporation of a identified content of cholesterol over a section of the intestine can be measured [12]. Therefore, the intake of cholesterol increase by 3% then, it shows a high risk of atherosclerotic cardiovascular disease (CVD) and leads to raise in LDL-c levels. In this context, dietary ways and problems regarding cardiovascular actions is considered. 50% of the cholesterol that is expended arrives in the body, while small amount of cholesterol absorption differs significantly among individuals (20%
80%) [13]. About 75% of the cholesterol absorbed through the intestines is derivative from the endogenous process which is then discharged toward the intestines with bile. Moreover, in humans, about 1 g/day secretion of biliary cholesterol is observed, which is far less than the quantity of cholesterol directly received by the intestines through the diet. This implies that cholesterol production inhibitors are required

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to possibly decrease the cholesterol content being absorbed through the intestines in a significant amount than that was achieved by controlling consumption of cholesterol. In addition, solubilization in the proximal abdominal lumen is the first step involved in cholesterol absorption due to its hydrophobic nature (cholesterol is insoluble in water). In other words, to allow for absorption, hydrophobic nutrients and cholesterol first need to be solubilized [14]. Therefore, the liver creates bile as a mediator. During digestion of a meal, the gallbladder receives the bile which is then ejected into the mutual bile duct and drifts through the intestines where it is received by the sphincter of Oddi. Therefore phospholipid concentration, cholesterol, and bile acids are contained in the bile which are in millimolar in the form of mixed micelles. These then act as an emulsifier and solubilize cholesterol and some of the lipids. In the inexistence of bile acids, as is often the case for cholesteric people and in situations where absolute bile deviation is done, absorption of cholesterol is normally not present. Thus, reducing fat consumption results in an improved percentage of cholesterol, that results in atherosclerosis expansion [15].

12.2.1 Cholesterol Oxidase 3β-Hydroxysteroid: oxygen oxidoreductase, which is also called cholesterol oxidase (ChOx). The isomerization of cholesterol to cholest-4-en-3-one is catalyzed by this flavoenzyme as well as oxidation, and this has been characterized chemically and structurally. Moreover, the enzyme is characterized as extracellular and arises in a secreted or cell surface-associated form depending on the fabricator growth and microorganism situation. ChOx is normally utilized as the

biosensing material during the modification of cholesterol biosensors. ChOx is a flavinadenine-dinucleotide (FAD) consisting of an element known as flavoenzyme. In the cholesterol system, dehydrogenation of C(3)-OH is catalyzed by this element [16]. A member of alcohol dehydrogenise/oxidase is then released in the methanol
glucose
choline oxido-reductase group. The steroid rings containing trans double bond Δ5
Δ6 which are contained in the 3β-hydroxysteroids are catalyzed by FAD in the form of isomerization and oxidation, the result in function Δ43-ketosteroid as well as H2O2. The process occurs after the consumption of ChOx. From alcohols, the major receiver of hydride is the oxidized FAD, and then the final acceptor is the reduced FAD where it conveys the redox equivalents to dioxygen (Fig. 12.3). The use of covalently linked flavin which stabilizes ChOx was reported by Caldinelli et al. In their work it was observed that when flavin is linked with protein moieties in ChOx it can be considered as a structural device that enhances stability in all-inclusive of the protein tertiary structure. At low concentrations of urea and temperature, in the nonexistence of a covalent link, the denaturation procedure is less cooperative [17]. In addition, the studies on cholesterol lipophilic nature was performed by the catalysis studies involving kinetic mechanism resulted in the formation of fatty acid. Therefore, the solubility of cholesterol decreases in aqueous medium and the same mechanism is used in the preparation of detergents [18]. The enzymatic characteristics of ChOx formed from brevibacteriu sp. are described by Salva et al. It was observed that ChOx can be easily removed by the detergent and its activity was based on the absorption of the Triton X-100 (detergent). At 0% concentration of detergent, it was found that the enzyme loses its motion completely [19].

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FIGURE 12.3 Schematic illustration of cholesterol oxidase enzyme reaction.

12.2.2 Cholesterol Esterase The pancreatic cholesterol esterase are the enzyme present in the pancreas that hydrolyzes cholesterol esters to free fatty acids and un-esterified cholesterol. Over a period of 30 years, extensive research has been carried out on a protein with similar properties that can be separated from homogeneous body fluids as well as tissues. One of the advantages is the nonspecificity of the enzyme and its ability to hydrolyze cholesteryl esters, triacylglycerol, vitamin esters, lysophospholipids, and phospholipids. During the first attempts of these determinations, it was not easy to identify that these many activities caused by enzymes were from a part of the similar or same proteins. In addition, theses enzyme were characterized as nonspecific lipase, phospholipase AI, lysophospholipase, bile-salt-dependent lipase, bile-saltstimulated lipase, carboxyl ester hydrolase, and carboxyl ester lipase [20]. Cholesterol esterase (ChEt) was calculated in order to

examine the correlation between the enzyme activity and absorption of dietary cholesterol. In addition, high concentrations of ChEt is present in breast milk and this helps in providing the enzymes to infants whose pancreases are not yet fully advanced [21]. ChEt plays a huge role in intestinal lipid absorption in the upper intestinal area as well as behaving as a cholesterol transfer protein. This directly occurs in the lipoprotein metabolism and enhances the levels of cholesterol in serum. In the group of proteins namely α/β-hydrolase fold, where ChEt is originated; and catalytic mechanism takes place during the hydrolysis of lipid substrates in the form of serine proteases. A functional site of serine residue is observed, and aspartic acid with a histidine produce a catalytic triad [22]. Pancreatic cholesterol esterase structures at 1.6 A resolution have been reported by Chen et al. In α/β-hydrolase fold, it was found that there are rare active sites contained within the catalytic

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triad. Hence, in the active part it is easy to notice when protein in the hydrophobic C terminus is blocked. When this occurs, it diverts the oxyanion hole far from the catalytic Ser194 and useful binding site. In cholesterol esterase structure, the amphipathic helical lid observed in triglyceride lipases is reduced, but this feature is not important for stimulating this lipase (Fig. 12.4). Therefore the new structural features proposed are in alignment with a bile-salt-dependent motion of the enzyme containing activation mode of lipase. A new activity for aliphatic chains contained within fatty acids was proposed which can be imitated by detergent in the structure, and this displayed an important role for the C terminus in the activation of the lipase [22].

12.3 ENZYMATIC SENSORS Enzymatic or nonenzymatic analytical techniques are used to classify cholesterol detection. Normally, higher selectivity is achieved with enzymatic techniques compared to the nonenzymatic techniques and this has been utilized in some of the routine clinical laboratory. The mechanisms were represented the determination of cholesterol by the enzymatic process. ChE

Cholesterol ester 1 water
! cholesterol 1 fatty acids ChOx

Cholesterol ester 1 oxygen
! 4-cholestene-3-one 1 H2 O2

FIGURE 12.4 Schematic illustration of the cholesterol assay principle.

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12.3 ENZYMATIC SENSORS

Hydrolysis of cholesterol ester by the use of cholesterol esterase (ChE) to obtain fatty acids and free cholesterol is shown in the first mechanism. In the second mechanism, oxygen is used to oxidize free cholesterol oxygen by utilizing cholesterol oxidase (ChOx) to give hydrogen peroxide (H2O2) and 4-cholestene-3one. The quantity of H2O2 is mostly utilized for the indirect process for quantifying cholesterol. A significant factor in the improvement of biosensors based on enzymes is the steadiness of the enzymes. In addition, preferred solid supports are used to achieve stabilized and strong enzymes with high activity for a long period of time. This procedure is known as immobilization. This method is mainly utilized in attaching some biological substances on a transducer surface. In biosensors, the performance of all the components is important because it can result in the proper functionality of a biosensor which also depends on how the enzymes bond with the desired biosensor surface and continues with being active during the process. A T-cut quartz crystal sensor has been developed by Martin et al., in order to determine the real-time analysis of cholesterol concentration in serum and buffer by utilizing

321

ChOx, trienzyme structure of cholesterol esterase (ChE), and horseradish peroxidase (HRP). The H2O2 resulting from the reaction involving ChE
ChOx oxidizes diaminobenzidine (DAB) in the existence of HRP. An optimal response of the sensor to cholesterol was observed due to the existence of Triton X-100 with 0.1% (v/v) at 1 U/mL ChE and 0.2 U/mL ChOx. Moreover, the advantages of the quartz crystal method produces good quality results with less interference when compared to optical biosensing and conventional electrochemical strategies [23]. Alagappan et al. has been developed an electrochemical biosensor for cholesterol using ChOx enzyme which was immobilized with gold nanoparticles and functionalized MWCNTs/PPy nanocomposite. A two-step method was used to modify this sensor. First, the wet chemical process was utilized to construct Au NPs-f-MWCNT, and the second entails electropolymerization of pyrrole. A support matrix is utilized as PPy so as to hold ChOx and the existence of Au-fMWCNT results in an increase in electrical conductivity. The electrode fabrication is displayed in Fig. 12.5. The linear response due to the use of AuNPs-f-MWCNT-PPy-ChOx/GCE was between 2 3 1023 and 8 3 1023 M in

FIGURE 12.5 A schematic diagram of the fabrication of the Au-f-MWCNT-PPy-ChOx-GCE-based biosensor.

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amperometry with a detection limit and sensitivity of 0.1 3 1023 M and 10.12 μA/mM cm2, respectively [24]. Ahmadraji et al. reported a reproduced electrochemical sensors to quantify the amount of cholesterol in serum samples using amperometric method and the mechanism is illustrated in Fig. 12.6. An evaluation of surfactant Triton X-100 based silver paste screen-printed electrode with H2O2 for the improved detection of cholesterol from lipoprotein. Triton X100 with 0.5% (v/v) is employed in biosensors as a phosphate buffer solution with 156 U/ mL cholesterol esterase as well as 60 U/mL cholesterol oxidase. In addition, the total cholesterol measurement contained in serum is within the range of 0
10 mM and a good sensitivity of 2.24 3 1028 A/mM was observed, with a detection coefficient of 0.984, limit of detection of 2 mM, and the standard deviation was 10.8% (n 5 3) [25]. A mediator-free cholesterol and reusable biosensor was developed by Rahman et al. This was constructed using ChOx which was then modified by utilizing self-assembled monolayer (SAM) of thioglycolic acid (TGA) Lipoprotein Apo C

(covalent enzyme immobilization by dropping process) on biochips. A fabricated biochip, Au/TGA/ChOx, was then utilized in order to determine the cholesterol by means of the cyclic voltammetric method for Au/TGA/ ChOx at room temperature and pressure. The Au/TGA/ChOx fabricated biochip sensor displayed a good linearity within a linear range of 1.0
1.0 mM, with R2 5 0.9935, detection of limit was 0.42 nM, and it exhibited high sensitivity B74.3 μA/μM cm2. The fabrication scheme is provided in Fig. 12.7 [26]. Sekretaryova et al. reported a reagent with less cholesterol content based on human plasma utilizing a novel single-enzyme, selfpowered, membrane-free biosensor, where both anodic and cathodic bioelectrocatalytic reactions are motorized with the use of the same substrate. In this process, a sol
gel matrix was used to immobilize the cholesterol oxidase in the anode and cathode. In addition, this novel method showed an increase in the sensitivity of 26.0 mA/M cm2 as well as a good dynamic linear range of 4.1 mM. The schematic illustration of a single enzyme, membrane-free, self-powered, cholesterol FIGURE 12.6 A general method of amperometric determination for cholesterol. All lipoproteins are solubilized by the use of Triton X100 reagent, which allows access to cholesterol esters by cholesterol oxidase (ChOx) and cholesterol esterase (ChEs). The generated H2O2 can then be determined by utilizing the catalytic electrodes at
0.1 V against Ag/AgCl.

Solublised lipoprotein Apo B-100

Triton X-100 Unesterified cholesterol Phospholipid

Esterified cholesterol

Trigly ceride Inkjet printed Triton X-100 PBS H2O2 electrocatalyst

O2

ChOx red

Esterified cholesterol

ChEs

Free cholesterol H2O2

ChOx oxi

Silver paste screen printed electrode

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323

FIGURE 12.7 Schematic diagram representing the fabrication technique for a cholesterol biosensor utilizing tiny biochips.

FIGURE 12.8 Schematic illustration of single enzyme
based, self-powered, membrane-free, cholesterol biosensor.

biosensor is shown in Fig. 12.8 [27]. A novel cholesterol sensor where the CNT electrodes were aligned vertically with more than one process of electrochemical polymerized and immobilized enzymes was developed by Wisitsoraat et al. In this process, the vertically aligned CNTs are involved which are then selectively grown on a 1 mm2 window of Aucovered Si/SiO2 substrate by thermal chemical vapor deposition (CVD) with water-assisted etching and gravity impacts. The CNTs were then functionalized separately and enzymes

were immobilized by the electrochemical polymerization of cholesterol enzymes and polyaniline. The linear relation between the current response and the concentration of cholesterol can then be seen within the range of concentration between 50 and 300 mg/dL with a good sensitivity of 0.22 μA/(mg/dL) [28]. A highly sensitive and versatile biosensor platform was established by Sharma et al. Electrochemical-enzymatic redox cycling is involved in this process where the selective enzyme is induced and immobilized on nanosized carbon interdigitated electrodes which are fabricated with the use of AuNPs. The existing AuNPs
carbon interdigitated electrodes based cholesterol biosensor displayed a broad sensing range between 0.005 and 10 mM, with a high sensitivity of B993.91 μA/ mM cm2, and limit of detection of B1.28 μM. Fig. 12.9 shows the biosensor configuration. At the enzyme surfaces, the steps involved in the reaction mechanisms are: (1) the substrate is oxidized to the product; and (2) ferricyanide ([Fe(CN)6]32) is reduced to ferrocyanide ([Fe (CN)6]42). Therefore ferrocyanide is produced through enzymatic reaction and later becomes oxidized at both combs (1 and 2) of the AuNP/carbon interdigitated electrodes [29]. In cholesterol determination, a distinctive ChOx

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A diagram representing the sensing procedure based on the redox cycling of [Fe(CN)6]32/[Fe(CN)6]42 on the carbon electrode surfaces and enzymes at nanoscale gold nanoparticle (AuNP) which are the functionalized selectively with ChOx.

FIGURE 12.9

liquid crystal biosensor based on the disruption of orientation in liquid crystal was reported by Tyagi et al. In this method, a self-assembled monolayer of (3-aminopropyl) trimethoxy-silane and dimethyloctadecyl [3-(trimethoxysilyl) propyl] ammonium chloride was constructed utilizing a glass plate by simple adsorption. For a period of 12 hours, ChOx is immobilized on the self-assembled monolayer surface before using the film for biosensing applications. In addition, a liquid crystal-based biosensing study showed a self-assembled monolayer/ChOx/liquid crystals (5CB) cells for cholesterol concentrations ranging between 10 mg/dL and 250 mg/dL. The main drawback of this device was the shelf life which is very low (10
15 days) [30]. Yunianto et al. developed a fiber optic biosensor in order to examine cholesterol in serum and blood samples. Furthermore, this fiber optic sensor is engineered with a special sensing part where the fiber optic is grated by 5 scratches followed by bending treatment of 5 cm. UV-Vis spectrometer was the first test to

be conducted for this fiber and it gave good linearity of 0.96. Then a second test was conducted by the use of light spectrometer with 0.94 of linearity in white LED. Moreover, these optical fiber sensors were designed to work well in the range of serum or blood concentrations between 140 and 250 mg/dL [31]. Dey et al. described an improved and sensitive amperometric biosensor involving hybrid elements derived from graphene and nanoscale Pt particles (nPt) for proper detecting of cholesterol and H2O2. The fabricated electrode with PtNP-based hybrid element was successfully catalyzed by electrochemical oxidation of peroxide at the potential range of 0.4 V, which is greater than 100 mV less positive compared to the bare Pt electrode. Furthermore, increased sensitivity is observed in the sensing platform, which results in a linear response concerning H2O2 of 12 mM with the limit of detection 0.5 nM in the in-existence of enzyme or redox mediator. A schematic illustration of the biosensor is showed in Fig. 12.10 [32].

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FIGURE 12.10 Schematic diagram showing the biosensing of the cholesterol ester based on the GNS-nPt-based biosensor.

FIGURE 12.11 An electrocatalytic reaction of cholesterol utilizing ChOx/MWCNT transformed electrodes.

A cholesterol biosensor consisting of ChOx and MWCNTs was reported by Yang et al. A glassy carbon electrode (GCE) was successfully used to prepare this sensor. Good performance was shown by this sensor with a linear range of 4.68 3 1025
2.79 3 1024 M, good sensitivity of 1261.74 μA/mM cm2, and limit of detection of 4.68 3 1025 M. The film made on the electrode surface can be seen in Fig. 12.11. A carbon dot/hemoglobin (CD/Hb) composite utilized as a bioreceptor in an optical cholesterol biosensor was described by Bui et al. These visual sensors determine cholesterol via fluorescence improvement of CD, which is usually reduced through the π
π interactions between Hb and CD in the CD/Hb complex. Moreover, the CD/Hb complex allows the cholesterol to be selectively detected in the linear range of 0
800 μM, with good detection limit of 56 μM, and a response time of

# 5 minutes in human blood plasma samples. The fluorescence quenching of CD with Hb and improvement of cholesterol is displayed in Fig. 12.12 [33]. Pakapongpan et al. reported a cholesterol biosensor involving direct electron of ChOx which was covalently functionalized on MWNTs and fabricated SPEs. In addition, this CNT can enhance the immobilization of enzymes as well as develop the direct electrochemistry of ChOx, which has good functional uses related to cholesterol biosensor [34]. Aydogdu Tig et al. developed an amperometric based cholesterol biosensor by immobilizing ChOx/SnO2NPs/nafion on carbon paste enzyme electrode. In addition, the detection of cholesterol was performed by electrochemical oxidation of H2O2 at 0.6 V versus Ag/AgCl. The CPE/SnO2NPs-ChOx/Naf displayed a linear range between 0.20 and 4.95 μmol/L and low limit of detection of 0.04 μmol/L [35].

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FIGURE 12.12 Schematic representation of fluorescence quenching of carbon dots with hemoglobin to improve the detection of cholesterol. FIGURE 12.13

Proposed biochemical reaction on the nanocomposite and bioelectrode.

Cholesterol oleate

ChEt ChOx (Oxi)

[ Fe(CN ) ] 3–

E

6

Fatty acid + Cholesterol

L

MWCNT-PPY nanocomposite

E

e–

C T R

Cholest-4-ene-3-one

[Fe(CN ) 6 ] 4–

ChOx (Red)

Sing et al. developed a nanocomposite electrode with carboxy functionalized MWCNT and PPy which was fabricated onto an indium
tin
oxide (ITO) electrode for the detection of p-toluene sulfonic acid (PTS). This was utilized for the examination of cholesterol. The PPY
MWCNT/ITO nanocomposite electrode was used to immobilize ChEt and ChOx utilizing N-hydroxy succinimide and N-ethyl-N- (3-dimethylaminopropyl) carbodiimide, and for evaluation of esterified cholesterol. This ChEt
ChOx/ PPY
MWCNT/PTS/ITO nanocomposite has a linearity between 4 3 1024 and 6.5 3 1023 M/L, with a response time of approximately 9 s for cholesterol oleate concentration and thermal stability of up to 45  C, and Km of 0.02 mM [36]. The proposed

O D E

biochemical reaction on the nanocomposite and bioelectrode is shown in Fig. 12.13. Fig. 12.14 displays various enzymatic schemes developed by many research groups on electrochemical and optical biosensors to determine cholesterol as well as free cholesterol by immobilizing PANI on to the electrode surface. This was the first attempt in utilizing PANI for cholesterol biosensors on electrochemically polymerized PANI films. Electrochemically deposited PANI film was developed by Singh et al. by ChOx, HRP enzymes, and ChEt, which was examined at different pathways, pathways 4 and 5 shows that the total cholesterol is detectable as shown in Fig. 12.14. The polyaniline/cholesterol oxidase/cholesterol esterase films gave a good limit of detection and sensitivity of

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327

FIGURE 12.14 Different biochemical routes adopted for developing electrochemical and optical electrochemical biosensors for cholesterol determination by utilizing PANI as the enzyme immobilization platform.

about 25 mg/dL and 0.042 μA mg/dL, respectively [37]. The first proposal for the application of electrophoretically nanostructured PANI films related to cholesterol biosensor improvements was done by Dhand et al. Route 1 (see Fig. 12.14) included a o-dianisidine strategy for examining fabricated bioelectrodes in terms of photometric response. A low value was exhibited due to an increase in cholesterol interactions with oxidase
cholesterol on PANI matrices which are fabricated electrophoretically. The ability and productivity of these matrixes fabricated electrophoretically are convincing, and they have been used to examine the possible usage of electrophoretically dropped PANI nanotubes film for cholesterol determination [38]. For biosensor development for cholesterol

determination, the films involving Langmuir
Blodgett (LB) PANI-stearic acid (SA) were evaluated by Matharu et al. for possible biosensor applications. They observed that bioelectrodes based on ChOx/Glu/PANI-SA LB film/ITO showed high affinity toward cholesterol due to the sequence preparation of the PANI molecules that enables steady sharing of ChOx in the desired shape or structure [39]. A cholesterol biosensor with mediator support by electropolymerized PANI-MWCNT matrix containing potassium ferricyanide (K3 [Fe (CN)6]32) was used to immobilize with ChOx by Nguyen et al. A good linearity of cholesterol concentration ranging from 0.02
1.2 mM was displayed by the bioelectrode, with a correlation coefficient of 0.9985 [40]. A silica-PANI bienzyme cholesterol

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FIGURE 12.15 (A) Schematic representation of electrospraying technique. (B) The enzymatic response between ChOx and cholesterol on the GR/PVP/PANI modified paper-based biosensor.

biosensor was described by Manesh et al. This system was achieved by electrochemical frame of HRP and ChOx during the polymerization of poly (N [3-(trimethoxysilyl)propyl]aniline) (PTMSPA). At low values of electrochemical potential, that is, 2150 mV. This type of sensor has good working ability due to the fabricated PANI matrix that enables fast electron transfer among the redox center of HRP and the electrode. Furthermore, the good analytical performance showed by PTMSPA-HRP/ChOx-ME gave a range of 1
25 mM in detection limit, with high selectivity and sensitivity [41]. Ruecha et al. developed a nanocomposite made of GR, PANI, and PVP using a highpotential assisted electrospraying method (Fig. 12.15) for the preparation of a cholesterol biosensor using paper. GR/PVP/PANI nanostructures with 160 6 1.02 nm were used to enhance the developed electrode, which was then utilized for immobilization of ChOx, and the resultant response given by the biosensor was expected. Good electrocatalytic activity of the bioelectrodes toward oxidation of H2O2 was observed, resulting in improved electrodes in terms of conductivity and better sensitivity [42]. Shin et al. described an electrospun PANI/polystyrene blended fiber by layer-by-layer using electrostatic adsorption of ChOx for the amperometric cholesterol biosensor [43]. Rodrigues et al. described a

stable and highly sensitive amperometric biosensor which was developed by using cholesterol oxidase, chitosan, and MWCNTs to examine the cholesterol amount in serum samples. ChOx was covalently immobilized on MWCNTs and this enzyme was coated over a GCE and then followed by a layer of chitosan deposited to prevent enzyme leaching (Fig. 12.16). In addition, the optimized biosensor possessed a minimum limit of detection of 0.13 μM (S/N 5 3), with a quick reaction time (B5 seconds), broad linear range from 0.16 to 9.69 mM (25
500 mg/dL) (N 5 10, R2 . 0.99), and 92% stability for up to 7 months [44]. Lin et al. described a novel fabricated electrode by the usage of a nanocomposite mixture containing of AuNPs and molybdenum disulfide. Its application as a cholesterol biosensor was also examined. An amperometric i
t technique was utilized for determining the amount of cholesterol in a sample, and a linear calibration graph for cholesterol was observed within the range of 0.5
48 μM with a limit of detection of 0.26 6 0.015 μM. The observed values for apparent app Michaelis
Menten constant (KM ) and sensitivity were B0.325 mM and B4460 μA/mM cm2, respectively. Fabrication of the transformed electrode is shown in Fig. 12.17 [45]. A platinized Pt electrode was utilized as a support in the development of cholesterol

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FIGURE 12.16 A diagram representing the fabrication of ChOx/MWCNTs/ chitosan-modified GCE.

FIGURE

12.17 A diagram representing of the possible reaction mechanism of cholesterol and modification of the Chox/MoS2AuNPs/GCE at the modified GCE.

amperometric biosensor. The support was used for electropolymerization of polypyrrole film in which ferrocene monocarboxylic acid and cholesterol oxidase were coentrapped by Vidal et al. The detection limit and sensitivity was 12.4 μM and 88.51 nA/mM, respectively. Chatterje et al. prepared a biosensor for the investigation of cholesterol based on CNT and AuNP hybrid materials. The limit of detection found was 3 mmol/L, which was higher than

many reported biosensors based on other CNTs dispersion [46]. A sensitive and highly selective cholesterol electrochemical biosensor established based on upon electron transfer of a hemoglobinencapsulated chitosan fabricated GCE and successfully utilized for the analysis of serum samples by Zhao et al. The linear reaction of cholesterol concentrations was within the range of 1.00 3 1025
6.00 3 1024 mol/L, with

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12. CHOLESTEROL-BASED ENZYMATIC AND NONENZYMATIC SENSORS

0.9969 correlation coefficient, and an expected cholesterol limit of detection 9.5 μmol/L at a signal/noise ratio of 3. The cholesterol concentration changes with the slope of the calibration graph, the sensitivity was 0.596 A/M, and the computed standard deviation was found to be below 4.0% (n 5 5) for the detection of real samples [47]. Dervisevic et al. developed an amperometric cholesterol biosensor by using electropolymerization technique with polymers on pencil graphite electrode and immobilization of apo-cholesterol oxidase in flavinadenine-dinucleotide as a monolayer. Two well-known materials including 3-aminophenyl boronic acid and thiophene-3-boronic acid for electropolymerization onto pencil graphite electrode and immobilized with ChOx for the detection of cholesterol. The construction of the fabricated electrode is shown in Fig. 12.18 [48]. An amperometric electrochemical biosensor was developed by the process of coimmobilization of catalase and ChOx onto the ionic liquid/graphene fabricated on glassy carbon electrode (GR
IL/GCE) for the detection of cholesterol was reported by Gholivand et al. The fabricated sensor showed a good

FIGURE 12.18

A schematic diagram representing the fabrication of the PABA/FAD/apo-ChOx electrode.

response time below 6 seconds, good sensitivity of 4.163 mA/mM cm2, and a linear range of 0.25
215 μM with R2 . 0.99 [49]. A modified voltammetric sensor was utilized for the examination of total cholesterol and this was reported by Dernia et al. Coimmobilization of different enzymes was utilized to modify the sensor using HRP and ChOx on porous graphite. This revealed that the sensor had high sensitivity and high stability (16 μA/mM cm2) [50]. A novel cholesterol based on electrochemical
photoelectrochemical (PEC) dual-mode biosensor was described by Komathi et al. This sensor involves graphene (G) sheets linked together and fixed firmly with titanium nanowires (TiO2(G)-NWs) into 3D nanostacks (specified as G/Ti(G) 3DNS). This was achieved by using the favorable features of graphene and TiO2-NWs and resulted in high sensitivity and good selectivity for cholesterol examination. The G/Ti(G) 3DNS/CS/ChOx bioelectrode showed selectivity toward cholesterol with a low detection limit of 6 μM and significant sensitivity (3.82 μA/cm2 mM). The fabrication method is illustrated in Fig. 12.19 [51]. Screen-printed electrodes were used by Cinti et al. for the optimization and development of a cholesterol biosensor. Inkjet-printed Prussian blue nanoparticles were utilized for the fabrication. The developed biosensor showed a sensitivity related to cholesterol of 2.1 μA/mM cm2 (R2 5 0.97, n 5 5) with a linear range of 0
15 mM [52]. Due to good catalytic activity and conductivity of ZnO2 and silver nanowires, these were fabricated on an ITO electrode with chitosan and graphene oxide to produce a novel amperometric biosensor for the detection of cholesterol by Wu et al. The modified cholesterol biosensor showed a linear response in the range of 0.25
400 mg/dL (6.5
10) mM, sensitivity of 9.2 mA/mM cm2, and a limit of detection of 0.287 mM (S/N 1/4 3) [53]. The schematic representation is provided in Fig. 12.20.

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12.4 NONENZYMATIC SENSORS

FIGURE 12.19 (A) Utilizing G/Ti(G)-3DNS/ CS/ChOx occurring in different pathways in the fabrication of electrochemical
photoelectrochemical dual-mode biosensor. Stage (i); TiO2(G)-NWs altered to G/Ti(G)-3DNS Stage (ii); CS used to transform G/Ti(G)-3DNS then followed by immobilization of CS. (B) Electrochemical (i) photoelectrochemical (ii) sensing of cholesterol at G/Ti(G)-3DNS/CS/ChOx based biosensor. (i) Enlarged current signals shown by cyclic voltammogram in the detection of cholesterol (ii) Photocatalytic group existed upon light radiation at G/Ti(G)-3DNS/CS/ChOx biosensor as well as differences of amperometric photocurrents in the cholesterol concentrations.

FIGURE 12.20 A schematic fabrication procedure in the development of cholesterol biosensor.

12.4 NONENZYMATIC SENSORS A new sensor for nonenzymatic cholesterol determination utilizing a functionalized nanographite has been described by Bhawana et al. Under the optimized detection state, the fabricated sensor had a linear range of 50
500 mg/dL for cholesterol with 0.99784 correlation coefficient and sensitivity

of 1.0587 μA/mg [54]. Yang et al. has developed a platinum nanoparticles which were placed in a layer-by-layer accumulation with thin film carbon nanotubes through the spontaneous reduction of H2PtCl6 at the effective sites of the carbon nanotube surface. This fabricated electrode can also be utilized for the nonenzymatic determination of cholesterol by the use of chronoamperometry in neutral phosphate

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buffer solution. The sensor with 24 bilayers of CNTs displayed a low limit of detection with a broad linear range of 0.005
10 mM and a sensitivity of 8.7 μA/mM cm2 [55]. Li et al. developed AgNPs fabricated on GCEs as a working electrode and found that the sensor was good for nonenzymatic cholesterol determination, as it exhibited noticeable enhanced catalytic activity toward the cholesterol oxidation associated with that of the solid silver nanorods. In addition, under optimal detection conditions, the fabricated sensor had a linear response in a range of 2.8 3 1024 M
3.3 3 1022 M, the detection limit was 1.8 3 1024 M, and a signal-tonoise ratio of 3 [56]. A cover-type nonenzymatic free-cholesterol sensor was reported by Yoon et al. to determine cholesterol based on utilizing stainless steel microneedle patches and nanoporous platinum sensing electrodes. The developed sensor showed high sensitivity of 305 nA/mM cm2 and 0.964 correlation coefficient in 0.1 M phosphate buffer solution [57]. A new Schiff base (Z)-2-((pyridin-2-yl) methyleneamino) benzenethiol (2PMAB) was

developed by Goswami et al. via the reaction of picolinaldehyde with 2-aminobenzenethiol, which is a good electrode altering agent for cholesterol determination. The response time with GC/styrene/2PMABAg electrode was 3 seconds with a detection limit (3σ) 1.99 3 1025 M, linear range of 3.96 3 1025
37.03 3 1025 M, and sensitivity of 0.0722 V/mM. Rengaraj et al. demonstrated a high-quality graphene composite and nonenzymatic cholesterol sensor on a nickel oxide (NiO), for the detection of cholesterol. High sensitivity of 40.6 mA/μM cm2 was observed, with a quick response time of 5 seconds and low detection limit of 0.13 μM. The construction of the fabricated electrode is illustrated in Fig. 12.21 [58]. A nonenzymatic glucose biosensor based on monodispersed SiO2-coated magnetic Fe3O4 dispersed multiwalled carbon nanotubes (Fe3O4@SiO2/MWNT) was developed and described by Baby et al. This was successfully tested for quick examination of cholesterol and glucose. This fabricated electrode showed high sensitivity with a limit of

FIGURE 12.21 Diagram showing the modification of the NiO/graphene composite electrode for cholesterol sensing.

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333

FIGURE 12.22 The mechanism of cholesterol sensing utilizing Grp-β-CD as the working matrix.

detection of 1 PM for the determination of glucose and cholesterol [59]. Saha et al. developed a nonenzymatic cholesterol biosensor from coconut oil which was modified from the surface of a CNT electrode. The linearity in the concentration of cholesterol was in the range of 1 to 50 lM, with sensitivity of around 15.31 6 0.01 μA/μM cm2 and response time of around 6 seconds [60]. Another method for cholesterol determination was reported by Agnihotri et al. which demonstrated that the sensing process was based on electrochemical nonenzymatic process having several possess advantages as compared to conventional enzymatic processes. The detection limit was below 1 μM and the sensitivity was 0.01 μA/μM. Fig. 12.22 shows the reaction mechanism of cholesterol sensing utilizing Grp-β-CD as the working matrix [61]. A 3D-structured photonic quartz carbon inverse opal rods was utilized by Zhong et al. to constructed for the detection of nonenzymatic cholesterol sensor. More significantly, photonic crystal carbon inverse opal rods were utilized to determine cholesterol in human blood as well as in human serum. The

resultant reflection peak shifts of photonic crystal carbon inverse opal rods exhibited high linearity regarding the concentrations of cholesterol. The developed photonic crystal carbon inverse opal rods possess unique catalytic and conductive properties which make them more favorable for applications relating the examination of cholesterol in human blood [62]. For molecular sensing, colloidal gold is widely utilized due to the extensive flexibilities it provides when applied in the transformation of the gold nanoparticles surface with the changing of functional groups. An enzyme-free assay for the determination of cholesterol was reported by Raj et al. for the estimation of cholesterol in human serum samples. The linearity of cholesterol was within the range of 100
800 ng/mL, with a 0.9958 correlation coefficient and detection limit of 7075 ng/mL. The schematic illustration of etching of tomatine functionalized gold nanoparticles in the presence of cholesterol to form small gold nanoparticles is provided in Fig. 12.23 [63]. Saheny et al. developed a conductometric

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FIGURE 12.23

Schematic illustration of etching of tomatine functionalized gold nanoparticles in the presence of cholesterol to form smaller gold nanoparticles.

sensor for the detection of cholesterol in blood serum samples. This device was also utilized for the detection of cholesterol in the concentration range of 0.453
468.5 mg/dL in a solution and the obtained limit of detection was 1.2489 μg/dL [64]. The quantification of cholesterol, utilizing the principle of laser-beam propagation was performed by Budiyanto et al. using optical fiber bundle intensity profile through a solution containing cholesterol concentrations between 0 and 300 ppm. The result obtained showed that the output voltage is inversely proportional to the concentration of cholesterol with a sensitivity of 0.0004 mV/ ppm and a linearity above 97 %. The schematic representation of the fiber optic sensor used to

detect the concentration of cholesterol using fiber bundle is shown in Fig. 12.24 [65]. A selective and sensitive chemiluminescence sensor from peroxidase-like activity of copper nanoclusters was established for the determination of cholesterol by Xu et al. The chemiluminescence sensor intensity increases as the concentration of cholesterol increases over a broad range of 0.05
10 mM, with a detection limit of 1.5 μM. Fig. 12.25 shows the proposed scheme of the chemiluminescent cholesterol sensor which occurs in three steps [66]. An ITO glass plate was utilized for the immobilization of ChOx where the chitosan-tin oxide nanobiocomposite film was placed on it by Ansari et al. for the examination of

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12.4 NONENZYMATIC SENSORS

FIGURE 12.24

A setup diagram for a fiber optic sensor that investigates the absorption of cholesterol utilizing a fiber bundle.

FIGURE 12.25 Schematic illustration of the Cu NCs
based electrochemical sensor for cholesterol.

cholesterol. With the use of the modified cholesterol sensor, after 4
6 weeks at a temperature of 48 C, about 95% of the enzyme activity was retained with a rapid response time of 5 s, as well as detection limit and sensitivity of 5 mg/dL and 34.7 mA/mg dL cm2, respectively [67]. ZnO nanorods which were aligned vertically were utilized to modify the highperformance cholesterol sensor, as reported by Ahmad et al. The sensor was based on solution-gated field-impact-transistor where

the nanorods were grown selectively on a prepatterned substrate in solution. A real-time response was shown by the fabricating sensor in a broad range of the concentration of cholesterol and sensitivity of 0.001
45 mM and 10 μA/cm2 mM, respectively, and also showed good selectivity [68]. An efficient matrix of cerium oxide
graphene composites was utilized to construct sensitive and a simple electrogenerated chemiluminescence cholesterol biosensor by Zhang et al. A linear range

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based on the concentration of cholesterol was within 12 μM
7.2 mM and 4.0 μM [69]. A monolithic molecular imprinting sensor involving ceramic carbon electrode was reported by Tong et al. This was modified by the mixture of siliconalk-oxide, graphite powder, and MWCNTs to prepare molecularly imprinted polymer, followed by immobilization of the resultant complex mixture of components into the electrode cavity of a Teflon sleeve. Good sensitivity displayed by sensor with a linear range of 10
300 nM with a detection limit of 1 nM and a signal-to-noise ratio of 3 [70]. A new simple and sensitive technique for the detection of cholesterol involving development of electrochemical and enzymatic assay was reported by Nantaphol et al. The linear relation in the cholesterol concentration reduction current was found in the range of 3.9
773.4 mg/dL with a detection limit of 0.99 mg/dL [71].

12.5 FUTURE PROSPECTS Development in the biosensor sciences, signal transduction apparatus, sensor modification techniques, and data administration have been the key components in designing highly selective biosensors. In the near future, microelectromechanical system and nanotechnology could be the promising areas in developing next generation devices for point-of-care applications. Investigation on cholesterol-based sensors as active diagnostic applications is a massive challenge. Important use of many binder as well as immobilization mediators, biological ingredients, enzymes, or solids as electron collectors for designing of enzymatic identifying tools is vital to produce suitable current from biocatalytic methods. Recently, evolution in the field of advanced nanomaterials is greatly affected on the redox reactions as well as kinetic of transportation of electrons. Other critical points in the intention

and manufacture of new attractive sensor devices are the life time of enzymatic electrodes, reduction of enzyme electrodes, and rate of electron transfer. Complex sensing preparations with single and composite selective agents can be anticipated to have an effective influence in ecological monitoring as well as in wide-ranging clinical chemistry. It is clear that the cholesterol-based sensor arena offers exciting collaborative opportunities to address the existing challenges encountered by cholesterolbased electrochemical sensors.

12.6 CONCLUSION In this chapter, applications of different nanomaterial-based cholesterol sensors have been summarized extensively. The adoption of some novel methods, derived from biochemistry, physical chemistry, thin and thick-film physics, electronics, and materials science with the expected expertise has exposed the potential use in the improvements of feasible cholesterol sensors. The enhanced biocompatibility of nanomaterials related to the electrode surfaces are supported in extending the lifespan of the biosensor. Combination of nanomaterials in immobilization matrices for ChOx reduces the working potential of the bioelectrodes and, thus, removes the need for destructive chemical mediators. Moreover, the improvement of precise and inexpensive cholesterol sensors market, similar to the glucose sensor will eventually increase the market of biosensors for purpose of point-of-care applications.

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C H A P T E R

13 Recent Trends in Sensors for Health and Agricultural Applications Rajasekhar Chokkareddy1, Niranjan Thondavada1, Surendra Thakur2 and Suvardhan Kanchi1,* 1

Department of Chemistry, Durban University of Technology, Durban, South Africa, 2 eSkills CoLab, Durban University of Technology, Durban, South Africa

13.1 INTRODUCTION Present healthcare are structured and even enhanced to mitigate setbacks and spreading of infection of diseases still face a variety of challenges such as the fast rising and aging population with increasing healthcare expenditure. The 1990 band population of 357 million will multiply in 2025 to B761 million [1]. The fast-developing rate of healthcare and the aging population have produced several concomitant challenges to governments, healthcare workers, and the manufacture of healthcare products or equipment. These overwhelmed stakeholders use eHealth methods to expand healthcare service distribution, analytical monitoring, diseasetracking, and related medical techniques [2]. Moreover, the use of wireless technologies to passively and unobtrusively monitor patients in a discreet manner has received wide-ranging appeal since it is a consistent, cost-effective,

and ecofriendly method to assist patients [3]. Wearable gadgets is a category of technology that can be worn by a consumer and often include tracking information related to health and fitness. Other wearable gadgets include devices that have small motion sensors that take photos and link with mobile devices. A critical feature of wireless body area network (WBAN) is its aptitude to offer consistent communication for health procedures, particularly WBANs inserted in the human body. WBANs consist of light-weight, low-cost, small sensors which may be situated on the body as small smart coverings, incorporated into clothing (smart clothing), affixed under the membrane, or deeply implanted into body tissue [4,5]. These assist medical staff and doctors in key diagnosis and carefully and continually screen the health outlook of patients. Patientassociated data from WBANs are also linked to a central healthcare source for stable, safe,

*Corresponding author.

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and reliable records. The data is accessible to physicians to determine the patient’s state of health and can assist in notifying patients of necessary actions via alarm, SMS, or notice emails [1,6]. Over the past few years, the number of wearable, health-observing strategies has increased, alternating from pulse displays, movement monitors [1], and moveable Holter monitors to hi-tech and expensive implanted sensors [7]. Moreover, the latest technology developments include a combination of physical sensors, fixed microcontrollers, and wireless networking as well as radio edges on a single chip. Micromodifications have also assisted a novel group of wireless sensor networks for various uses such as in smart grids, health, and farming would be a potential development of thousands of sensors nodes or actuators. In this chapter, we will investigate how these methods have been incorporated into present day agro industry and health industries, and deliberate how the development of such equipment can assist our ability to feed the world’s population [8]. There are a number of physical and biological sensors that can dynamically display signs, as well as ecological sensors that display light, humidity, and

temperature, which may be successfully combined into wearable wireless personal/body area networks (WWBANs). Fig. 13.1 illustrates the WWBAN, which is an essential part of a multitier telemedicine system. It is predicted that the world’s population will increase from 7.4 billion in 2018 to 9.55 billion by 2050 and to 11.2 billion by 2100. To meet the demand for food without ecological destruction, however, temporary agricultural techniques will need to change drastically. Nowadays agriculture sector is a business which depend on a plethora of chemical and their physical quantities which harm environment, due to the usage of toxic agro-chemicals [9,10]. Many types of measurements are available, inter alia, soil superiority and biochemical configuration (nitrogen, pH, and moisture content) as well as atmospheric conditions which effect crop development (comparative humidity, wind, rain, solar energy, and airstream and direction) [11,12]. In the literature, a wide range of new technologies are focused on the agrifood chain and strategies for the sustainability of farming systems. For example, solar panels produce less charge consumption and enhances the battery performance. These

FIGURE 13.1 A schematic illustration of the wearable wireless body/personal area network linked with telemedicine for health monitoring.

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FIGURE 13.2 The Waspmote Plug & Sense! unit with three sensors.

technology uses wireless communication and limit the physical data collection which directly reaches to the consumer [13], for example “Waspmote Plug & Sense” by Spanish manufacturer, Libelium, a solar-powered sensor (Fig. 13.2). This self-powered machine can be connected with various data transmission channels including WiFi, Bluetooth Low Energy, a radio frequency between 868 and 900 MHz, and ZigBee. A 3 G/GPRS receiver may also be utilized to direct data from the sensor, even through mist, in the absence of transitional access. Further these type of sensors can easily coexist with other types of sensors, such as soil wetness content, soil temperature, solar energy, atmospheric conditions, leaf dampness, and direction of wind [14,15]. Recently healthcare-based methods develops greatly to help precise health and highquality healthcare. In healthcare, several devices are used based on their features, usability, and effectiveness. Smart sensors have a number of advantages:

• • • • • •

easy to design high reliability scalable, flexible system minimum interconnecting cables high performance small, rugged packing

Farmers, researchers, and manufacturers are joining forces in the agricultural industry by investigating new effective solutions for resolving various problems. The exponential improvement of sensors abilities and technologies allow these capabilities to attain above objectives. Furthermore, pH-sensitive, soil-based sensors allow to characterize different variables in the soil concerned with the crop productivity. An FPGA-based fused smart sensor for real-time plant transpiration monitoring is also possible.

13.2 SMART HEALTH SENSORS Kaiqi Su et al. [16] reported an enhanced efficient biochemical determination technique

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FIGURE 13.3 (A) The schematic illustration of Bionic e-Eye; (B) Bionic e-Eye image; and (C) the chief line of the home-based software iPlate.

called Bionic e-Eye. This smartphone system is used to examine marine toxins in various environmental conditions. The point-of-care based smartphone platform is a quick process that communicates with interface device spontaneously. Bionic e-Eye consists of approximate features with less measurement parameters, it is inexpensive, highly effective, and works properly for field-test submissions with the assistance of enzyme-linked immunosorbent. Moreover, Bionic e-Eye is a device used to monitor moveable objects by acting as a radiance provider. The transferrable attachment comprises a dark hood, broad-angle lens, and a section of shorter-power electro luminous which offer constant, low-power light. Fig. 13.3A is an illustration of the Bionic e-Eye device, while Fig. 13.3B is an image of a homebased Bionic e-Eye. The iPlate home-based software is shown in Fig. 13.3C [17]. Normally a standard curve is obtained by clients with a simple form of standardization. The standard curve permits easy calculation for sample concentration. Finally, the data may be shared among the users of iPlate through the upload function considering the exact condition. The iPlate is utilized as an iOS application which can be introduced by Object-C and Swift 1.0 in Xcode7 (Mac OS, Apple). This also consists of a

feature for image capture and data dispensation, loading, and transmission. The iPlate workflow entails a sample amount, calibration, and the sharing of data as displayed in Fig. 13.4. This iPlate was available in two color modes red, green and blue (RGB) and hue, saturation, value (HSV). In addition, these two color methods are utilized for data examination as well as in the metrics presentation and can be compared in phases of standardization and the sample quantity. In the calibration phase, iPlate was blank plate and essentially used for calibration curve detection [17]. Ling Zou et al. [18] reported an electrochemical immunosensor for the detection of okadaic acid using the Love-wave sensor with the immunogold stain technique. Moreover, okadaic acid BSA act as a anti-BSA and it easily forms antibody bond through staphylococcal protein A, this reaction was complete on the sensor superficial via Fc areas. The Love-wave sensor contains a dualchannel, namely, a reference and working channel, and is displayed in Fig. 13.5. The results confirmed that the electrochemical immunosensor definitely reacted to the okadaic acid with a linear range of 10
150 ng/mL and detection of limit 5.5 ng/mL [18]. Okadaic acid, is a diarrheic and highly harmful pollutant that has extensive circulation.

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FIGURE 13.4 Schematic representation of iPlate which contains sample quantity, calibration, and information distribution.

Hence, high-throughput, low-cost, wide-ranging, and moveable determination of okadaic acid is highly desirable for food protection and ecological monitoring. Kaiqi Su et al. [16] developed a new and transferable smartphone-based label-free biosensor using cell feasibility for the long-term monitoring of cell possibility. Moreover, this scheme was useful to process the mixture of image investigation and the cell contain kit-8 assay (CCK-8) to display the replication of the sensor (Fig. 13.6). The developed biosensor showed excellent performance to several okadaic acid absorptions, with a detection limit range from 10 to 800 μg/L [16]. Oncescu et al. [19] have described a self-diagnostic, smartphone accessory and software for the quantification of blood cholesterol. In addition, the smartCARD has an attachment that was present around the camera works as a detector of the smartphone as displayed in Fig. 13.7. The quantification of cholesterol by the electrochemical colorimetric reaction, which happens on a dry substance test strip over the complete range of physical cholesterol standards. Moreover, they have examined several

projects and categories of light sources to increase the strength of the arrangement and capability to deal with misalignment of the trial strip. Preechaburana et al. [20] developed a surface plasmon resonance (SPR; a traditional label-free determination technique) approach with the help of smartphones. In addition, they developed a initial angle-resolved superficial plasmon resonance based on a single-use device constructed using trained radiance and visual recognition from smartphones. SPR determination is demonstrated with a profitable analyze for β2 microglobulin (β2 M), which is a recognized marker for inflammatory illnesses, cancers, and kidney disease, which are estimated applicants for corresponding monitoring in dispersed situations. Furthermore, SPR determination is also possible with a made-to-order chip with the surrounded standardization method. Fig. 13.8A illustrates the performance of the SPR method with a commercial test chip. In addition, the use of fixed standardizations is critical in the universal arrangement of methods, hence the last sample (Fig. 13.8B) contains a

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FIGURE 13.5 (A) Schematic diagram of the Love-wave sensor wrapped on a reproduced circuit board. (B) SAW device with the reference network and working channel. (C) Poly(dimethyl siloxane)-based microfluidic chip with four air voids and two microfluidic cavities. (D) Image of a poly(dimethyl siloxane) microfluidic channel constructed Love-wave sensor.

customized chip with calibration stations combined with the device. Gallegos et al. [21] reported a smartphone as the determination tool for a label-free photonic crystal biosensor. In this determination, a custom-designed cradle embraces the phones in static arrangement with photosensitive mechanisms, for repeatable quantities of changes in the resounding wavelength of the biosensor. Moreover, they used as an iPhone 4 model

(Apple, Inc., Cupertino, CA, USA) which required modification of a framework that steadily embraces the cellphone rear camera in arrangement with the PC and the determination optics. Fig. 13.9 shows a schematic illustration of the detection system. Point-of-care testing refers to any analytical test completed at (or near) the site of the patient. In addition, the capability to transfer diagnosis data via WiFi is a challenging and

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FIGURE 13.6 (A) Construction of cell viability biosensor; (B) an image of a transferrable smartphone-based device; and the (C) chief interface of the home-based software iPlate screen.

FIGURE 13.7 Schematic illustration for cholesterol testing on a smartphone with precision equivalent to marketable point-of-care strategies.

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(A) Inter-facial device for a commercial Biacore CM5 test chip functionalized for the determination of β2 microglobulin (B) The test chip with fixed calibrations (H and L for high and low references, respectively) with a straight quantification of the unidentified test value (T), the two solid outlines in (b) specify the error margin.

FIGURE 13.8

FIGURE 13.9

(A) A schematic representation of the visual mechanisms inside the smartphone cradle. (B) The image of the cradle with a PC biosensor slide introduced into the discovery slot. ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

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FIGURE 13.10 (A) Gene-Z image. (B) Fluorescence image of a single 15-well showed after extension of fluidic cross-talk among reaction wells. (C) DRn for the wells revealed in (B) as measured at the end of the reaction using the Gene-Z, close-fitting active optical separation of the response wells.

significant features of upcoming point-of-care strategies. Stedtfeld et al. [22] developed an inexpensive, accessible, and solid scheme (Gene-Z) for fast, measurable determination of various genetic indicators with good selectivity. The Gene-Z is functionalized device via an iPod Touch, which also collects information and transfers reports through a WiFi boundary (Fig. 13.10). Moreover, this investigation offer data relating to the device with good selectivity and reproducibility through the genomic DNA Staphylococcus aureus and Escherichia coli. Generally, the Gene-Z signifies an important phase in genetic research to reduce the cost for point-of-care applications.

13.3 SMART AGRICULTURE SENSORS Due to the rapid advancement of the Internet, real-time raincloud calculations have become possible and have catapulted a phenomenon known as smart agriculture. Smart agriculture, established through the combination of material and devices in cultivation to gather a large capacity of information [23]. In addition, smart agriculture trusts the statistics diffusion and the care of data in isolated

packing to be used in some systems, which allows the mixture of investigation of several range of data sets for the generation of good results. Smart agriculture is the future of the food-producing sector along with novel procedures to affirm the world’s food security. Traditional agriculture was regularly undertaken within a family or village and acquisitive farming knowledge and information were passed down over the generations. Smart agriculture (also known as the precision agriculture) allows farmers to increase their yields while simultaneously reducing resources. Resources saved include fertilizer, water, and seeds [24,25]. Through strategically organizing sensors and plotting fields, farmers can initiate a process to understand their yields at a micro level which, in turn, preserves resources and decreases the negative influences on the environment. Smart agriculture has origins going back to the 1980s when global positioning system (GPS) ability became available for civilian use [26,27]. Furthermore, when farmers precisely map their crop fields, they might screen and apply compost and unwanted plant activities only to applicable zones. Throughout the 1990s, early care farming users accepted crop monitoring to produce fertilizer and pH alteration references. For

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example, huge changes could be measured and input into the crop model, that is more precise approvals for fertilizer application, irrigation, and uniform peak crop harvesting could be made. In smart cultivation many new trends has been followed by the farmers and have overcome several issues namely shortage of labor and in-time usage and name a few [28,29]. Generally, the smart agriculture can help to minimize environmental pollution and ultimately decreasing the prices of agricultural products [30]. Smart farming is a concept that was designed using software and computer science methods [31] attained through the novel methods by analyzing other data from farming effectively [32]. These calculated features are implanted in substances and unified with other Internet sources. Moreover, smart farming includes other relationships with comparable significance to smart agriculture. Hence, overlying edges and machineries occur such as accuracy farming and supervision material schemes in cultivation, which resulting the awareness of the farmhouse organization material system [31]. The farm organization information system is precise by gathering, storing, processing, and distributing data in a vital format to achieve processes and plays a key role in rural areas [33]. Several detection approaches are used in smart agriculture, so that data could assist farmers to monitor and enhance crop production, as well as adjust ecological features, including sensors signals from GPS satellite television to regulate longitude, latitude, and height. For this determination, three satellites are required to triangulate a location. Specific placing is the keystone of smart agriculture [34]. In addition, many GPS united circuits exist, but the NJRNG1157PCD-TE1 is a good example of a location-based sensor. Furthermore, optical sensors are also used for the determination of soil properties. These types of sensors measure several frequencies of light reflectance in nearinfrared, mid-infrared, and separated light

spectra. Hence, optical sensors are in automobiles or floating platforms such as drones or satellites. The soil reflectance and plant color information are two objective variables from visual sensors that can be combined and processed [35]. Optical sensors are mostly used to determine clay, organic matter, and humidity content of the soil. Although many hundreds of photodiodes and photodetectors are commercially available, most farmers choose a basic building block for optical sensors (Fig. 13.11). Xue-fen et al. [36] reported a smartphonebased LoRa in-soil spread quantity for wireless underground sensor networks. Generally, LoRa offers novel statement solution for wireless alternative sensor networks in soil. The features of LoRa in-soil circulation are mostly associated to its physical layer. In addition, LoRa has an option to forward error correction code and cyclic redundancy code, good LoRa data (Fig. 13.12). The scientists established a solid smartphone-based spectral analyzer that can be used in-field to collect reproduction spectra from soil. In addition, they anticipated to find correlation among the soil organic fillings and the restrained reflection spectra from soil models. Meng et al. [37] developed a smartphone-based device for the determination of soil organic matter. In addition, the modified smartphone-based spectrometer contains of a 3D-printed cage and some low-cost optical mechanisms. The compressed spectrometer seems to be good and simple spectrometer for the detection of soil organic matter. In comparison to the commercial spectrometers, the footprint of the scheme’s smartphone-based component is significantly reduced. Moreover, the reliability and precision of the reflectance dimension was achieved by the researcher after modifying the model chamber in combination with the smartphone-based spectrometer (Fig. 13.13).

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FIGURE 13.11 Vishay photo IC sensor.

FIGURE 13.12

Schematic illustration

of the LoRa Wizard.

13.4 GLOBAL MARKET FOR SMART SENSORS Over the past few years, smart sensors have progressed to convert completely a smart phone as a stand alone or a interface systems,

compared to the conventional techniques. Initially, they did simple functions, like changing impulses into electrical energy. Rapid progress in technology have led to the development of reduced types of smart systems with improved sensitivity and reliability

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FIGURE 13.13

metrics. In addition, smart sensors offer sensing and calculating abilities in a distinct device. The scope of these sensors is quickly fading, with modifications in nanotechnology and measurable science. Silicon is the mostused material in the development of smart sensors. Since, smaller varieties of sensors are more flexible, they can be inserted in a change of devices to attain real-time information. The “aptitude” element of these sensors separates them from new alternatives in the market. These sensors have advanced from being distinct devices calculated to determine definite properties to fully combined sensor systems, offering improved calculating capabilities. Moreover, there has been a stable shift to the miniaturization of these devices, using microelectrical mechanical systems. The global smart sensors market was estimated at US$25.96 billion in 2017 and is predicted to reach US$72.39 billion by the end of 2023, with a CAGR of 18.64% through the projection period (2018
2023). The overall demand for smart phones is estimated to rise gradually due to the advances options and availability of sensors in the device. Similarly, the expanding level of smartphone usage has assisted customers to opt for solutions that are more appropriate. Smart homes not only improve energy preservation by using smart meters, but also

Smartphone-based spectrometer.

have the ability to transform a home into a high-tech home using the internet of things.

13.5 FUTURE PROSPECTS Investigation of Unused Sensors. There are several built-in sensors in smartphones which have yet to be applied to agriculture. Physical sensors, for example, barometer and humidity sensor, can be rather simply extended to the agricultural field due to their unique future and functions for ecological sensing. This leaves an open chance for researchers and inventors to discover suitable applications that are required to improve these sensors. Extension of Sensor Usage from Present Research. Furthermore, the sensors that have already been discovered can promotly be explored by adopting previous methods to smartphones. Computer image procedures have been investigated in the context of agriculture broadly, numerous of which were offered as offline calculations, for example, weed identification [38] and cotton foreign matter inspection [39]. Smartphone Battery Usage Reduction. We observed that the majority of the literature studied did not focused on the batteryeffective feature to improve its applications [40]. One characteristic of farm work is that

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REFERENCES

farmers are usually away from a power source, preventing them from refreshing their phones once they run out of battery. So, strategy and enlargement of a battery effectiveness application is vital for the practicality of any smartphone application and might be alternative future way for investigation in this zone. Solutions for Exterior Factors Affecting the Efficiency of Smartphone Applications. Many problems have been described in using smartphone applications, for example, poor mobile network coverage [41]. It is necessary that network signals will be constantly improved to increase the effective usage of smartphone sensors in remote areas. Despite the challenges facing the active utilization of wireless sensing methodologies in biomedicine, there are positive visions for the future of wireless sensing in biomedicine. Many researches are constantly developing novel technologies to overcome the main problems of energy, isolation and safety in wireless sensor networks. Upcoming smart home locations will have multimodal sensor solutions. Universal healthcare uses will become abundant. The efforts are ongoing to improve the sensors efficiency to determine cancerous tumors cells among various types of cells, thereby identifying cancerous cells and other diseases [42].

13.6 CONCLUSION In the fields of agriculture and health, the advancement of mobile sensor technology remains somewhat limited. There are several scientific challenges. The challenges facing the mobile devices include input (difficulties entering characters), form factor (small size of the window), skills (uncertainty of device usage), communication (link speed and availability), and understanding (presentation and applicability of applications). In this chapter, we outlined recent developments in smartphone-

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based sensors related to healthcare and agricultural applications. Some of these applications discussed previously were used for medical purposes, but with some additional integration presented for novel applications. This can be upgraded to sensor-based documentation using applications and wireless networks. There are, however, many potential safety setbacks that need to be considered relative to the placement of mobile sensors for healthcare. Despite this, the prospective of smartphones, improved by many built-in and add-on sensors is enticing and limitless. The main development is established to endure for the predictable upcoming, and a request related to the growth of food remaining, agriculture will definitely organize an important and essential segment of a complete sensor at market place.

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C H A P T E R

14 Hybrid Carbon Nanostructures for Chemical and Biological Sensors Pratibha1, Supriya Singh1, Sudesh Kumar1, Kakarla Raghava Reddy2,*, S. Naveen3 and Veera Sadhu4,* 1

2

Department of Chemistry, Banasthali University, Banasthali Vidyapith, Vanasthali, India, School of Chemical & Biomolecular Engineering, The University of Sydney, Sydney, NSW, Australia, 3 School of Basic Sciences, Jain University, Bangalore, India, 4School of Physical Sciences, Banasthali University, Banasthali Vidyapith, Vanasthali, India

14.1 INTRODUCTION Graphene is a light-weighted universal material [1]. It is springy and more efficiently conducts electricity at room temperature than the other carbon materials. It is more rigid than the diamond [2]. Graphene is made of two-dimensional (2D) sheets of carbon atoms bonded by sp2 bonds in a hexagonal configuration [3]. Graphene is the most current associate of the multidimensional carbonnanomaterial family. For example, fullerene is a zero-dimensional material, single-walled carbon nanotubes (SWCNTs) are one-dimensional (1D) nanomaterials, and graphite is a three-dimensional (3D) substantial material as shown in Fig. 14.1 [5]. The best material among these for electrochemistry is graphene [6
9] because it has a large surface area, low cost, and large 2D electrical conductivity [6].

In the Dirac (relativistic) equation, the charge exporter conveyance in graphene is candidly explained. Hence, charge exporters are named as massless Dirac fermions [10]. Graphene can be defined as “one of the forms of carbons consisting of planar nanosheets with atoms organized in a 3D frame” [11]. Carbon materials form with different structures as shown in Fig. 14.2.

14.2 SYNTHESIS OF 3D STRUCTURED GRAPHENE NANOSHEETS 14.2.1 Exfoliation Process Exfoliation increases the reachable surface area of the material. It significantly improves the chemical and physical reactivity of surface-active

*Corresponding authors.

Advanced Biosensors for Health Care Applications DOI: https://doi.org/10.1016/B978-0-12-815743-5.00014-7

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FIGURE 14.1 Applications of graphene-based electrochemical sensors [4].

FIGURE 14.2 Different nanostructured carbons [5].

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or catalytic materials [12]. Thickness of graphene sheets will be changed during the exfoliation process [13]. As a thermally expansive material, a layered material can be used when heat causes exfoliation. Hence, for the fire obstacle in paints and firestop pillows, graphite and vermiculite are used because they produce a residue of lowthermal conductivity and reduce density upon heating. After the exfoliation process, a graphene monolayer can be formed from bulk graphite material [14]. Graphite is converted into a graphene sheet structure due to changes in the chemical and electronic properties as shown in Fig. 14.3. 14.2.1.1 Liquid Exfoliation of Layered Materials The layered materials have 2D nanoplate structures that are feebly fixed and convert to 3D structures, for example, the graphite formed from graphene monolayers. MoS2 and layered clays form more unusual materials such as MoO3 and Bi2Se3. These materials show an extensive series of optical, mechanical, electronic, and electrochemical properties. Over the past few years, many approaches have been established to exfoliate layered materials to yield monolayer nanosheets. After the exfoliation process, the surface area of graphene sheets will be increased. According to an advanced discovery, layered gemstones can be exfoliated in liquid. A number of procedures are used for this, such as

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oxidation, ion exchange, or apparent passivation by diluters. The development of thin films and complexes is endorsed by liquid exfoliation which is potentially measurable [15]. According to previous studies, the ultrasonication process has an effect on the number of layers of graphene [16]. Graphene sheets are well dispersed in the good solvents and in the ultrasonication process (Fig. 14.4). 14.2.1.2 Graphite Exfoliation via the Surfactant-Assisted Emulsion Process For the development of graphene, a largescale production method has been used and there has been significant interest in it. The exfoliation of graphene sheets in the aqueous system is superior to the organic phase. Oxidation and subsequent graphite exfoliation is the most common technique for preparing graphene oxide [17
20]. However, in terms of yield and amount, this process has significant disadvantages. The most significant disadvantage is the formation of structural defects as the result of the oxidation processes is evidenced by Raman spectroscopy [17
20]. The structural defects in the graphene sheets are known to significantly affect their electronic and chemical properties [21]. Even after strengthening at 1000 C, it is impossible to remove these defects completely. Some structural defects can be removed by chemical oxidation of graphene [22].

FIGURE 14.3 Crystal assemblies, indeed arising methods, and exfoliated yields for four examples of coated constituents. (A) Graphite involves interchanging piles of hexagonally organized carbon atoms (black domains), (B) is an indeed arising crystal, and (C) exfoliates to single atomic coats of carbon termed graphene [15].

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FIGURE 14.4 Graphic explanation of the main liquid exfoliation appliances. (A) Ion intercalation: ions, which are yellow spheres, are introduced among the coatings in a liquid atmosphere, bumping the gemstone and waning the intercoat fascination. Then, anxiety (like shear, ultrasonication, or thermal) can entirely distinct the coatings, causing an exfoliated scattering. (B) Formation of complex with graphene sheets and metal ions. (C) Well dispersion of graphene nanosheets through an ultrasonication process. In “good” diluters like suitable apparent energy—the exfoliated nanosheets are steadied beside reaggregation. Otherwise, for “bad” diluters reaggregation and sedimentation will happen. This appliance also defines the distribution of graphene oxide in glacial solvents like water. NB: solvent fragments are not displayed in this image [15].

Exfoliation of graphite in the liquid phase has been shown by two independent groups to result in defect-free monolayered graphene [23,24]. Depending on the properties of solvents used in the exfoliation process, the surface area of graphene sheets will be changed [24]. However, there are some drawbacks in this process such as these solvents being expensive require special care during handling. It is difficult to produce graphene monolayers using solvents that have a high boiling point. Therefore a liquid phase process is needed, which results in a reasonably high yield of graphene from the exfoliation of

graphite. The technique should be nonoxidative and do not need high-temperature processes or chemical postactions. Solvents with a low boiling point are favorable to produce monolayered graphene [24
29]. 14.2.1.3 Exfoliated Graphite Oxide via the Deoxygenation Process in an Alkaline Solution Large-scale graphene sheets can be prepared using the chemical reduction process with reducing agents such as hydrazine and dimethylhydrazine [20,30,31]. But these reducing agents are very hazardous. Some

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polymeric surfactant molecules may cause agglomeration of graphene sheets [30]. One method is to simply heat an exfoliated-GO suspension to get a stable graphene suspension under intensely alkaline conditions and moderate temperature (50 C
90 C) (Fig. 14.5A). The concentration of alkyl compounds influences the physicochemical properties of graphene sheets [33]. After treatment with NaOH, the color of graphene dispersion was changed from yellow to black. The exfoliation of graphene oxide and formation of graphene sheets with high yield can be achieved in the presence of strong alkaline solution (Fig. 14.5B) [32]. Let us consider the scenario where negatively charged oxide functional assemblies are completely removed. Negatively charged graphene will be formed under alkaline conditions; this permanency can be endorsed to a reinforced electrostatic stabilization as underneath higher pH values, and the revulsion among negatively charged graphene sheets should grow. Wallace et al. observed similar

FIGURE 14.5 (A) Conversion of graphene sheets from GO through deoxygenation process. (B) photograph of exfoliated GO diluted solution (0.5 mg/mL), and (C) exfoliated GO at 90 C [32].

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results [18]. The results from the atomic force microscopy showed that graphene has singlelayered structure. The thickness of graphene was 0.8 nm, which is equivalent of singlelayered graphene [18,34,35]. The reduction or deoxygenation process has been used for the conversion of GO to graphene sheets [36,37].

14.2.2 Graphene Synthesis by Microwave Irradiation Different strategies such as micromechanical breaking and thermal treatment have been used for the production of graphene sheets [32,38
40], epitaxial evolution on SiC exteriors [41,42], and chemical lessening of GO which is exfoliated [31,43
46]. High temperature and extensive processing eras are essential in most of these methods. For example, the thermal exfoliation of GO requires temperatures over 1000 C [39,40]. The microwave plasma enhanced chemical vapor deposition (MWPE-CVD) method has been used to prepare four- to six-layered graphene at 700 C [47]. The synthesis of three dimensional structured graphene has been established through a substrate-free, microwaveplasma process [48]. Chemical reduction using strong reducing agents like hydrazine is an efficient technique for conversion of GO to graphene sheets [43
45]. Microwave irradiation (MWI) process has been used for the production of various nanostructured materials such as inorganic metallic particles, metal oxides, metal alloys, and semiconducting materials [49
52]. MWI was also used to prepare soluble SWCNT derivatives and intercalated graphite [53,54]. MWI has some advantages in preparing quality graphene sheets compared to other methods [55].

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14.2.3 Graphene Synthesis via Chemical Vapor Deposition The CVD technique is used to prepare graphene sheets with good yield [56]. This is a remarkable option for the large-scale production of low-cost, high-quality graphene. The effective production of constant multicoated 3D graphene thin films on the surface of Si substrates is managed through the MWPE-CVD process. Graphene-based materials have been applied as electrochemical sensors [1]. Bhuvana et al. [57] developed novel carbon materials having petal morphological structures (Fig. 14.6). CVD method has been used to fabricate graphite or doped-graphite thin films [57,59]. In 1990s Johansson et al. effectively developed few-layer graphite by CVD [59]. More recently CVD has been extensively used for the production of graphene [56,60]. There are several methods to attain single-coat graphene or fewcoat graphene. The CVD process with nickelbased catalyst produced 20-layered graphene [61]. Concentration of nickel catalyst in CVD process influences a number of layers of graphene. Single-layered graphene showed good electrical and electrochemical properties than multilayered graphene sheets [62]. The thermal chemical vapor deposition (TCVD) process was used to prepare a high surface area of graphene in the presence of transition metal
based catalysts such as copper,

FIGURE 14.6 The catalyst-free production of three dimensional carbon nanosheets or petal structures on the graphite fibers via the MWPE-CVD process [58].

iron, and nickel [63,64]. In the presence of Ni-based catalyst, monolayered and bi layered graphene sheets were formed with the yield of 87% [64,65] and 95% of monolayered graphene was formed in the case of copper-based catalyst [66]. Classic CVD process was used to develop vaporous hydrocarbons in the presence of methane, ethylene at 1000 C [67
71], and acetylene at 650 C [72]. Liquid-type precursors have been used to fabricate carbon nanotubes (CNTs) [73
76]. Single-layered coated graphene produced on Ni halts in an argon environment through refrigeration after the CVD process. The formation of graphene film on the surface of metallic substrate was found [77]. It has been reported that graphene layers were developed from pentene and ethanol below atmospheric compression [78]. The ethanol was recycled to yield graphene sheets through the CVD process on the surface of a stainless brace [79].

14.3 GRAPHENE-BASED ELECTROCHEMICAL SENSORS AND ELECTRODES FOR DETECTING BIOMOLECULES Papakonstantinou et al. [1] used graphenebased nanostructured materials for electrochemical sensing applications. As shown in Fig. 14.7, they produced graphene platelets on the surface of silica substrates through catalyst-free technique. GNP-based electrode materials were fabricated from glassy-carbon electrodes and used as electrochemical sensors for the detection of dopamine, ascorbic acid, and uric acid.

14.3.1 Field Effect Transistors
Based Biosensors Using Functional Graphene Materials The FET or field effect transistor is a three terminal device that uses an electric field to

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FIGURE 14.7 Graphene platelets (GNPs) intended for electrochemical detection. (A and B) Low-magnification electron microscopic images of GNPs at various extensions; (C) high-magnification TEM image of GNPs; and (D) EDS spectrum of GNPs coated on substrates [1].

control the current flowing through the electronic device. FET activates via possessions of an electric field on the drift of one category of the electronic transporter (electrons/holes) from basis to channel via a single category of semiconducting materials. Tao et al. [80] have developed highly electrically conductive graphene sheets for electrochemical applications [80]. Furthermore, depending on graphene FETs, Frei tag et al. [81] explored graphene thermal transportation possessions via Raman scattering microscopy of a 2D

phonon ensemble. Graphene FET was used as an efficient candidate for sensing applications [82,83]. Biosensors based on FETs can compromise benefits as related to supplementary electrochemical sensors in the expressions of greater sensitivity, excellent selectivity through debauched retort interval, and nanoscale invention techniques and expedient proportions. As shown in Fig. 14.8, graphene is functionalized with biomolecules and utilized as FET-based biosensor [84].

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14.3.3 Cytochrome Biosensor

FIGURE 14.8 Graphic representation of glucose oxidase functionalized graphene FET [84].

14.3.2 Saccharides-Based Biosensor It is known that type 1 and type 2 diabetes is linked with saccharide content (e.g., glucose) in the human system, and causes diseases such as heart failure, blindness, kidney damage, tissue degeneration, and even death. 3D structured graphene has been applied to detect various saccharides such as sucrose, fructose, galactose, maltose, glucose, and others. Graphene-based functional hybrid materials have been applied to detect various saccharides such as glucose, fructose, lactose, galactose, xylose, maltose, sucrose, and others. Bacteria, fungi, and virus are causes for serious diseases [85]. It is necessary to determine the glucose concentration in the blood. Presentation of graphene in costeffective and sensitive biosensors can support the aforesaid virus [86]. Wang et al. [87] used graphene sheets as electrode material for developing glucose sensors. Various graphene-functionalized metal and metal oxide nanoparticles such as Ni, Cu, Ag, Au, CuO, SiO2, Co3O4, and NiCO2O4 have been well studied for the analytical and sensing performance of glucose. The developed graphene electrode showed good sensitivity and reproducibility for detection of glucose [88].

Cytochrome-c (Cyt-c) is a small metalloprotein that is present in the compartment between the inner and outer mitochondrial membranes, and functions to transport electrons between complexes III and IV of the respiratory system. Therefore Cyt-c plays a key role in determining the cell apoptosis process. The electron transfer takes place between electrodes and Cyt-c and is getting increasing investigation consideration due to its biological functions. However, pure glassy carbon electrodes cannot detect Cyt-c [89]. Therefore carbon materials have been used to modify their surface. It was reported that a novel graphene material consisting of graphene/ Ru(II) metal complex-CeO NPs-chitosan can be an outstanding sensing platform for the construction of electrochemiluminescence (ECL) sensors [90]. Finally, the surface of the electrode was incubated with bovine serum albumin to block nonspecific binding sites. This apta-detector showed a wide dynamic range from 0.2 to 100 nM with a low sensing limit of 25 pM that is applicable for the analysis of Cyt-c in order to find early-stage apoptotic cells. Likewise, GO-based materials have been widely used to construct nanovehicles and modulators of different enzymes such as Cyt-c, chymotrypsin, and serine protease through the interactions between graphene and biomolecules (proteins).

14.3.4 Fluorescent Biosensor Nicotinamide adenine dinucleotide (NADH) is a significant coenzyme that proceeds in more than 300 dehydrogenase enzymatic retorts [91]. D-lactase electrochemical biosensor was fabricated by modifying graphene electrode with biomolecules

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[92
94]. At an extraordinary overpotential (c.0.5 V), its reduction at simple glass carbon electrodes in neutral resolutions transpires. This is due to gentle electron transfer kinetics and electrode entangling. So, at low potential, the active oxidation of NADH would support the expansion of NADH-based biodevices [9]. The corrosion of NADH assists as the anodic indicator. It was used as sensor for detecting alcohols, glucose, lactic acids, and others [95]. Akhtar et al. [96] developed an amperometric sensor for the detection of NADH using activated graphene oxide modified with polyethyleneimine (PEI) and ethylenediaminetetraacetic acid (EDTA). This electrode material (GCE/AGO/PEI-EDTA) showed electrocatalytic properties toward sensing NADH at a low oxidation potential due to strong interaction between NADH and EDTA polymer modified activated carbon. The dynamic range of NADH is between 0.05 and 500 μM with a detection limit (LD) of 20 nM. This study demonstrates that the sensor device has potential in the biomedical field for sensing NADH in tumorigenic lung epithelial cells.

14.3.5 Horseradish Peroxidase Biosensor H2O2 has been used for environmental protection and clinical device [97]. The electrochemical approaches are more appropriate for this rather than supplementary predictable procedures. An horseradish peroxidase (HRP)-improved electrode was established for the exposure of H2O2 and demonstrated extraordinary electrocatalytic movement to H2O2 with great sensitivity, squat detection limit, extensive linear variety, and wild amperometric reaction. HRPbased electrodes showed good reproducibility, sensitivity, and selectivity [98].

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14.3.6 Lipoprotein-Based Biosensor Lipoprotein (Lp) contains large molecules that are a combination of proteins, fats, and cholesterol, and is present in human blood. High levels of Lp can lead to the risk of atherosclerosis and calcification in the arteries to the heart, heart valves, brain, and legs. It also represents risks for heart attack, stroke, and blockage of heart valves. They are also originators of bioanalytes like steroid hormones and bile acid. However, the high level of cholesterol can cause heart diseases, arthrosclerosis, and other diseases, so that it is very important to develop sensors for the detection of cholesterol. The sensitive amperometric biosensor was developed by a coating of Pt nanoparticles on the surface of graphene electrode. The sensitivity and detection limit of the electrode for detection of cholesterol were 2.07 μM/cm2, and 0.2 μM, respectively. The excellent sensitivity was achieved due to synergistic effect of graphene with Pt nanoparticles [99].

14.3.7 Iron-Based Biosensor Hemoglobin (Hgb) is a protein biomolecule that is a component of red blood cells and plays a key role in carrying oxygen from the respiratory organs (lungs) to other parts of the body. The structure of Hgb consists of four peptide chains and each chain is connected with a heme group that consists of an iron (Fe) chelated to a protoporphyrin ring. The determination of Hgb level in the blood is useful for the diagnosis of several diseases such as heart diseases, anemia, and leukemia. A chitosan
graphene hybrid electrode has been developed and used as a sensor for the detection of Hgb. The response of Hgb over glassy carbon electrode was in the range of 30
150 mV/s [100]. The influence of magnetic nanoparticles on Hgb exposure was studied.

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H2O2 sensor with lower detection limit has been developing using graphene based electrodes [101]. Zhan et al. [102] reported the fabrication of a 3D graphene/carbon nitride/ layered double hydroxide hybrid (Gr-LDHC3N4) by a hydrothermal-calcination process. This hybrid and chitosan were used to immobilize Hgb on a carbon ionic liquid electrode for the fabrication of a trichloroacetic acid biosensor. It was reported that this electrode showed a wide linear sensing range from 0.2 to 36 mM with a low DL of 0.05 mM. Due to the effective electron transfer ability, large surface area of the conductive Gr-C3N4 hybrid, and the good biocompatible property of chitosan, the constructed sensor showed short response time, good stability, and high sensitivity.

14.3.8 Dihydroxy Aromatic Compound
Based Biosensor Catechol is an unsaturated six-carbon ring phenol compound (benzenediol). Bhat et al. [103] developed PdAg alloy nanoparticles (NPs)-decorated graphene in the presence of Pd(OAc)2 and AgNO3 as precursors of required metal ions. It was found that the PdAg
Gr alloy electrode exhibits remarkable electrochemical sensing properties for the detection of catechol (0.06 μM).

associated with graphite/glassy carbon compared to basic glass carbon electrodes [9].

14.3.10 Neuron-Based Biosensors Dopamine (DA) neurotransmitters are located in the deep-middle region of the brain. Dopamine is a significant neurotransmitter which plays a vital role in the nervous, hormonal, renal, and circulatory organisms [106]. Basiri et al. [107] developed reduced graphene oxide
coated Ag nanoparticles and used it as a sensor for the detection of DA. The limit of detection (LOD) of this electrode material for DA was estimated to be 16 nM.

14.3.11 Ribose-Based Biosensor Electrochemical DNA sensors have emerged as a promising class of biosensors capable of detecting a wide range of molecular analytes such as nucleic acids, proteins, small molecules, and inorganic ions [108,109]. Graphene-based electrodes were used for electrochemical sensing of DNA [107]. Han et al. [110] developed electrode with the combination of graphene and Ag nanoparticles, and used it as reusable surface-enhanced Raman scattering (SERS) sensors for the detection of DNA. The noncovalent interactions between DNA and graphene mediated trapping and release of DNA with efficient SERS detection.

14.3.9 Peroxide-Based Biosensors H2O2 is a vital moderator in clinical, industrial, food, environmental, and pharmaceutical analyses [9]. Graphene and other carbon nanomaterials have been used in creating biosensors for distinguishing H2O2 [9,91,92,104,105]. The electrochemical behavior of H2O2 on the graphene electrode was improved and displayed significant growth in the transfer of the electron proportion

14.3.12 Detection of Ascorbic Acid Ascorbic acid (AA) is a C vitamin and is easily soluble in an aqueous system. Hence, vitamin C is necessary for good health [111]. It is well known that AA is water-soluble antioxidant vitamin [112]. It plays a significant role in the mitigation of excessive cellular reactive oxygen species activities caused by number of abiotic stresses [113]. These consist of high-

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performance liquid chromatography [114,115], liquid chromatography
mass spectroscopy [116], spectrophotometry [117,118], and voltammetric methods [119
121]. Different electrode materials have been developed for the detection of AA [122
124]. Graphene-modified electrode used as electrochemical detection of ascorbic acid [125,126].

14.3.13 Electrochemical Sensors for the Detection of Various Chemicals Electrochemical sensing is an effective technique for the quick detection of various chemicals and biochemical species in various samples such as water, air, and food, as these sensors offer low cost, easy operation, and excellent sensitivity. The graphene-based electrode materials show excellent characteristics such as controllable size, morphology, electrical conductivity, mechanical properties, and template suitability for the dispersion of dopants and modifying agents. Graphene oxide
based hybrid electrodes were studied for sensing various chemicals such as alcohols, chloroform, phenols, acetone, different volatile organic compounds, and gases (e.g., NO2, H2S, NH3, CO, and LPG gas). A one-step hydrothermal method was employed to prepare GO coated with ZnO nanoclusters and used it as an electrochemical sensor for the detection of bisphenol A [120]. The electrode has a detection limit of 2.1 nM. Such hybrid electrodes can also be applied for the detection of bisphenols present in food products such as milk, among others. Li et al. [121] prepared reduced graphene oxide
anionic metalorganic framework hybrids via a cationic exchange and reduction process, and investigated the hybrid electrode for the determination of H2O2 released from living cells. They observed that the electrocatalytic properties of the electrode depended on the content of

367

anionic metal-organic frameworks. The hybrid electrode showed a linear detection range from 4 to 11,000 μM with a DL of 0.18 μM. Li et al. [122] developed a sensitive electrochemical sensor for the detection of paracetamol using Au-graphene@PEDOT ternary electrode, which showed high sensitivity in a wide range of concentration from 0.15 μM to 5.88 mM, with a DL of 41 nM. This sensor also showed long-term stability, demonstrating its potential for the detection of paracetamol present in various pharmaceutical products. A new rGO-Pt@SnO2 ternary electrode was developed by Peng et al. [123] and was utilized as a sensor for the detection of methanol (40 ppm) at low operating temperatures. This sensor exhibits a short response time of only 6 seconds. rGO coated with SnO2 nanoparticles were fabricated via a solvothermal process and were used for the detection of less than 25 ppm of formaldehyde vapors, with a short response time of 20 seconds [124]. Zhang et al. [125] developed porous structured GO-WO3 hybrid fibers via the electrospinning technique. They found that 1 mL GO-WO3 hybrid nanofibers displayed the response of 35.9
100 ppm acetone at 375 C, which is 4.3 times higher than that of pristine WO3 nanofibers. The enhanced sensing efficiency for the hybrid electrode is due to the presence of GO, which has high surface area and good gas adsorption properties. Another new type of hybrid electrode was developed based on rGO-GaN nanorods through a spin coating process and used for the sensing of H2 and H2S [126]. It was found that the hybrid electrode showed a higher response to those gases compared with pure GaN. This hybrid electrode exhibited excellent sensing performance at room temperature under UVlight (λ 5 365 nm). Various graphene-based hybrid functional materials have been researched for the detection of various gases and chemicals.

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14.3.14 Surface Classification of Carbon Paste Electrode-Based Sensors Cyclic voltammetry (CV) and electrochemical impedance spectroscopic (EIS) studies were used to describe sensing characteristics. Entirely CV and EIS extents remained supported in KCl (0.1 M KCl) encompassing 5.0 mM of ferrocyanide. External extents of CPEs were assessed via CV. Randles–Sevcik equation is useful to determine the effect of scan rate on the peak current in the cyclic voltammetry [127,128]. The direct comparative studies are shown below:
ip 5 90:79v 1 2:25 R2 5 0:988 Conferring to the above equation, the slope of logip versus logarithmic graph can be adjacent to 1/2 if progression is diffusion controlled. The kinetics of electrochemical reaction can be calculated using [129]: ip 5

nFQv 4RT

From the Randles
Sevcik equation, it is possible to obtain an apparent diffusion coefficient [130]. It was observed that the peak potentials were changed under different pH conditions. The half-wave potentials and other electrochemical parameters can also be calculated [131].

14.4 CONCLUSION Graphene-based electrode materials were fabricated through different strategies, and used them as electrochemical sensors and biosensors. Different approaches have been discovered intended for the fabrication of different biosensors based on graphene or other nanocarbon resources as essentials for electrochemical sensing; however, there is still a need for further investigation. High surface area of graphene

prepared by CVD process can be used to detect small biomolecules (DNA) and metal ions. Graphene was functionalized with conducting polymers and different metal oxides and used as electrode materials for sensor applications. 3D structured graphene used excellent sensing material due to their biocompatibility, electrical and electrochemical properties. According to the research studies conducted in the literature, numerous graphene-based functional hybrid electrodes with various nanostructures are successfully fabricated through different techniques, with excellent properties such as electrical conducting and electrochemical behavior. Since such new hybrid materials showed excellent electrochemical performances due to the synergistic effect between the graphene sheets and modifier, they have the potential for designing and fabricating eco-friendly, large-scale devices for future electrochemical sensing applications. The development of highly efficient, graphene-based, hybrid electrodes is still in the early stage. For some of the specific electrochemical and biosensing applications, sensing devices are required to be developed with outstanding electrochemical characteristics and other properties (bonding between graphene and dopant species). We believe that it is possible to significantly improve the electrochemical properties of novel graphene-based functional hybrid electrodes by improving their large specific surface area, and effectively functionalizing graphene with nanoscale conductive materials such as silver 1D nanowires, copper nanorods, and polythiophene nanotubes. Such novel graphene-based hybrid materials with different hierarchical nanostructures would be promising for fabricating nextgeneration wearable sensing devices for the electrochemical detection of various chemicals (e.g., hydrogen peroxide, alcohols, drugs, toxic

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7145. [110] B. Han, Y.L. Zhang, L. Zhu, X.H. Chen, Z.C. Ma, X.L. Zhang, et al., Direct laser scribing of AgNPs@RGO biochip as a reusable SERS sensor for DNA detection, Sensor. Actuat. B 270 (2018) 500
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carbon nanotube composite electrode, Food Chem. 135 (2012) 362
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[121] C. Li, R. Wu, J. Zou, T. Zhang, S. Zhang, Z. Zhang, et al., MNPs@anionic MOFs/ERGO with the size selectivity for the electrochemical determination of H2O2 released from living cells, Biosensor. Bioelectron. 116 (2018) 81
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5498.

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Further Reading T.N.D. Alpha, C. Johann, N.P. Tim, B. Carsten, M. Thomas, Structure of epitaxial graphene on Ir (111), New J. Phys. 10 (2008) 043033. O. Chuhei, N. Ayato, Ultra-thin epitaxial films of graphite and hexagonal boron nitride on solid surfaces, J. Phys. Condens. Matter 9 (1997) 1.

L. Elena, C.B. Norman, J.F. Peter, F.M. Kevin, Evidence for graphene growth by C cluster attachment, New J. Phys. 10 (2008) 093026. C. Johann, T.N.D. Alpha, E. Martin, B. Carsten, W. Dirk, B. Niemma, et al., Growth of graphene on Ir (111), New J. Phys. 11 (2009) 023006.

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C H A P T E R

15 Challenges and Future Prospects of Nanoadvanced Sensing Technology Christina G. Siontorou1, Georgia-Paraskevi Nikoleli2, Marianna-Thalia Nikolelis3 and Dimitrios P. Nikolelis3,* 1

Laboratory of Simulation of Industrial Processes, Department of Industrial Management and Technology, School of Maritime and Industry, University of Piraeus, Piraeus, Greece, 2Laboratory of Inorganic and Analytical Chemistry, School of Chemical Engineering, Department 1, Chemical Sciences, National Technical University of Athens, Athens, Greece, 3Laboratory of Environmental Chemistry, Department of Chemistry, University of Athens, Athens, Greece

15.1 INTRODUCTION Nanotechnology deals with the generation and alteration of materials to nanosize (1029 m). Nanomaterials have dimensions in the range of 1
100 nm. The size of these materials makes them very special because they have most of their constituent atoms located at or near their surface and all their vital physicochemical properties differ greatly from those of their bulk materials. Nanomaterial-based biosensors represent the integration of material science, molecular engineering, chemistry, and biotechnology and have improved sensitivity and specificity for biomolecule detection and great potentiality for applications such as molecular recognition, food analysis,

environment monitoring, biomedical and clinical analysis, and pathogen diagnosis. The potential applications of nanoscale science and technology in food, medicine, and the environment are growing. However, as with all new technologies, these applications will require rigorous safety testing and risk/ benefit analyses to ensure that public and environmental concerns are addressed. Nanotools are recent routes in detection. Downsizing involves the miniaturization of devices and the ability to detect cellular and intracellular processes or moieties. Such capabilities provide nanosensors which are more suitable than conventional analytical systems with a wide range of applications that include clinical diagnostics, environmental monitoring, and quality control of industrial processes. During the past

* Corresponding authors: To whom correspondence should be addressed

Advanced Biosensors for Health Care Applications DOI: https://doi.org/10.1016/B978-0-12-815743-5.00015-9

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© 2019 Elsevier Inc. All rights reserved.

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decade, novel concepts have materialized into molecular electronics, synthetic biochemistry, DNA-based self-assembly, or manipulation of individual atoms via scanning tunneling microscopy. Opportunities, possibilities, and prospects are vast, yet there exist several obstacles that need to be overcome before routine sensing is revolutionized. Nanomaterial-based devices have several benefits in sensitivity and specificity over sensors made from traditional materials. Nanosensors can have increased specificity because they operate at a similar scale as natural biological processes, allowing functionalization with chemical and biological molecules, with recognition events that cause detectable physical changes. Nanosensors can also potentially be integrated with nanoelectronics to add native processing capability to the nanosensor. Sensors, in general, translate a chemical or biochemical interaction into an signal, for example, voltage and current. The sector is very dynamic, continuously evolving and well-established, almost in all continents, with remarkable infrastructure and potential [1]. To no surprise, most research teams were fast to adopt nanotools, processes, and concepts in order to solve their technology problems and optimize their product. Yet, size reduction to nanodimensions is neither straightforward nor problem-free. The fabrication of sensor subsystems or materials with nanometer sizes may be adequately addressed with nanomanufacturing strategies (e.g., self-assembly or 3D printing), but the efficient communication between these subsystems might prove to be tricky [2]. For example, as a critical distance between electrodes should be maintained, problems arise in integration and sensor architecture, circuitry for multilayer or 3D platforms may produce high noise levels. As the interface between the sensor and the environment decreases, the significance of certain phenomena, such as diffusion or dispersion, is

enhanced which may reduce the reliability of modeling. Nanosensing currently involves many research areas, of which the most important is the field of nanomaterial-based biosensors and chemical sensors. The use of nanomaterials has harmonized the scale between biological species and transduction platforms, thus resulting in the development of devices with higher rates of information flow [3]. The range of strategies, architectures, and materials for biomedical sensing, for example, increased drastically to include nonenzymatic catalysis schemes and enzyme wiring (Fig. 15.1). This chapter investigates the potentiality of nanosystems as a multidisciplinary science with the target of the design and development of smart-sensing technologies and devices that use micro/nanoelectrodes and novel nanosensing materials. It discusses their integration with miniaturized transduction systems and novel sensing strategies which perform, for example, diagnostics and personalized healthcare monitoring. Basic concepts are presented that pertain to nanobiosensor fabrication, developments in the field of smart nanomaterials, nanoenabling technologies, micro
nanohybrid platforms, and their applications in healthcare. In conclusion the challenges and future prospects of nanoadvanced sensing technology is presented and discussed.

15.2 NANOMATERIAL-BASED BIOSENSORS AND CHEMICAL SENSORS In the process to develop novel functional materials and devices with controlled and improved features on the nanometer scale is at the core of research and development innovation. The unique electronic, optical, and chemical properties of nanoscale materials make them attractive for the production of a new generation of devices. There is evidence that

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FIGURE 15.1

Nanosensor platforms for biomedical diagnosis. Traditional enzyme platforms have been optimized and nonenzymatic approaches have been developed, such as the direct adsorption of glucose onto graphene and its electron wiring through a Pt layer (top right panel). Peroxide detection follows a similar pattern, building an advantage over Fe3O4 (bottom left panel) or copper sulfide (bottom right panel) nanoparticles for enzymatic or nonenzymatic detection, respectively. Carbon platforms or polymer nanostructures can be converted to templates for detecting nitrites, sulfides, or heavy metals. Urea detection progresses along with chitosan chemistry, while cholesterol can be nanowired to graphene (top left panel). Adopted from C.G. Siontorou, V.T. Keramidas, G.-P. Nikoleli, D.P. Nikolelis, S. Karapetis, S. Bratakou, et al., Nano-enabled medical devices based on biosensing principles: technology basis and new concepts, AIMS Mat. Sci. 4 (2017), 250
266. Figure 7.

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15. CHALLENGES AND FUTURE PROSPECTS OF NANOADVANCED SENSING TECHNOLOGY

suggests novel behavior in nanomaterials that otherwise cannot be discovered by simple scaling laws. Semiconducting nanoclusters display interesting optical, electronic, and chemical properties which make them potential candidates for biosensors and optoelectronic nanodevices. While conductivity enhancement by several orders of magnitude occurs in some nanocrystalline oxides, there is also improved chemical and photochemical activity in other nanostructured oxides because they have enhanced surface areas. Because of unusual surface properties, materials with nanoscale features are particularly attractive for realizing fastresponding sensors with good sensitivity and selectivity for the detection of chemical species and biological compounds. Recent innovations in nanoprocessing that integrates cutting edge expertise and resources in materials processing, lithographic, and nonlithographic approaches in micro and nanofabrications, microscopy, and other advanced characterization techniques are described. Synthetic methodologies have been developed and produced fully customized materials according to their physicochemical properties, shape, and dimensions. A large class of chemical and biological sensors are based on the physicochemical characterization of interfaces. More specifically, electronic (bio)chemical sensing is often related to the characterization of interfaces between ion-based and electron-based conductive materials by means of electrical variables such as voltage, current, and charge. Also, recent improvements of integrated electronics and the development of nanotechnology have created a revolution in the field of biosensors allowing the shrinkage of very complex electronic systems into millimeter square sizes and this has prompted the development of novel nanosensing devices. This would allow implementing complex and sophisticated instrumentation in cheap and portable devices for the fast detection of harmful and toxic agents.

With the development of nanotechnology, many novel nanomaterials are being fabricated, their novel properties are being gradually discovered, and the applications of these materials in biosensors have also advanced greatly. For example, nanomaterial-based biosensors, which represent the integration of material science, molecular engineering, chemistry, and biotechnology, can markedly improve the sensitivity and specificity of biomolecule detection, and hold the capability of detecting or manipulating atoms and molecules. Nanobiosensors are based on merging nanotechnology with biosensors. Several nanomaterials have been investigated for the mechanism of their electronic and mechanical properties for their use in improved biological signaling and transduction mechanisms. Some of these materials include nanotubes, nanowires, nanorods, nanoparticles, and thin films made from crystalline matter. These can be as diverse as using amperometric devices for enzymatic detection of glucose to using quantum dots as fluorescence agents for the detection of binding and even using bioconjugated nanomaterials for specific biomolecular detection. The nanomaterials reported include colloidal nanoparticles which can be used to conjugate with antibodies for immunosensing and immunolabeling applications. These materials can also be used to enhance electron microscopic detections. Further, metal-based nanoparticles are excellent materials for electronic and optical applications that can be efficiently used for the detection of nucleic acid sequences through the exploitation of their optoelectronic properties. Various nanomaterials have been discussed to analyze their properties and applications in biosensors. The research in biosensor nanotechnology has prompted the construction of novel devices while interest is focused either in transducers or receptors, so as to enhance their multidetection capability, sensitivity, and selectivity. Nanomaterials contribute to either the

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15.2 NANOMATERIAL-BASED BIOSENSORS AND CHEMICAL SENSORS

biorecognition element or the transducer, or both. Nanosensors, nanoprobes, and other nanosystems have revolutionized in the fields of chemical and biological analysis, and enable the rapid analyses of multiple toxic and harmful substances and in many cases in vivo. In sensing, the most preferred platforms are still based on electrochemical or optical transduction [1]. Both formats are considered self-contained and low cost, while they are supported by a well-established research infrastructure, especially in universities. The main scope of research remains the development of reliable, highly sensitive and selective detectors with ultralow detectability and fast assay times [4].

TABLE 15.1 Type

With special interest to sensing, surface area is one property that significantly improves detection—the higher the surface area to volume ratio, the higher the rate of catalysis and response time and the lower the detection limit (DL) [5]. Further, optical, magnetic, and electrical properties are also improved and, more importantly, are amenable to control at the microscale [6]. The combination of functional nanomaterials, microfabrication techniques, and sensor integration technology yielded several electrochemical sensors and biosensors with nano, femtomolar, or even lower detection limits (Table 15.1). One-dimensional nanomaterials such as nanowires and nanotubes are well-suited for use

Nanomaterials Used in Biosensing and Chemical Sensing Platforms Properties/Advantages

Target Analyte

Detection Platform

DL

Refs.

Plasma S-nitrosothiol derivatives (SNOs)

Gold nanoparticles immobilized on ultramicroelectrode, in the presence of free thiols

100 nM

[7]

Hydroxylamine

Gold nanoparticles on metalmetalloporphyrin framework

4 nM

[8]

Adenosine triphosphate (ATP)

Gold nanoparticles/graphene oxide nanocomposites detected the binding of ATP to an aptamer

6.7 fM

[9]

Glycated hemoglobin (HbA1c)

Gold nanoparticles embedded Ndoped graphene nanosheets

0.2 μM

[10]

Hydrogen peroxide (H2O2)

Ni-doped Ag@C (Ni/Ag@C) nanocomposites

0.03 mM [11]

rGO-Nafion@Ag6 (rGO-Nf@Ag6) nanohybrid

0.5 μM

[12]

Moxifloxacin hydrochloride (MOXI)

CP modified with silver nanoparticles

2.9 nM

[13]

Glucose

Platinum decorated graphite and glucose oxidase

6.6 nM

[14]

Nitric oxide

Reduced graphene oxide
cobalt oxide nanocube@platinum nanocomposite

1.73 μM

[15]

METAL NANOMATERIALS Gold nanoparticles

Silver nanoparticles

Platinum nanoparticles

Adjustable physiochemical properties; high-surface area; high chemical stability; biocompatibility; wide electrochemical potential range

High conductivity; amplified electrochemical signal; biocompatibility

Distinctive electronic and electrocatalytic properties

(Continued)

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380 TABLE 15.1

15. CHALLENGES AND FUTURE PROSPECTS OF NANOADVANCED SENSING TECHNOLOGY

(Continued)

Type

Properties/Advantages

Target Analyte

Detection Platform

DL

Refs.

Palladium nanoparticles

Abundant material; amenable to sizeshape tailoring; unique electronic properties

Genosensor for Brucella

Potential-driven palladium nanoparticles deposited on a gold surface

27 zM

[16]

Levodopa (LD) and uric acid (UA)

Gold and palladium nanoparticlesmodified nanoporous stainless steel (Au-Pd/NPSS)

H2O2

Integration of single-walled carbon nanohorns with CeO2 (CeO2/ SWCNH) catalysts

Squamous cell carcinoma antigen (SCCA)

CeO2 nanocomposite (Co3O4@CeO2Au@Pt)

Copper oxide Various valence states; tunable nanomaterials electron-transport performance; hierarchical nanostructures; highsurface area

glucose

Nanoneedle-like CuO on N-doped rGO (CuO/N-rGO)

0.01 μM

[20]

Magnetic High accessible and active surface nanomaterials area; rapid electron-transfer; hierarchical porous structure

Methyl parathion

Methyl parathion hydrolase on the Fe3O4@Au nanocomposite

0.1 nM

[21]

Acetylcholine

Acetylcholine esterase and choline oxidase immobilized on the surface of Fe2O3 nanoparticles and poly(3,4ethylenedioxythiophene) (PEDOT)rGO nanocomposite

4 nM

[22]

Glucose

Genetically engineered M13@MnO2/ GOx nanowires

1.8 μM

[23]

Dopamine

CNTs-coated niobium (CNTs-Nb) microelectrode

11 nM

[24]

Ascorbic acid

Carbon nanotube fiber

259 μM

[25]

Glucose

Monodispersed Ni nanoparticles supported on functionalized MWCNT (Ni@f-MWCNT)

0.021 μM [26]

Acetaminophen

rGO

2.3 nM

[27]

Adenine dinucleotide Unscrewed Au nanoparticle/rGO (NADH) nanocomposite (Au NPs/rGO)

1.13 nM

[28]

Cholera toxin

Lipid films with incorporated ganglioside GM1 on graphene nanosheets

1 nM

[29]

Saxitoxin

Antisaxitoxin incorporated lipid films 1 nM on graphene nanosheets

[30]

[17]

METAL OXIDE NANOMATERIALS Cerium oxide Easy immobilization of proteinaceous nanomaterials moieties; supreme catalytic activities; low cost; high organic capture ability

0.1 mM

[18]

[19]

CARBON NANOMATERIALS Carbon nanotubes

Graphene

High surface-to-volume ratio; high electrical conductivity; chemical stability; biocompatibility; strong mechanical strength; easily tunable function

High sensitivity; great selectivity; good stability; low overpotential; wide potential window; negligible capacitive current; high electrocatalytic activity; transparency; excellent mechanical strength and flexibility; strong ambipolar electric field effect; good thermal and electrical conductivity

(Continued)

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15.2 NANOMATERIAL-BASED BIOSENSORS AND CHEMICAL SENSORS

TABLE 15.1 Type

(Continued) Properties/Advantages

Target Analyte

Detection Platform

DL

Refs.

DNA

Nanocomposite consisting of MWCNTs coated with polypyrrole (PPy) and redox poly-(amidoamine) dendrimers (PAMAM) (MWCNTs-PPy-PAMAM)

0.3 fM

[31]

Paracetamol

Nanocomposite consisting of MWCNTs coated with polypyrrole (PPy) and redox poly-(amidoamine) dendrimers (PAMAM) (MWCNTsPPy-PAMAM)

0.1 μM

[32]

Dopamine

AuNPs patterned on polyaniline nanowires (PANI)

0.08 μM

[33]

Ascorbic acid

AuNPs patterned on polyaniline nanowires (PANI)

0.01 μM

[33]

Serotonin

AuNPs patterned on polyaniline nanowires (PANI)

0.025 μM [33]

Uric acid

AuNPs patterned on polyaniline nanowires (PANI)

0.04 μM

[33]

Paracetamol

Macromolecular self-assembly and molecular imprinting technique employing paracetamol as a template molecule

0.3 μM

[34]

Creatinine

Magnetic-MIPs and a nanocomposite consisting of Ni NPs and PANI (Ni@PANI NPs)

0.2 nM

[35]

Thrombin

CNTs, aptamer, and horseradish peroxidase

0.05 pM

[36]

Cardiac biomarkers

Gold microelectrode functionalized with aptamers

Melamine

DNA on an indium tin oxide (ITO) electrode

Mucin 1 tumor marker

Dual signal-tagged hairpin structured DNA(dhDNA)-based ratiometric probe

POLYMER NANOMATERIALS Dendrimers

Conducting polymers

Molecularly imprinted polymers

Structural consistency; veracity; wellordered composition; biocompatibility

Unique electronic properties

Tailored architectures

BIOLOGICAL MATERIALS Aptamers

DNA

High affinity; stability

High selectivity; low cost; simple synthesis; reusability; high affinity; flexibility

[37] 0.43 nM

[38] [39]

CNTs, carbon nanotubes; CP, carbon paste; DL, detection limit; fM, femtomolar (10
15 M); MIP, molecularly imprinted polymer; mM, millimolar (1023 M); MWCNT, multiwalled carbon nanotubes; nM, nanomolar (10
9 M); pM, picomolar (10
12 M); rGO, reduced graphene oxide; zM, zeptomolar (10
21 M); μM, micromolar (10
6 M).

in nanosensors compared to bulk or thin-film planar devices. They can function both as transducers and wires to transmit the signal. Their high-surface area can cause large signal changes

upon binding of an analyte. Their small size can enable extensive multiplexing of individually addressable sensor units in a small device. Their operation is also “label-free” in the sense of not

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15. CHALLENGES AND FUTURE PROSPECTS OF NANOADVANCED SENSING TECHNOLOGY

requiring fluorescent or radioactive labels on the analytes. There are several challenges for nanosensors still to be explored, including avoiding fouling and drift, developing reproducible calibration methods, and applying preconcentration and separation methods to attain a proper analyte concentration to be detected and also interferences from other sample constituents. To date, materials science has reached a high degree of sophistication. As a result of continuous progress in synthesizing and controlling materials on the submicron and nanometer scales, novel advanced functional materials with tailored properties can be created. The boundary between materials science and biology has become a fertile ground for new scientific and technological developments. For the fabrication of an efficient biosensor, the selection of the substrate for dispersing the sensing material determines the sensor performance. Various kinds of nanomaterials, such as gold nanoparticles, carbon nanotubes (CNTs), graphene, magnetic nanoparticles, and quantum dots are being gradually applied to biosensors because of their unique physical, chemical, mechanical, magnetic, and optical properties and considerably enhance the sensitivity and specificity of detection. Potential applications for nanosensors include medicine, detection of contaminants and pathogens in the field or workplace, the environment, and in products such as food, as well as monitoring manufacturing processes and equipment, Medicinal uses of nanosensors mainly involve the potentiality of nanosensors to accurately identify particular cells or places in the body. By measuring changes in volume, concentration, displacement and velocity, gravitational, electrical, and magnetic forces, pressure, or temperature of cells in a body, nanosensors may be able to distinguish between or recognize certain cells, most importantly those of cancer, at the molecular level in order to deliver medicine or monitor development to specific places in the body.

15.3 NANOSTRUCTURES, NANOPARTICLES, NANOWIRES, NANOFIBERS, AND NANOPROBES A nanostructure is a structure of intermediate size between microscopic and molecular structures. In describing nanostructures, it is necessary to differentiate between the number of dimensions in the volume of an object that are on the nanoscale. Nanotextured surfaces have one dimension on the nanoscale, that is, only the thickness of the surface of an object is between 0.1 and 100 nm. Nanotubes have two dimensions on the nanoscale, that is, the diameter of the tube is between 0.1 and 100 nm. Finally, spherical nanoparticles have three dimensions on the nanoscale, that is, the particle is between 0.1 and 100 nm in each spatial dimension. The terms nanoparticles and ultrafine particles (UFP) are often used synonymously although UFP can reach into the micrometer range. Novel nanomaterials for use in bioassay applications represent a rapidly advancing field of science. Various nanostructures have been explored to define their properties and possible applications in biosensors. These structures include nanotubes, nanofibers, nanorods, nanoparticles, and thin films. Nanoparticles have numerous possible applications in biosensors. For example, functional nanoparticles (electronic, optical, and magnetic) bound to biological molecules (e.g., peptides, proteins, and nucleic acids) have been developed for use in biosensors to detect and amplify various signals. Some of the nanoparticle-based sensors include acoustic wave biosensors, optical biosensors, and electrochemical biosensors, as discussed next. Cui reported boron-doped silicon nanowires (SiNWs) to create highly sensitive, real-time electrically based sensors for biological and chemical species [40]. The amine and oxidefunctionalized SiNWs exhibited pH-dependent

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15.5 METAL NANOMATERIALS

conductance that was linear over a large dynamic range and can be understood in terms of the change in surface charge during protonation and deprotonation. Biotin-modified SiNWs were used to detect streptavidin down to at least a picomolar concentration range. In addition, antigen-functionalized SiNWs showed reversible antibody binding and concentrationdependent detection in real time. The small size and capability of these semiconductor nanowires for sensitive, label-free, real-time detection of a wide range of chemical and biological species can be exploited in array-based screening and in vivo diagnostics. The nanoscale size of this new class of sensors allows for measurements in the smallest of environments such as individual cells. This provides opportunities for in vivo monitoring of processes within live cells. Cullum et al. (2000) used optical fibers with a distal-end diameter of less than 1 μm, coated with antibodies, to detect the presence of toxic chemicals within single cells. They were able to measure the concentration of benzopyrene tetrol within human mammary carcinoma cells and rat liver epithelial cells. Vo-Dinh fabricated nanoprobes with optical fibers pulled down to tips with the distal ends having sizes of approximately 30 2 50 nm (2002). Using these nanobiosensors, it has become possible to probe chemical species at specific spots. Nanocontrolled release systems have been devised for optical biosensing of peroxide concentration (Choi et al., 2001).

15.4 TUBULAR AND POROUS NANOSTRUCTURES A common use of tubular and other porous nanostructures in biosensors is to increase the quantity and activity of the immobilized biomolecules. However, in view of their unique properties, these nanostructures provide opportunities for the development of novel designs of biosensors.

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15.5 METAL NANOMATERIALS Metallic nanoparticles have fascinated scientists for over a century and are now heavily utilized in biomedical sciences and engineering. They are a focus of interest because of their huge potential in nanotechnology. Today these materials can be synthesized and modified with various chemical functional groups which permits them to be conjugated with enzymes, antibodies, receptors, ligands, and drugs of interest and, thus open a wide range of potential applications in biotechnology, targeted drug delivery, vehicles for genes, and more importantly diagnostic imaging. Moreover, various imaging modalities have been developed over time such as magnetic resonance imaging, ultrasound, surface-enhanced Raman spectroscopy, and optical imaging as aids to provide images of various disease states. These imaging modalities differ in terms of techniques and instrumentation and, more importantly, require a contrast agent with unique physiochemical properties. This has led to the discovery of nanoparticles, that is, magnetic (Fe3O4), gold (Au), and silver (Ag) nanoparticles. This review aims to provide an introduction to these nanoparticles by discussing their synthesis, physiochemical properties, and some recent applications in the diagnostic imaging and therapy of cancer. AuNPs demonstrate a high chemical stability and adjustable physical chemistry besides their wide electrochemical potential range (Table 15.1). Various sensors have been developed with AuNPs, used either alone [7], or in combination with metalloporphyrins [8], aptamers [9], or graphene nanomaterials [10]. AgNPs exhibit higher conductivity and are more amenable to doping with other metals [11] or mixing with carbonbased materials [12,13]. Platinum (Pt) is easy to decorate and more suitable for proteinaceousbased composites [14]. Palladium (Pd) is an abundant material amenable to size or shape tailoring [41] and PdNPs have been extensively

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FIGURE 15.2 Overview of the methods commonly used for the fabrication of nanoparticles.

used for the development of genosensors (see e.g. [16]) or for clinical biomarkers such as levodopa and uric acid [17]. There are generally two approaches for the preparation of nanoparticles: top
down processes involve the reduction of bulk materials into nanoparticles, while bottom
up processes build-up the particles starting from molecules or atoms (Fig. 15.2). Depending on the application, the selection of the synthetic route is very important in order to avoid polydispersity and wide size distributions [42]. Dispersion at the bulk-scale is a mechanical process, involving, mainly, the coagulation of the particles and their spatial distribution into

the dispersed phase. At the nanoscale this is not exclusively mechanical; as the size of the particles decreases, the significance of the physicochemical interactions increases. A tradeoff should be carefully considered between coagulation and dispersion, the former should be kept to a minimum or avoided entirely, if possible, whereas the latter should be controlled with mixing [2]. In that way, however, critical phenomena for the orientation and the size distribution of the particles are underplayed. Coagulation leads to the decrease of the entropy of the system that is presented as elasticity upon shear deformation, which, in turn, forces the particles to orient themselves along

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15.6 METAL OXIDE NANOMATERIALS

the axis of the shear stress. At mixing, the elastic shear strain falls to zero, thus the entropy increase due to the disorientation of the particles leads to wide size distributions. Chemical methods (e.g., deposition, reduction, or thermal decomposition) are generally more complex and energy-demanding since the reaction should occur far from equilibrium so that the nucleation rate remains high [43]. Low-temperature fabrication yields particles at micrometers sizes, requiring further processing (e.g., with arc discharge) for improving quality. Efficiency, however, remains low as the products have high impurity contents.

15.6 METAL OXIDE NANOMATERIALS Metal oxides (Table 15.1) present better catalytic activities at lower costs, can be fabricated into hierarchical nanostructures, offer a range of valence states, have tunable performance, and exhibit a better compatibility to proteinaceous moieties [44]. Iron oxide nanoparticles are more compatible with enzymatic platforms, while copper sulfide crystal nanostructures work better in nonenzymatic schemes [3]. Iron oxide nanoparticles have been used for the harvesting and separation of biological moieties or as drug delivery vehicles [3]. Bioconjugation proceeds via amino groups, polymer entrapment, or cross-linking. Cerium and copper oxides are wellintegrated with graphene nanomaterials, and such hybrid platforms have been used for the detection of hydrogen peroxide [18], cancer biomarkers [19], or glucose [20]. Magnetic nanoparticles present a significant advantage for biomedical applications as they retain their properties even when their size reduction goes below critical dimensions [45]. They exhibit an enhanced magnetic moment that affords them almost paramagnetic properties. The exploitation of the magnetic properties

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of metal oxides, especially of ferrous or manganese alloys, facilitates the coimmobilization of more than one biological species. For example, Chauhan et al. [22] have immobilized acetylcholine esterase and choline oxidase on the surface of Fe2O3/reduced graphene oxide nanocomposites for the development of an acetylcholine sensor. Han et al. [23] fabricated easily modified, one-dimensional manganese oxide nanowires using genetically engineered phages as nucleation templates. The developed platform could monitor the electrooxidation of hydrogen peroxide at neutral pH and used for the nonenzymatic detection of glucose. Magnetic nanomaterials offer alternative analytical capabilities for the detection of lipophilic substances. Zhao et al. [21] developed a Fe3O4@Au nanocomposite biosensor platform for the detection of organophosphorus pesticides. Using methyl parathion hydrolase, for example, methyl parathion could be detected at 0.1 nM levels. A critical problem with magnetic nanoparticles is their synthetic routes. Fig. 15.3 shows an overview of the common methods for the preparation of ferrous/ferric-based nanoparticles. The synthetic pathways are generally simple and fast, yet control over shapes and size distribution is not optimized yet, and yields remain low [46]. Particle size and size distributions are very important for the magnetic behavior of the particles. At the microscale, several domains of homogeneous magnetization can be formed. At the nanoscale, energetic parameters prevent the formation of these domains and the particle exhibits spontaneous magnetization [47]. The production of magnetic nanoparticles with satisfactory magnetic behavior has another obstacle to address, that is, the steric stabilization of the dispersions is challenging due to their tendency to form agglomerates. Despite this, the implementation of these particles in sensor development impacts adversely the reliability of detection unless their size and size distribution could be accurately adjusted.

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FIGURE 15.3 A comparative schematic of the advantages and disadvantages of the common methods for the preparation ferrous/ferric-based nanoparticles.

15.7 CARBON-BASED NANOMATERIALS The era of carbon nanomaterials was initiated with the first reports on fullerenes and related compounds in the mid-1980s. A tremendous increase of research in this field has been observed ever since. New classes of carbon materials have been produced which include

carbon nanotubes, carbon onions, and nanoscale diamonds and diamondoids. The appearance of graphene as a nanomaterial available for investigation was spurred by the development of reliable production methods. The progress in terms of understanding the properties and chemistry of carbon nanomaterials has opened a whole new dimension of applications for nanomaterials.

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15.7 CARBON-BASED NANOMATERIALS

By now, all types of carbon nanomaterials are available to the scientific community with excellent quality and suitable amounts for the investigation of fundamental properties and prospective applications. This has led to the emergence of a new community of scientists working in an interdisciplinary area involving materials science, organic chemistry, and physics. Carbon can be found in several different hybridization states, each having unique properties as shown in Fig. 15.4 [48]. In fact, the electrical, thermal, mechanical, and chemical properties of the different allotrope forms are directly correlated to their hybridization state and structure, opening up the possibility to use the same material for a wide range of applications [48]. Various techniques have been reported in the literature for the synthesis of 0D, 1D, 2D, and 3D carbon nanomaterials. The most common techniques are laser ablation [49
52], arc-discharge [52,53], and chemical vapor deposition (CVD) [54,55]. CVD is the most commonly employed thin-film deposition technique used to synthesize nanomaterials. The various forms of carbon-based nanomaterials, that is, single-walled or multiwalled carbon nanotubes, carbon nanohorns, fullerenes, buckypaper, or graphene are provided in Table 15.1. Carbon-based nanoconstructs have

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been converted into sensors for anions or heavy metals [3]. The high surface-to-volume ratio and their mechanical strength are considered the most prominent advantages for sensor development (for a recent comprehensive review, see [56]). Carbon nanotubes exhibit high and tunable electrical potential for the selective detection of electroactive species, such as dopamine [24], ascorbic acid [25], or glucose [26]. The low overpotential and capacitive current, combined with the wide potential window of graphene have been exploited for lowering detection limits to less than nanomolar levels [27,28]. Also, the flexibility and mechanical strength of graphene nanosheets have advanced considerably the ruggedness of lipid-based biosensor platforms and enabled the immobilization of complex receptors [29] and engineered antibodies [30] for the detection of toxins. Graphene is a one-atom thick layer of carbon and is considered to be a new wonder molecule. Its production became possible only in 2018 and graphene is now available for various applications. The term graphene is often applied to many members of the family of graphene-based materials, the two most important members being graphene and graphene oxide (GO). Graphene is transparent, flexible, and very stable on a molecular level. Various future uses of graphene and graphene

FIGURE 15.4 Hybridization states of carbon-based nanomaterials. Reproduced with permission from E.B. Bahadir, M.K. Sezgintu¨rk, Applications of commercial biosensors in clinical, food, environmental, and biothreat/biowarfare analyses. Anal. Biochem. 478 (2015), 107
120. Copyright (2008) American Chemical Society.

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15. CHALLENGES AND FUTURE PROSPECTS OF NANOADVANCED SENSING TECHNOLOGY

FIGURE 15.5 Overview of graphene and carbon nanotube syntheses.

oxide are expected, from applications in the fields of electronics, photonics, composite materials, and energy generation and storage, sensors, and metrology to uses in biomedicine. An obvious disadvantage of graphenebased nanostructures is the nonscalable costeffective synthesis (Fig. 15.5). For experimentation, easy and simple methods such as mechanical exfoliation may be adequate, yet for mass production other routes should be developed. Other issues have also attracted much attention lately, such as the problems of elastic modulus and tensile strength of carbon nanoconstructs in composites with polymers or the dependence of the electrical percolation thresholds on morphology, purity, and aspect ratio [57].

15.8 POLYMER NANOMATERIALS Polymer nanomaterials (Table 15.1) can be fabricated to sizes, shapes, and properties fitted for the immobilization of a variety of biochemical species [44]. An important aspect of this technology is the interface between the abiotic substrate and the biotic layer. Biofunctionalization depends on the nature of the biological moieties and the extent of their exposed surface. Ruggedness depends on the efficiency of polymer-bioelement binding, and signal transduction depends on how the biochemical interaction affects the properties of the nanomaterials. Dendritic structures with built-in amine groups have the ability to covalently bind to many bioelements. Thus detection

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15.8 POLYMER NANOMATERIALS

limits may be reduced to a few femtomoles [31]. The construction of the intended architecture may be quite complicated, but the suitability of these sensors for the detection of pharmaceuticals has been demonstrated [32]. Conducting polymers have been widely used for biosensor development [1], thus the experience gained during the past few decades is vast. Nanosynthesis has enabled the fabrication of the polymeric matrices into nanowires, reducing the detection limit for various electroactive species manifold [33]. A molecularly imprinted polymer (MIP) is a polymer that has been constructed using the molecular imprinting technique. This route leaves cavities in the polymer matrix with an affinity for a chosen “template” molecule. The procedure involves initiating the polymerization of monomers in the presence of a template molecule that is extracted afterwards, leaving behind the complementary cavities. These polymers have affinity for the original molecule and have been used in applications such as chemical separations, catalysis, or molecular sensors. Molecular imprinting is, in reality, an artificial tiny lock for a specific molecule that serves as a miniature key. The imprinted polymer has great selectivity and grabs only specific chemicals. Many basic biological processes, from sensing of odors to signaling between nerve and muscle cells, rely on such lock-and-key combinations. Presently the elegance of molecular imprinting in nature has inspired many scientists to build the locks themselves by etching a material to create specific cavities which in size, shape and functional groups, fit the target molecule. Molecular design has prompted a revolution in this field of science. The stability, flexibility, and other properties are freely modulated according to need. Even functional terminal groups which are not found in nature can be employed in these artificial compounds.

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Furthermore, the activity in response toward outer stimuli (photo-irradiation, pH change, electric or magnetic field, etc.) can be provided by using appropriate functional groups. The spectrum of functions is broader than those of naturally occurring compounds. In a molecular imprinting processes, one needs a: (1) template; (2) functional monomer(s); (3) cross-linker; (4) initiator; (5) porogenic solvent; and (6) extraction solvent. There are two main methods for the fabrication of MIPs. The first is called self-assembly and involves the formation of the polymer by combining all the elements of the MIP and allowing the molecular interactions to form the cross-linked polymer with the template molecule bound. The second method of formation of MIPs involves covalent linking of the imprint molecule to the monomer. After polymerization, the monomer is cleaved from the template molecule. The selectivity is greatly influenced by the kind and amount of crosslinking agent used in the synthesis of the imprinted polymer. The selectivity is also determined by the covalent and noncovalent interactions between the target molecule and monomer functional groups. The careful choice of functional monomer is another important parameter to provide complementary interactions with the template and substrates. The cross-linker is important in controlling the morphology of the polymer matrix, whether it is gel-type, macroporous, or a microgel powder; the cross-linker also serves to stabilize the imprinted binding site. High cross-linking is generally preferred in order to obtain permanent porous materials and in order to generate materials with adequate mechanical stability. The self-assembly method offers a few advantages as it forms a more-natural binding site, and also offers additional flexibility in the types of monomers that can be polymerized. The covalent method has advantages in offering a high yield of homogeneous binding sites,

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15. CHALLENGES AND FUTURE PROSPECTS OF NANOADVANCED SENSING TECHNOLOGY

but first requires the synthesis of a derivatized imprint molecule and may not imitate the “natural” conditions that could be present elsewhere. Over recent years, interest in the technique of molecular imprinting has increased rapidly, both in the academic community and in the industry. Significant progress has been achieved in the development of polymerization methods that produce adequate MIP formats with relatively good binding properties expecting an enhancement in the performance or in order to suit the desirable final application, such as beads, films, or nanoparticles. An issue that has limited the use of MIPs in practical applications so far is the lack of simple and robust methods to synthesize MIPs in the optimum formats required by the applications. Most investigators in this field are making MIPs with heuristic techniques such as the hierarchical imprinting method. The technique for the first time was used for making MIP [58] for imprinting small target molecules. With the same concept, a technique was developed [59], called the polymerization packed bed, to obtain a hierarchically structured, highcapacity, protein-imprinted porous polymer beads by using silica porous particles for protein recognition and capture. Molecularly imprinted polymers have the advantage of selective organization of functional groups in controlled architectures [43]. They have been extensively proposed for the detection of pharmaceuticals (see, e.g. [34]) or disease markers (see, e.g. [35]). There are generally two methods for the preparation of molecularly imprinted polymers, namely, bulk polymerization and grafting, and both have many variations (see [60] for a recent review). Both methods are complicated, requiring accurate optimization of imprinting conditions and high volumes of solvents that increase costs considerably. Despite this, yields are generally very low, whereas the thickness of the deposited film cannot be

controlled as do the shape and the distribution of the pores [61]. A technique for the construction of polymer nanomaterials with incorporated lipid film similar to molecularly imprinting technique, but much more simplified, has been described in the literature (Nikolelis and Mitrokotsa, 1992). Microporous filters composed of glass fibers Gf/F (nominal pore sizes 0.7 and 1.0 μm) were used as supports for the formation and stabilization of these polymer nanomaterials. Methacrylic acid was the functional monomer, ethylene glycol dimethacrylate was the cross-linker, and 2,20 -azobis-(2-methylpropionitrile) was the initiator. The mixture was sparged with nitrogen for about 1 min and sonicated for 30 min. This mixture could be stored in the refrigerator. For the preparation of the polymer films, 0.15 mL of this mixture was spread on the microfilter and was left at 60 C for 12 h. This polymerization technique is described in Fig. 15.6. A technique for the construction of these nanosensors took place by using UV irradiation instead of heating the polymerization mixture at 60 C [62,63]. This process retains the activity of a “receptor” (i.e., enzyme, antibody, natural or artificial receptor), whereas heating may deactivate it. The results indicated that the polymerization is completed within 4 hours by using physicochemical methods such as differential scanning calorimetry (DSC), IR, or Raman spectrophotometry. The preparation of these devices makes possible the practical use of these nanodevices based on lipid membranes to perform chemical sensing, because it allows incorporation of a “receptor” molecule such as an enzyme, antibody, or ion-channel receptor without their deactivation and, more importantly, these devices are stable outside of the solution in the air for more than 48 hours. Various efforts have been made to circumvent these disadvantages and improve the

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15.8 POLYMER NANOMATERIALS

O

R = fatty acids

O O

R

OH

R O

O

OO

N+(CH3)3

P

O O

O

HO

Phosphatidylcholine

Methacrylic acid Prearrangement

O O OH

O

O

R

R

O-

O

O

– O

O

P O

O

N+(CH3)3

O

HO

O O

Polymerization

+ AIBN O

O Ethylene glycol dimethacryate O O OH

O

O O

R

R O

O P

O

O

–

O-

N+(CH

O 3)3

O

OH

FIGURE 15.6 A schematic of the polymerization stage and preparation of polymerized lipid membranes.

performance of sensors based on polymeric nanoplatforms. Fig. 15.7 presents an overview of the main strategies used, as well as their

benefits and problems. These three strategies include solid-phase extraction, chemical sensors and artificial antibodies owing to their

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FIGURE 15.7

Overview of the strategies currently adopted for optimizing nanopolymer-based sensing.

REFERENCES

desired selectivity, physical robustness, thermal stability, as well as low cost and easy preparation. Currently, there are three strategies involving signal amplification by labeling or doping with functional materials that are directly applied on the polymer network or the coupling of the polymer platform with amplification mechanisms [64]. Certain problems in this technology have been also stressed, including the interaction of nanopolymers with proteins or macromolecules, low reproducibility, steric hindrance from labeling, and less-than-optimal signal-to-noise ratios (see e.g., [65]).

15.9 BIOLOGICAL MATERIALS Nanosynthesis has facilitated the integration of biological materials with nanoplatforms and are particularly suitable for healthcare applications (Table 15.1). Engineered materials, such as aptasensors or DNA microstructures, can be easily incorporated within a nanomaterial frame to provide detection for thrombin [36], creatinine kinase [37], or tumor markers [38,39], for example. Their integration with other nanoparticles remains challenging due to the denaturation of biological species under harsh operation conditions and the impact of bioelements on the charge and density of the nanomaterial network [3]. For biomedical applications, the coating of biolelements with inorganic nanoparticles could reduce nonspecific protein adsorption. However, immunogenicity has not been studied yet [39].

15.10 CONCLUSION Looking into the framework of the relevant research, the multidisciplinary developed becomes evident. At present, the synthesis of nanomaterials is complicated, care-intensive, condition-sensitive, expertise-demanding, and

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costly, while yields are low and processes are not scalable. Once the fabrication issues are satisfactorily solved, issues regarding the analytical performance of the sensors are expected to minimize as most of them are directly linked with the fabrication of nanoplatforms and the assembly of the devices. Despite this, opportunities are vast, especially with regard to sensing architectures. Efforts should be focused toward optimizing sensitivity and selectivity to allow for multianalyses using bimetallic or trimetallic nanoparticles or multilayered nanocomposites. Enzyme mimics and artificial biomaterials present substantial potential for future applications. At present, knowledge on cellular functions is limited, let alone their full synthetic replication. Still, the ability to design biological moieties may advance sensing and support personalized medicine, point-of-care monitoring, and disease prevention and control.

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1571.

ADVANCED BIOSENSORS FOR HEALTH CARE APPLICATIONS

Index

Note: Page numbers followed by “f” and “t” refer to figures and tables, respectively.

A AA. See Ascorbic acid (AA) AA-p. See Ascorbic acid 2-phosphate (AA-p) Ab. See Antibodies (Ab) Ab
Ag interaction, 149 ABTS. See 2,20 -Azino-bis(3ethylbenzothiazoline-6-sulfonic acid) (ABTS) Acetone sensing for diabetes diagnosis, 218
219 Acetylcholinesterase (AChE), 288
289, 302
303 Acinetobacter baumannii, 273
274 Acquired immunodeficiency syndrome (AIDS), 270
271 Acute inflammatory response, 250 Adenosine triphosphate (ATP), 282
283 Adsorption, 250 Advanced hybrid materials, 171 Advanced inorganic hybrid materials, 172
176 alloy and metal oxide hybrids, 174
175 colloidal clusters, 173 mesoporous-silica hybrid materials, 176 self-assembled inorganic nanorods, 175
176 titanium-oxo clusters, 173
174 Advanced nanoparticle-based biosensors bioreceptor types, 6
10 established and traditional diagnostic tools, 2
4 foodborne pathogens, 1
2 membrane-based biosensors, 22
26 multiplex biosensors, 26
29

NPs and biosensors, 4
5 performance of nanoparticle-based biosensors, 29
33 transducer types, 10
22 Advanced organic hybrid materials, 188
190 Advanced organic
inorganic hybrid materials, 176
188 carbon
organic hybrid materials, 177
179 functionalized clays and silica, 186
187 ionic liquid hybrid materials, 187
188 magnetic materials, 182
186 metal nanoparticles, 179
182 AEGIS-SELEX. See Artificially expanded genetic information systems-SELEX (AEGIS-SELEX) Affinity sensors, 229 Aflatoxin B1 (AFB1), 290 AFP. See Alpha fetoproteins (AFP) AFP antibody (Ab1(AFP)), 293
294 AIDS. See Acquired immunodeficiency syndrome (AIDS) Albumin nanoparticles (AlbNPs), 192
193 Alcohol consumption, 72 Alcohol oxidase (AOx), 214 Alcohols detection, 213
214 ethanol, 214 Alexa Flour 488 (dye), 290
291 Alkaline phosphatase (ALP), 305 Allergic reaction, 269 Alloy, 172
175 ALP. See Alkaline phosphatase (ALP) Alpha fetoproteins (AFP), 191
192, 282
283, 293
294

397

ALVs-J. See Subgroup J of avian leukosis viruses (ALVs-J) ALX-0171 nanobody, 270 American Cancer Society, 147
148 Amino propyl triethoxysilane, 185
186 Amino-modified tail-clamp bioreceptor, 233
234, 234f 2-Aminobenzenethiol, 331
332 4-Aminophenol (4-AP), 286 3-Aminophenyl boronic acid, 329
330 (3-Aminopropyl)triethoxysilane (APTES), 19, 110
111 Ammonium ion, 214
216 Ammonium-selective membrane, 214
216 Amperometric biosensors, 10
12, 85, 235
236 cholesterol biosensor, 329
330 detection, 25
26 i
t technique, 328
329 method, 62 sensors, 301
303 Angiogenic drugs, 251
252 Anodic stripping voltammetry technique (ASV technique), 110
111 Anodic SWV, 135 Antagonistic anti-EGFR nanobodies, 266
267 Anti-CEA Ab, 160
162 Anti-von Willebrand factor humanized nanobody, 268
269 Antibiotic resistance, 274 Antibodies (Ab), 6
7, 33, 80
81, 81f, 149 antibody-dependent detection procedures, 50
52 aptamers over, 230
232, 231t

398 Antifluorescein-conjugated alkaline phosphatase, 7
8 Antigen (Ag), 149 antigen-binding site, 263 antigen-functionalized SiNWs, 382
383 antigen-specific nanobodies, 264
266 Antiinflammatory drugs, 251
252 Antioxidants, 105 biosensors for polyphenols analysis in beverages, 120
135 for ROS determination, 108
115 classification of major antioxidants, 107f electrochemical biosensors, 115
120 mechanism of action, 106 Antiviral drugs, 270
271 AOAC. See Association of Official Analytical Chemists (AOAC) AOx. See Alcohol oxidase (AOx) 4-AP. See 4-Aminophenol (4-AP) APE-1. See Apurinic/apyrimidinic endonuclease 1 (APE-1) Apple stem hole virus (ASPV), 57
58 Aptamer, 8
9, 33, 230, 379t bioreceptor element, 230 health monitoring, 241 merits over antibodies, 230
232, 231t technology, 55. See also Transducer—technology biosensors, 53
56 classification of aptamer-based biosensors, 56
64 foodborne pathogens, 46, 47t foodborne pathogens detection, 50
53 future prospects of biosensors, 64 NoV, 46
50 selection, 55
56 Aptamer Selection Express SELEX (ASExp SELEX), 56 Aptamer-based biosensors, 56
64, 393 efficacy in combination with nanomaterials, 232 electrochemical aptasensors, 62
63 impedimetric detection, 62 optical biosensors, 57
62 Aptamer-based LFIAs, 24
25

INDEX

Aptamer-mediated nanobiosensing for health monitoring biosensors, 228
232 causes of death in United States, 228f types of nanobiosensors, 232
244 Aptasensors. See Aptamer-based biosensors Aptazymes, 82
83 APTES. See (3-Aminopropyl) triethoxysilane (APTES) Apurinic/apyrimidinic endonuclease 1 (APE-1), 163
165 Artificially expanded genetic information systems-SELEX (AEGIS-SELEX), 56 Artificially intelligent nanosensors for multiple disease detection, 221
222 Ascorbic acid (AA), 135, 164
165, 286, 296, 366
367 detection, 366
367 Ascorbic acid 2-phosphate (AA-p), 164
165 ASExp SELEX. See Aptamer Selection Express SELEX (ASExp SELEX) ASPV. See Apple stem hole virus (ASPV) Association of Official Analytical Chemists (AOAC), 13 Asthma, 249 diagnosis, 220
221 ASV technique. See Anodic stripping voltammetry technique (ASV technique) Ataxia-Telangiectasia Mutated (ATM), 72
73 ATP. See Adenosine triphosphate (ATP); p-aminothiophenol (ATP) Au electrode modified by cysteine (Au/Cys), 112
113 Au-f-MWCNT-PPy-ChOx-GCE-based biosensor, fabrication of, 321
322, 321f Au-graphene@PEDOT ternary electrode, 367 Au/TGA/ChOx fabricated biochip sensor, 322 AuNCs. See Gold nanoclustures (AuNCs) AuNFs. See Gold nanoflowers (AuNFs)

AuNPs. See Gold nanoparticles (AuNPs) Avenanthramides (AVs), 115 AVs. See Avenanthramides (AVs) 2,20 -Azino-bis(3-ethylbenzothiazoline6-sulfonic acid) (ABTS), 121 Azo-bis-isobutyronitrile, 185
186 2,20 -Azobis-(2-methylpropionitrile), 390

B B-cell hybridomas, 80
81 Bacillus cereus, 14 Backscattering interferometry (BI), 235 Bacteria, nanobodies against, 272
274 Bacteria detection, microfluidics chips in, 285 Bacterial expression system, 263 Bacteriophages, 17 Bad breath. See Halitosis BCO. See Brevibacterium sterolicum (BCO) Beer, 121 BEIA. See Bioluminescent enzyme immunoassay (BEIA) β-cyclodextrin-tagged glucose oxidase, 194 β-galactosidase (Gal), 156 3β-hydroxysteroid, 318 β2 microglobulin (β2 M), 345
346, 348f Beverages, biosensors for polyphenols analysis in, 120
135 antioxidant capacity for healthcare purposes, 120
124 electrochemical biosensing of polyphenols, 125
135 BI. See Backscattering interferometry (BI) Bioanalytical applications, 175 Biobarcode platinum nanoparticles-Gquadruplex/heme, 293 Biochemical receptor, 172 Biocompatibility, 250
252, 254 Biodegradability, 254 Biodistribution studies, 267 Biofouling, 250 Biofunctionalization, 388
389 Biohybrids, 186
187 Biological damage, 132 materials, 393 MEMS, 254
255 processes, 218 recognition elements, 279

INDEX

Biological microelectromechanical systems (bioMEMS), 254
255 Bioluminescent enzyme immunoassay (BEIA), 3 Biomarkers, 148, 210 biomarker-based expression techniques, 148
149 breast cancer, 149
165, 153t for breast cancer detection, 75
78 microfluidics chips in biomarkers detection, 281
283 Biomedical MEMS, 254
255 Biomedical microelectromechanical systems (bioMEMS), 254 Biomedicine, 188 bioMEMS. See Biological microelectromechanical systems (bioMEMS); Biomedical microelectromechanical systems (bioMEMS) BioMEMS technology, 258
259 Biomimetic receptors, 9 Bionic e-Eye, 343
344, 344f Biopsy, 73 Bioreceptor element (BRE), 53, 229
230 Bioreceptors, 4, 6
10 antibodies, 6
7 aptamers, 8
9 enzymes, 7 nucleic acids, 7
8 Biorecognition element (BRE), 53, 79
84 antibodies, 80
81, 81f cells and tissues, 83 enzymes and proteins, 83 molecular imprints, 83
84 nucleic acids, 82
83 Biosensing biosensing-based drug delivery system, 250 strategy, 184 Biosensing
drug delivery systems, 250 applications, 256 adaptation to in vivo applications, 250
254 case studies, 257
259 image of wearable diabetes monitoring and therapy system, 259f materials and devices as, 254
257

biomedical or biological MEMS, 254
255 electrochemical biosensors, 256
257 microdevices, 256 responsive hydrogels, 254 Biosensors, 4
5, 53
56, 53f, 79
80, 172, 209, 228
232, 250, 296, 315, 316f aptamers, 230 bioreceptor elements, 229 biosensor-integrated drug delivery systems, 249
250 classification, 80
86 BRE, 80
84 transducer technology, 84
86 features of, 53
54, 54t future prospects, 64 in health care, 86
87 merits of aptamers over antibodies, 230
232 merits over conventional methods, 54 nanobiosensing, 232 parts, 229 for polyphenols analysis in beverages, 120
135 for rapid breast cancer detection, 88
98 BRCA1, 88
91, 88t, 89t CA 15
3 biosensors, 96 CEA, 93 ERα, 92 HER2, 93
95 miRNA 155, 97
98 miRNA 21, 96
97 MUC1, 95 PR, 92
93 for ROS determination, 108
115 sensitivity, 252 technology, 228 types, 54
55, 229 Biotin
streptavidin interaction system, 14
15 Bipolar electrode (BPE), 282
283 Bis(2-ethylhexyl) sebacate (DOS), 216
217 Blood glucose level, 249
250 Blood-based disease diagnosis, 210 BMIPF6. See 1-N-butyl-3methylimidazolium hexafluorophosphate (BMIPF6) BN. See Boron nitride (BN)

399 BN/PDDA/AuNCs nanocomposite, 162
163 BoNTs. See Botulinum neurotoxins (BoNTs) Boron nitride (BN), 162
163 Boron-doped silicon nanowires, 382
383 Bottom
up processes, 384 Botulinum neurotoxins (BoNTs), 273 Bovine serum albumin (BSA), 294 Bovine serum albumin-CLB (BSACLB), 300 BPE. See Bipolar electrode (BPE) BRCA1. See BReast CAncer 1 (BRCA1) BRE. See Bioreceptor element (BRE); Biorecognition element (BRE) Breast cancer, 71, 147
148 biomarkers, 149
151 biosensors, 79
87 for rapid breast cancer detection, 88
98 electrochemical immunosensing of breast cancer protein biomarkers, 151
165, 153t electrochemical immunosensors, 149, 150f signal amplification strategies, 151, 157f future prospects and challenges, 165
166 risk factors for, 71
73, 72f types, 73
75 biomarkers, 75
78 conventional detection methodology, 78
79 IDC, 74
75 noninvasive or in situ, 73
74 BReast CAncer 1 (BRCA1), 72
73, 76, 88
91, 88t, 89t BReast CAncer 2 (BRCA2), 72
73, 76, 162 Breast Cancer Research Foundation, 147
148, 162 Breath water vapor sensing for respiration monitoring, 217
218, 218f Breath-acetone monitoring, 218
219 Breathalyzers, 214 Brevibacterium sterolicum (BCO), 316 BSA. See Bovine serum albumin (BSA) BSA-CLB. See Bovine serum albuminCLB (BSA-CLB) Bulk wave sensor (BW sensor), 20

400 C CA. See Cancer antigen (CA) CA 15
3. See Cancer antigen 15
3 (CA 15
3) CA 27.29. See Cancer antigen 27.29 (CA 27.29) CA549, 151
156 cAbs. See Capture antibodies (cAbs) Caffeic acid, 121, 122t, 128
131 Calcium ion, 217 Calorimetric biosensors, 86 Camelidae nanobodies, 273 Campylobacter jejuni, 2, 20
21, 49, 273 Cancer, 147, 227, 249 cancer stage
based diagnostic biomarkers, 149
151 early stages of development, 148 Cancer antigen (CA), 151
157 Cancer antigen 15
3 (CA 15
3), 76
77, 91t, 96, 151
156, 159
160, 160f Cancer antigen 27.29 (CA 27.29), 76
77, 151
156 Cantilever-based biosensors, 238 Caplacizumab, 268
269 Capture antibodies (cAbs), 18
19, 149, 151 Carbon carbon-based nanomaterials, 386
388 hybridization states, 387f carbon-doped golden wattle-like TiO2 microspheres, 175
176 carbon
organic hybrid materials, 177
179 carbon
polymer hybrid platform, 179 materials, 177
178 Carbon dot/hemoglobin (CD/Hb), 325 Carbon ionic liquid electrode (CILE), 159
160 Carbon nanotubes (CNTs), 4, 151, 174, 210
212, 252, 280, 379t, 382 synthesis, 388f Carbon nitride (C3N4), 301 Carbon paste electrodes (CPE), 286 surface classification of CPE-based sensors, 368 Carboxylated g-C3N4, 303
304 Carcino embryonic antigen (CEA), 76
77, 89t, 93, 160
162, 293
294

INDEX

Cardiovascular disease (CVD), 249, 317 Carotenoids, 105
106 CAT. See Catalase (CAT) Catalase (CAT), 113, 289
290 Catalytic sensors, 229 Catechin, 121, 122t, 128
131 Catechol, 128
131, 366 Cathepsin D, 78 CB. See Conduction band (CB) CCK-8. See Cell contain kit-8 assay (CCK-8) CD-105. See Cluster of differentiation 105 antigen (CD-105) CD-146. See Cluster of differentiation 146 antigen (CD-146) CD/Hb. See Carbon dot/hemoglobin (CD/Hb) CDC. See Center for Disease Control and Prevention (CDC) CDR. See Complimentary determining regions (CDR) CdSe nanocrystals (CdSe NCs), 293
294 CdTe nanocrystals (CdTe NCs), 293
294 CdTe-graphene oxide, 290 CEA. See Carcino embryonic antigen (CEA) CEA antibody (Ab1(CEA)), 293
294 Cell contain kit-8 assay (CCK-8), 344
345 Cell surface protein receptors, 149
151 Cell viability biosensor, 344
345, 347f Cells, 83 Cellular bioreceptor, 9
10 Center for Disease Control and Prevention (CDC), 45
46 CeO2 nanocrystallines (CeO2 NCs), 297
298 Cerium, 385 Cerium oxide nanomaterials, 379t Chaotropic agents, 26
27 Charge exporters, 357 Chemical methods, 385 Chemical sensors, 172, 376
382, 379t Chemical vapor deposition (CVD), 295, 322
323, 387 graphene synthesis via, 362 Chemically modified electrodes (CMEs), 116
117

Chemiluminescence (CL), 113
114, 191 materials, 191
192 sensor intensity, 334 Chemokines, 266
267 receptors, 266
267 ChEt. See Cholesterol esterase (ChEt) Chitosan (Chi), 10
12, 210
212 Chlorogenic acid, 122t Cholera toxin (CT), 26 Cholesterol absorption, 317
320 amperometric determination for, 322f assay principle, 320f biosensor, 325, 331f detection, 315
316 electrocatalytic reaction, 325f enzymatic sensors, 320
330 fabrication technique, 323f future prospects, 336 nonenzymatic sensors, 331
336 oxidization, 316 structures, 316, 317f testing, 344
345, 347f Cholesterol esterase (ChEt), 319
322 Cholesterol oxidase (ChOx), 302, 318, 321, 325 enzyme reaction, 319f fabrication of ChOx/MWCNTs/ chitosan-modified GCE, 329f ChOx. See Cholesterol oxidase (ChOx) Chronic respiratory diseases, 269 Chronopotentiometric detection, 13 Chymotrypsin, 364 CILE. See Carbon ionic liquid electrode (CILE) CK. See Creatine kinase (CK) CL. See Chemiluminescence (CL) Clarck Electrode, 79 Clenbuterol (CLB), 300 Clostridium botulinum, 273 Cluster of differentiation 105 antigen (CD-105), 162 Cluster of differentiation 146 antigen (CD-146), 162 CMEs. See Chemically modified electrodes (CMEs) CNPs, 23
24 CNTs. See Carbon nanotubes (CNTs) Co3O4 nanorods, 299 Coagulation, 384
385 Colloidal clusters, 172
173

INDEX

Colorimetric biomimetic, 9 Colorimetric biotransducers, 148
149 Commonwealth Scientific and Industrial Research Organization (CSIRO), 46
49 Complimentary determining regions (CDR), 263
264 Concanavalin A, 14, 83 Conducting polymers, 188, 252, 379t, 389 conducting polymer-based nanohybrids, 189 Conduction band (CB), 293, 305 Conductometric aptasensors, 62 Conductometric biosensor, 14
15, 85 Conventional detection methodology, 78
79 Conventional methods, 50
53 in health care, 86
87 Copper oxides, 385 nanomaterials, 379t Copper sulfide and reduced graphene (CuS-rGO), 151
156 Copper(II) (Cu21), 307 Coupling inorganic materials, 176
177 Covalent, 177 bonding, 253 covalently linked flavin, 318 method, 389
390 Cover-type nonenzymatic freecholesterol sensor, 331
332 CPE. See Carbon paste electrodes (CPE) Creatine kinase (CK), 282 Cross-linker, 389 CSIRO. See Commonwealth Scientific and Industrial Research Organization (CSIRO) CT. See Cholera toxin (CT) Culture-based methods, 50 Cu
NG nanocomposites, 297 Cupric reducing antioxidant capacity (CUPRAC), 121 Cuprous oxide (Cu2O), 174 CuS-rGO. See Copper sulfide and reduced graphene (CuS-rGO) Cut carbon nanotubes, 295
296 CV. See Cyclic voltammetry (CV) CVD. See Cardiovascular disease (CVD); Chemical vapor deposition (CVD) Cyanamide, 302

Cyclic voltammetry (CV), 116, 368 Cyclin D1, 77 Cyclin E, 77 Cyclodextrin, 179 Cystic fibrosis, 216
217 Cyt c. See Cytochrome c (Cyt c) Cyt c
based biosensors, 114 Cytochrome biosensor, 364 Cytochrome c (Cyt c), 112, 364

D DA. See Dopamine (DA) dAb. See Detector antibody (dAb) DAB. See Diaminobenzidine (DAB) DAEC. See Diffusely adherent E. coli (DAEC) DC. See Downconversion (DC) DCIS. See Ductal carcinoma in situ (DCIS) Deep eutectic solvents (DESs), 187
188 DEFT. See Direct epifluorescence technique (DEFT) Degraded biosensors, 250 Dendrimers, 379t Deoxygenation process in alkaline solution, exfoliated graphite oxide via, 360
361 Deoxynivalenol (DON), 286 DESs. See Deep eutectic solvents (DESs) DET. See Direct electron transfer (DET) Detection of alcohols, 213
214 of ions, 214
217 of organic compounds, 212
214 of saccharides, 210
212 Detection limit (DL), 364
365, 379 Detector antibody (dAb), 18
19 Diabetes mellitus, 210
212, 249
250, 257 diagnosis, 218
219 Diagnostic biomarkers, 148 Diaminobenzidine (DAB), 321
322 Diazirines, 194 Diet-derived antioxidants, 106 Differential pulse voltammetry (DPV), 116, 149 Differential scanning calorimetry (DSC), 390 Diffusely adherent E. coli (DAEC), 1
2

401 Digital microfluidics, 284 Dihydroxy aromatic compound
based biosensor, 366 Dimethyl N-oxide pyrroline (DMPO), 109 (Dimethyl-4-phenylenediamine) (DMPD), 121 2,2-Diphenyl-1-picrylhydrazyl (DPPH), 116, 121 Direct electron transfer (DET), 289 Direct epifluorescence technique (DEFT), 22 Disease-causing agents, 45
46 Dismutase enzymes, 105
106 Dispersion at bulk-scale, 384
385 Disposable microfluidic immunoarray device (DμID), 156
157 Dithiocarbamate DNA (DTC-DNA), 299 DL. See Detection limit (DL) DMPD. See (Dimethyl-4phenylenediamine) (DMPD) DMPO. See Dimethyl N-oxide pyrroline (DMPO) DNA, 7, 379t biosensors, 132
135, 134t guanine oxidation to 8oxoguanine, 133f degradation, 132 DNA-triplex affinity capture method, 233
234 hybridization, 25 detection with microfluidic chips, 284 microstructures, 393 minor groove, 285 sensors, 299 DNFB. See 1-Fluoro-2,4-dinitrobenze (DNFB) DON. See Deoxynivalenol (DON) Dopamine (DA), 286, 296, 366 Doping polypyrrole, 180
181 DOS. See Bis(2-ethylhexyl) sebacate (DOS) Downconversion (DC), 194
195 Downsizing, 375
376 Doxorubicin-conjugated CdTe QDs (Dox-QDs), 290
291 DPV. See Differential pulse voltammetry (DPV) Droplet-based microfluidic systems, 286

402 Drug delivery systems, 249 DSC. See Differential scanning calorimetry (DSC) DSN. See Duplex-specific nuclease (DSN) DTC-DNA. See Dithiocarbamate DNA (DTC-DNA) Ductal carcinoma in situ (DCIS), 73 Duplex-specific nuclease (DSN), 305
307 Dye-doped polystyrene NPs, 25 DμID. See Disposable microfluidic immunoarray device (DμID)

E E. coli of serogroup O157:H7 (E. coli O157:H7), 1
2 EAEC. See Enteroaggregative (EAEC) ECD. See Extracellular domain (ECD) ECL. See Electrochemiluminescence (ECL) ECL resonance energy transfer (ERET), 290
291 EDC. See 1-Ethyl-3-(3 dimethylaminopropyl)carbodiimide hydrochloride (EDC) EDC/N-hydroxy-succinimide coupling, 165 EDC/NHS. See 1-Ethyl-3-(3dimethylaminopropyl)carbodiimide/Nhydroxysuccinimide (EDC/NHS) EDTA. See Ethylenediaminetetraacetic acid (EDTA) EGFR. See Epidermal growth factor receptor (EGFR) EHEC. See Enterohemorrhagic (EHEC) EIA. See Enzyme immunoassay (EIA) EIEC. See Enteroinvasive (EIEC) EIS. See Electrochemical impedance spectroscopy (EIS) Electroanalytical approaches, 110 Electrochemical aptasensors, 62
63, 63f amperometric method, 62 conductometric aptasensors, 62 impedimetric detection, 62
63 potentiometric detection, 62 Electrochemical bioactive small molecule biosensing, 296
297

INDEX

Electrochemical biosensing applications, 181
182 of polyphenols in beverages, 125
135 DNA biosensors, 132
135 enzymatic biosensors, 126
132, 129t Electrochemical biosensors, 10
15, 11f, 108, 115
120, 117f, 229, 233, 256
257, 279. See also Optical biosensors amperometric biosensor, 10
12 based on oxidase enzymes, 116
117 for cholesterol, 321
322 conductometric biosensor, 14
15 g-C3N4 based nanomaterials, 301
307 graphene, 294
301 impedimetric biosensor, 13
14 microfluidics chips, 280
287 potentiometric-based biosensors, 12
13 QDs, 287
294 Electrochemical cell biosensing based on graphene, 300
301 Electrochemical DNA biosensing based on graphene, 299 sensors, 366 Electrochemical enzyme biosensing based on graphene, 297
299 based on QDs, 287
289, 288f Electrochemical gene biosensing based on QDs, 289
292 Electrochemical immunobiosensing based on graphene, 299
300 Electrochemical immunosensing, 149, 151 based on QDs, 293
294 of breast cancer protein biomarkers, 151
165, 153t BRCA-1 and BRCA-2, 162 cancer antigen, 151
157 CD-146 and CD-105, 162 CEA, 160
162 EGFR, 157
158 HER 2 and 3, 158
160 IL-6 and IL-8, 162
163 VEGFRs, 160 Electrochemical immunosensor approach, 93
95 Electrochemical immunosensors, 149 working principle, 150f

Electrochemical impedance spectroscopy (EIS), 13, 190, 285, 368 Electrochemical impedimetric spectroscopy (ESI), 236 Electrochemical nanobiosensors, 97
98, 232
233, 235
237 Electrochemical sensing, 367 Electrochemical sensors, 303
305 for detection of chemicals, 367 Electrochemical synthesis, 189 Electrochemical transducers, 116
117 Electrochemical-enzymatic redox cycling, 323
324 Electrochemically deposited PANI films, 326
327 Electrochemically reduced graphene oxide (ERGO), 160
162 Electrochemiluminescence (ECL), 187, 213 enzyme biosensing, 289 materials, 191
192 sensors, 279, 364 Electrode, 236 modification, 14 Electromotive force (EMF), 12 Electron spin resonance (ESR), 109 Electronic (bio)chemical sensing, 378 Electronic tongue, 132 Electrophoresis, 125 Electropolymerization process, 190 Electropolymerized membranes, 253 Electropolymerized PANI-MWCNT matrix, 327
328 Electrospun PANI/polystyrene blended fiber, 328 Electrostatic repulsion, 184, 253 ELFA. See Enzyme-linked fluorescent assay (ELFA) ELISA. See Enzyme-linked immunosorbent assay (ELISA) Embedded biosensor, 250 EMF. See Electromotive force (EMF) Emulsion droplets, 184 Endogenous antioxidant, 106 Energy transfer (ET), 290 Enteroaggregative (EAEC), 1
2 Enterobacteriaceae, 2 Enterohemorrhagic (EHEC), 1
2 Enteroinvasive (EIEC), 1
2 Enteropathogenic (EPEC), 1
2 Enterotoxigenic E. coli (ETEC), 1
2, 272
273

INDEX

Entrapment of enzyme, 118 Enzymatic/enzymes, 7, 83 biosensors, 119t, 126
132, 135 detection method, 96
97 enzyme-based biosensors, 252 enzyme-based electrochemical biosensors, 256 enzyme-catalyzed redox reaction, 235
236 immobilization, 116
118, 253 sensors, 320
330 Enzyme immunoassay (EIA), 3 Enzyme-linked fluorescent assay (ELFA), 3 Enzyme-linked immunosorbent assay (ELISA), 3, 148
149, 228, 282 EPEC. See Enteropathogenic (EPEC) EPI. See Exciton
plasmon interactions (EPI) Epicatechin, 122t, 128
131 Epicatechingallate, 122t Epidermal growth factor receptor (EGFR), 157
158, 266
267 Epigallocatechin-gallate, 122t Epigenomic DNA, 149
151 ERET. See ECL resonance energy transfer (ERET) ERGO. See Electrochemically reduced graphene oxide (ERGO) ERα. See Estrogen-alpha (ERα) Escherichia coli, 1
2, 285 ESI. See Electrochemical impedimetric spectroscopy (ESI) ESR. See Electron spin resonance (ESR) Estrogen, 149
151 Estrogen-alpha (ERα), 75, 89t, 92 ET. See Energy transfer (ET) Et3N. See Triethylamine (Et3N) ETEC. See Enterotoxigenic E. coli (ETEC) Ethanol, 214, 215f 1-Ethyl-3-(3 dimethylaminopropyl)carbodiimide hydrochloride (EDC), 165 1-Ethyl-3-(3-dimethylaminopropyl)carbodiimide/Nhydroxysuccinimide (EDC/NHS), 6
7 Ethylene glycol dimethacrylate, 390 Ethylenediaminetetraacetic acid (EDTA), 364
365

Ethyleneglycol dimethacrylate, 185
186 Exciton
plasmon interactions (EPI), 305
307 Exfoliated g-C3N4, 301 Exfoliation process, 357
361 graphite exfoliation via surfactantassisted emulsion process, 359
360 graphite oxide via deoxygenation process in alkaline solution, 360
361 liquid exfoliation of layered materials, 359 Exhaled breath analysis, 210 Exhaled breath gas sensors, 222 Exhaled VOC monitoring, 218
222 acetone sensing for diabetes diagnosis, 218
219 artificially intelligent nanosensors, 221
222 hydrogen sulfide detection for halitosis diagnosis, 219
220 NO gas detection for asthma diagnosis, 220
221 Exogenous antioxidants. See Dietderived antioxidants Extracellular domain (ECD), 158
159

F FAD. See Flavin adenine dinucleotide (FAD) Farm organization information system, 349
350 FBDs. See Foodborne diseases (FBDs) FC. See Folin-Ciocalteu (FC) FC method. See Folin-Ciocalteau method (FC method) Fc-labeled aptamers. See Ferrocenelabeled aptamers (Fc-labeled aptamers) Fe3O4@Au nanocomposite biosensor platform, 385 Fe3O4NPs. See Iron oxide nanoparticles (Fe3O4NPs) Fenton reaction, 132, 132f Ferredoxin, 114 Ferric reducing ability of plasma (FRAP), 121 Ferric reducing antioxidant power (FRAP), 115
116 Ferricyanide ([Fe(CN)6]32), 323
324 Ferrocene (Fc), 192

403 Ferrocene-labeled aptamers (Fclabeled aptamers), 303
304 Ferrocyanide ([Fe(CN)6]42), 323
324 Fiber optic biosensor, 324 Field effect transistors (FETs), 210
212, 362
363 FETs
based biosensors using functional graphene materials, 362
363 FITC. See Fluorescein isothiocyanate (FITC) Flagellum, 273 Flavin adenine dinucleotide (FAD), 298
299, 318 Fluorescein isothiocyanate (FITC), 25 Fluorescence resonance energy transfer (FRET), 18, 93, 148
149, 192
193 Fluorescent-silica NPs (FSNPs), 4, 5t Fluorescent/fluorescence biosensor, 364
365 dual-channel biosensor model, 88
91 fluorescence-based assay, 3 materials, 192
195 hybrid materials for surface plasmon resonance, 195
196 luminescent optical labels, 194
195 PEC materials, 193
194 method, 113
114 phenomenon, 240
241 quenching of CD, 325, 326f tags, 7 Fluorine-doped tin oxide (FTO), 110
111 1-Fluoro-2,4-dinitrobenze (DNFB), 268 Fluorometric sensing, 240
241 FMDV. See Foot-and-mouth disease virus (FMDV) Folic acid, 296 Folin-Ciocalteau method (FC method), 125 Folin-Ciocalteu (FC), 115
116 spectrophotometric, 118 Food, 45
46 poisoning, 45
46 Foodborne illnesses, 46 pathogens, 1
2, 24, 46, 47t conventional methods, 50
53, 51t, 52t detection, 45
46, 50
53, 51f

404 Foodborne diseases (FBDs), 45
46 Foods antioxidant capacity assessment, 115
120 Foot-and-mouth disease virus (FMDV), 269 Fourier transform infrared spectroscopy (FT-IR spectroscopy), 18
19 FPGA-based fused smart sensor, 343 Framework regions (FR), 263
264 FRAP. See Ferric reducing ability of plasma (FRAP); Ferric reducing antioxidant power (FRAP) French paradox, 120 FRET. See Fluorescence resonance energy transfer (FRET) FSNPs. See Fluorescent-silica NPs (FSNPs) FT-IR spectroscopy. See Fourier transform infrared spectroscopy (FT-IR spectroscopy) FTO. See Fluorine-doped tin oxide (FTO) Functional nanoparticles, 382 Functionalized advanced hybrid materials, 171 advanced inorganic hybrid materials, 172
176 advanced organic hybrid materials, 188
190 advanced organic
inorganic hybrid materials, 176
188 optical multifunctional advanced hybrid materials, 190
196 Functionalized clays and silica, 186
187 Functionalized graphene nanocomposites, 296

G g-C3N4. See Graphitic carbon nitride (g-C3N4) G-PANI. See Graphene-polyaniline (GPANI) G-quadruplex, 289 3G/GPRS receiver, 342
343 GA. See Glutaraldehyde (GA) Gadolinium diethylenetriamine pentaacetic acid (Gd-DTPA), 78 Gal. See β-galactosidase (Gal) Gallbladder, 318 Gallic acid, 121, 122t, 128
131

INDEX

Ganglioside, 26 Gas chromatography (GC), 286 Gas sensors for healthcare, 217
224 breath water vapor sensing for respiration monitoring, 217
218 exhaled VOC monitoring, 218
222 ingestible sensors for gut-gas monitoring, 222
224 GC. See Gas chromatography (GC) GCE. See Glassy carbon electrode (GCE) Gd-DTPA. See Gadolinium diethylenetriamine pentaacetic acid (Gd-DTPA) Gene Expression Omnibus (GEO), 78 Gene-Z, 346
349, 349f Genomic DNA, 149
151 Genomic SELEX, 56, 58f Genosensor platforms, 7
8 GEO. See Gene Expression Omnibus (GEO) Glassy carbon electrode (GCE), 88
91, 178, 297
298, 325, 362 Global positioning system (GPS), 349
350 Glucose biosensor, 232 meter, 281 Glucose, 96
97, 210
212, 296 glucose-responsive hydrogel-based insulin delivery systems, 257 glucose-responsive system, 257
258 Glucose oxidase (GOx), 160
162, 194, 210
212, 286
289 enzymes, 257 Glutaraldehyde (GA), 19 Glycoprotein (gp), 270
271 GNP. See Gold nanoparticles (AuNPs); Graphene platelet (GNP) GNP-SPCE. See Gold nanoparticlemodified screen-printed carbon electrode (GNP-SPCE) GO. See Graphene oxide (GO) GOD. See Glucose oxidase (GOx) Gold (Au), 10 Gold nanoclusters protected by bovine serum albumin (AuNCs-BSA), 110 Gold nanoclustures (AuNCs), 162
163

Gold nanoflowers (AuNFs), 29
33, 196 Gold nanoparticle-modified screenprinted carbon electrode (GNPSPCE), 62
63 Gold nanoparticles (AuNPs), 4, 5t, 88
91, 110, 131
132, 151, 159
162, 191
192, 232
233, 238
239, 241f, 280, 290, 379t, 382
384 AuNP-based LFIA, 24 AuNP-modified surface, 156
157 AuNPs-decorated g-C3N4 nanosheets, 302
303 AuNPs/Cu-MOF nanoconjugate, 160
162 AuNPs
polypyrrole composite, 180
181 GO
TiO2. See Graphene
titanium dioxide (GO
TiO2) GOx. See Glucose oxidase (GOx) GPS. See Global positioning system (GPS) GR. See Graphene (GR) Gram-negative bacteria, 2 Grapes, 120
121 Graphene (GR), 177, 178f, 182, 294
301, 379t, 382, 387
388 electrochemical bioactive small molecule biosensing, 296
297 cell biosensing based on, 300
301 DNA biosensing based on, 299 enzyme biosensing based on, 297
299 immunobiosensing based on, 299
300 graphene-based electrodes, 366 synthesis, 295
296, 388f Graphene nanosheets (GS), 300 Graphene oxide (GO), 114, 156, 177, 219
220, 234
235, 294, 387
388 GO
based hybrid electrodes, 367 Graphene platelet (GNP), 362, 363f Graphene sheets -Nafion film, 300 Graphene-based functional hybrid materials, 357 applications, 358f electrochemical sensors and electrodes for detecting biomolecules, 362
368 synthesis of 3D structured graphene nanosheets, 357
362

405

INDEX

Graphene-polyaniline (G-PANI), 286 Graphene
graphene oxide (G-GO), 14 Graphene
titanium dioxide (GO
TiO2), 173 Graphite, 10 exfoliation via surfactant-assisted emulsion process, 359
360 Graphite-Teflon-peroxidase-ferrocene electrode, 25
26 Graphitic carbon nitride (g-C3N4), 280 based nanomaterials, 301
307 amperometric sensors, 301
303 electrochemical sensors, 303
305 PEC sensors, 305
307 GS. See Graphene nanosheets (GS) Guanine oxidation to 8-oxoguanine, 133f Gut-gas monitoring, 222
224

H HA. See Hemagglutinin (HA); Hyaluronic acid (HA) Hairpin DNA1 (H1), 290
291 Hairpin DNA2 (H2), 290
291 Halitosis, 219
220 HAT. See Hydrogen atom transfer (HAT) HCAbs. See Heavy-chain-only antibodies (HCAbs) HCV. See Hepatitis C virus (HCV) HDT. See (Z)-2-((Xhexanedithiol (HDT) Healthcare, gas sensors for, 217
224 Heavy-chain-only antibodies (HCAbs), 263 generation, 264
265 HEMA-co-DMAEMA. See Poly(2hydroxyethyl methacrylate-coN,N-dimethylaminoethyl methacrylate) (HEMA-coDMAEMA) Hemagglutinin (HA), 269
270 Hemoglobin (Hgb), 114, 365
366 Hemolytic uremic syndrome (HUS), 1
2, 49 Hep3B liver cancer cells, 300
301 Hepatitis, 214
216 Hepatitis A virus, 2 Hepatitis C virus (HCV), 271
272 Hepatocellular carcinoma, 271
272 HER1. See Human epidermal growth factor receptor 1 (HER1) Hgb. See Hemoglobin (Hgb)

High-operating temperature, 171
172 HIV-1. See Human immune deficiency virus-1 (HIV-1) Hoechst 33258 redox molecules, 285 Hormonal therapy, 74 Horseradish peroxidase (HRP), 7, 93, 112, 116
117, 321
322 biosensor, 365 Horseradish peroxidise-labeled antibody functionalized with gold nanoparticles (HRP-Ab2/ Au NPs), 300 HRP. See Horseradish peroxidase (HRP) HRP/anti-CEA/AuNPs/ZnONPs/Au sensing platform, 160
162, 161f Human epidermal growth factor receptor 1 (HER1), 157
160 Human epidermal growth factor receptor 2 (HER2), 75, 90t, 93
95, 149
151, 158
160, 266
267 ECD, 159
160, 160f Human epidermal growth factor receptor 3 (HER3), 158
160 Human immune deficiency virus-1 (HIV-1), 269 Human serum albumin, 269 Huperzine-A (hupA), 302
303 HUS. See Hemolytic uremic syndrome (HUS) Hyaluronic acid (HA), 14 Hybrid, 171 materials, 171 for surface plasmon resonance, 195
196 Hybrid phenothiazine platform-based cyanine dye (MPT-Cy2), 109
110 Hybridization, 174 Hydrazine, 361 Hydrogels, 254 Hydrogen atom transfer (HAT), 115
116 Hydrogen peroxide (H2O2), 286
287, 296, 321, 365
366 biosensors for, 113
115 Hydrogen sulfide detection for halitosis diagnosis, 219
220 Hydrolysis of cholesterol ester, 321 Hydrophobic effect, 176 Hydroquinone, 128
131

4-Hydroxy-2,2,6,6tetramethylpiperidine-N-oxyl (Hydroxy-TEMPO), 109 Hydroxyl radical (HO ), 109, 132 biosensors for, 109
111 Hypertension, 249 Hypervariable regions, 263
264 Hypoxanthine, 256



I IAs. See Immunoassays (IAs) IBC. See Inflammatory breast cancer (IBC) IDC. See Invasive ductal carcinoma (IDC) IgG. See Immunoglobulin G (IgG) IL-6. See Interleukin-6 (IL-6) ILA. See Immunoligand assay (ILA) ILC. See Invasive lobular carcinoma (ILC) ILs. See Ionic liquids (ILs) Imaging modalities, 383
384 IMID. See Immune-mediated inflammatory diseases (IMID) Immobilization, 6
7, 117
118, 321 Immune-based assays, 228 Immune-mediated inflammatory diseases (IMID), 268 Immunoassays (IAs), 10 Immunodetection-based methods, 3
4 Immunoglobulin E neutralization (IgE neutralization), 269 Immunoglobulin G (IgG), 263, 280 Immunohistochemistry methods, 148
149 Immunoligand assay (ILA), 12 Immunological techniques, 3 Immunosensors, 299 Impedance spectroscopy, 25 Impedimetric biosensor, 13
14 detection, 62
63 immunosensor, 159
160 Imprinted microporous
MOF, 182 In situ carcinoma, 73
74 In vivo applications, 250
254 biocompatibility, 250
252 selectivity, 252
254 sensitivity, 252 In2O3-based FET fabricated on PET substrate, 210
212, 211f Indium oxide (In2O3), 210
212

406 Indium
tin-oxide (ITO), 14, 88
91, 284, 326 electrode, 114, 156 glass plate, 334
336 Induce immune reactions, 265 Inflammation, 268 Inflammatory breast cancer (IBC), 74 Inflammatory diseases, 268
269 Influenza virus, 269
270 Infrared (IR), 15 Ingestible sensors for gut-gas monitoring, 222
224 InlB. See Internalin B (InlB) Inorganic characteristics, 171 compounds, 171
172 membranes, 22 nanoparticles, 175 sensing materials, 171
172 Intelligent hydrogels, 254 Intensifier, 213 Interleukin-6 (IL-6), 162
163, 281 Interleukin-8 (IL-8), 162
163 Internalin B (InlB), 14 International Organization for Stabilization (ISO), 4 Intraoral halitosis, 219
220 Invasive ductal carcinoma (IDC), 74
75 Invasive lobular carcinoma (ILC), 74 Invasive method, 222
224 Ion channels, 268 Ion detection, 214
217 ammonium ion, 214
216 calcium ion, 217 sodium ion, 216
217 Ion-selective, field-effect transistors (ISFETs), 12, 236 Ionic liquids (ILs), 187
188, 212 hybrid materials, 187
188 Ionic-covalent, 177 iPlate, 344, 344f, 345f chief interface of home-based software iPlate screen, 347f iPod Touch, 346
349 IR. See Infrared (IR) Iridium oxide nanoparticles (IrO2 NPs), 62
63 Iron oxide nanoparticles (Fe3O4NPs), 159
160 Iron oxide nanoparticles, 385 Iron-based biosensor, 365
366

INDEX

ISFETs. See Ion-selective, field-effect transistors (ISFETs) ISO. See International Organization for Stabilization (ISO) ITO. See Indium
tin-oxide (ITO)

J Janus particles, 173 Joule magnetostriction principle, 21

K Kanamycin, 241 Kaolinite, 186
187

L Lab on chip, 280 Label-free detection, 8 electrochemical nanobiosensors, 236, 236t immunosensor, 93
95 nanobiosensors, 232
238 electrochemical nanobiosensors, 235
237 mechanical transducer
based nanobiosensors, 238 optical nanobiosensor, 233
235 optofluidic ring resonator sensor, 96 potentiometric aptasensor, 13 Labeled nanobiosensors, 238
244 labeled electrochemical nanobiosensors, 242
244, 243t labeled optical nanobiosensors, 240
241, 242t nanozyme-based turn-off/turn-on approach, 241 Laccase, 116
117, 126, 126f Lactate sensor, 212f Lactose, 212 electrochemical reactions at gate electrode, 213f Laminin (LN), 294 LAMP. See Loop-mediated isothermal amplification (LAMP) Langmuir
Blodgett PANI-stearic acid (LB PANI-stearic acid), 326
327 LAPSs. See Light-addressable potentiometric sensors (LAPSs) Lateral-flow immunoassay (LFIA), 22
25, 23f

reader device detection, 24
25 visual detection, 23
24 Layered materials, liquid exfoliation of, 359 LB PANI-stearic acid. See Langmuir
Blodgett PANIstearic acid (LB PANI-stearic acid) LC. See Liquid chromatography (LC) LCIS. See Lobular carcinoma in situ (LCIS) LDL. See Low-density lipoprotein (LDL) LDL-c. See Lipoprotein cholesterol (LDL-c) Lectins, 83 LEDs. See Light-emitting diodes (LEDs) LFIA. See Lateral-flow immunoassay (LFIA) Light-addressable potentiometric sensors (LAPSs), 12 Light-emitting diodes (LEDs), 12 Limit of detection (LOD), 29, 156
157, 366 Linear sweep voltammetry (LSV), 149, 285 Lipoprotein (Lp), 365 lipoprotein-based biosensor, 365 Lipoprotein cholesterol (LDL-c), 317 Liquid chromatography (LC), 286 Liquid exfoliation of layered materials, 359 Listeria monocytogenes, 13 Listeriolysin of Listeria monocytogenes, 8
9 Lithium niobite (LiTaO3), 20 LN. See Laminin (LN) Lobular carcinoma in situ (LCIS), 73 LOD. See Limit of detection (LOD) Loop-mediated isothermal amplification (LAMP), 285 LoRa Wizard, 350, 351f Love-wave sensor, 344, 346f Low-density lipoprotein (LDL), 115
116 Low-temperature fabrication, 385 Lp. See Lipoprotein (Lp) LSV. See Linear sweep voltammetry (LSV) Luminescent optical labels, 194
195 pyrene-diazirine photoreactive platform, 195f

INDEX

M mAbs. See Monoclonal antibodies (mAbs) Magnetic beads (MBs), 110
111, 151 sandwich immunoassay using, 151, 152f streptavidin-coated, 157
158 Magnetic core MIP particles, 185 Magnetic materials, 182
186 working principle of magnetic nanofluid sensor, 185f Magnetic nanofluid
based nonenzymatic sensor, 184 Magnetic nanomaterials, 379t Magnetic nanoparticles (MNPs), 18
19, 25, 182, 382
385 Magnetic resonance imaging (MRI), 74, 78, 86
87, 193 Magnetic silica particles, 184 Magneto-genosensing, 26
27 Magnetoelastic biosensor (ME biosensor), 21
22 Mammalian antibodies, 263 MammaPrint, 79 Mammographies, 86
87 Mammography, 74, 78 Mass change
based biotransducers, 148
149 Mass spectrometry (MS), 125, 286 Mass-based biosensors, 20
22, 229 magnetoelastic biosensor, 21
22 piezoelectric biosensor, 20
21 Mass-sensitive nanobiosensors, 232
233 Matrix metallopeptidase 9 (MMP-9), 163
164 MB. See Methylene blue (MB) MBs. See Magnetic beads (MBs) MCF-7. See Michigan cancer foundation-7 (MCF-7) MDM-2. See Murine double minute 2 (MDM-2) ME biosensor. See Magnetoelastic biosensor (ME biosensor) Mechanical biosensors, 233 Mechanical exfoliation, 295, 388 Mechanical transducer
based nanobiosensors, 238 Mediator-free cholesterol and reusable biosensor, 322 Medical imaging, nanobodies application in, 267
268

Membrane-based biosensors, 22
26 LFIA, 22
25 Membrane-based multiplex biosensors, 29, 29f Membrane-proximal external region (MPER), 270
271 MEMS. See Microelectromechanical systems (MEMS) Menopause, 71
72 MEQ-LAMP. See Microfluidic electrochemical quantitative loop-mediated isothermal amplification (MEQ-LAMP) 3-Mercaptopropionic acid (MPA), 293 Mercaptosuccinic-acid-capped gold (MSAG), 190 Mesoporous graphitic carbon nitride (mpg-C3N4), 301
302 Mesoporous-silica hybrid materials, 172
173, 176 Mesoporous-SiO2 nanoparticles (MSN), 176, 191
192 Metabolism sensors, 229 Metal metal-based nanoparticles, 378
379 nanomaterials, 383
385 nanoparticles, 179
182, 252, 383
384 Metal organic frameworks (MOFs), 160
162 Metal oxide hybrids, 172
175 nanomaterials, 385 Metal
organic frameworks (MOFs), 179
182, 183f, 183t Methacrylic acid, 390 Methicillin-resistant S. aureus (MRSA), 9
10 Methoxy poly(ethylene glycol)-bpoly (ε-caprolactone) (mPEG-bPCL), 193 Methyl parathion, 385 1-Methyl-3-(3-sulfobutyl) imidazolium p-toluenesulfonate, 187 Methylene blue (MB), 96
97, 285 Michigan cancer foundation-7 (MCF7), 95 breast cancer cells, 300
301 Microdevices, 256, 256f Microelectromechanical systems (MEMS), 254
255 Microfluidic electrochemical quantitative loop-mediated

407 isothermal amplification (MEQLAMP), 285 Microfluidic paper-based analytical devices (μPADs), 280
281, 284 Microfluidics chips, 280
287 in bacteria detection, 285 in biomarkers detection, 281
283 in nucleic acid detection, 283
284 for small molecule detection, 285
287 microfluidics-based immunoassays, 156 Microporous filters, 390 MicroRNA-34a (miRNA-34a), 299 MicroRNAs (miRNA), 149
151, 284, 290
291, 292f miRNA 21, 92t, 96
97 miRNA 155, 92t, 97
98 Microsensors, 294 Microthrombi, 268 Microwave irradiation (MWI), 361 graphene synthesis by, 361 Microwave plasma enhanced chemical vapor deposition method (MWPE-CVD method), 361 MIP. See Molecularly imprinted polymer (MIP) miRNA. See MicroRNAs (miRNA) miRNA-34a. See MicroRNA-34a (miRNA-34a) MMP-9. See Matrix metallopeptidase 9 (MMP-9) MNPs. See Magnetic nanoparticles (MNPs) MOFs. See Metal organic frameworks (MOFs); Metal
organic frameworks (MOFs) Molecular imprinting, 83
84, 84f, 389 Molecularly imprinted polymer (MIP), 9, 83
84, 96, 179, 286, 379t, 389
390 fabrication, 389 Molybdenum disulfide (MoS2), 180
181 Monoclonal antibodies (mAbs), 6, 80
81, 265 Monosaccharides, 210 Morin (mor), 110
111 MoS2. See Molybdenum disulfide (MoS2) MPA. See 3-Mercaptopropionic acid (MPA)

408 MPA-CdS/RGO nanocomposites, 293 mPEG-b-PCL. See Methoxy poly (ethylene glycol)-bpoly (ε-caprolactone) (mPEG-b-PCL) MPER. See Membrane-proximal external region (MPER) mpg-C3N4. See Mesoporous graphitic carbon nitride (mpg-C3N4) MPT-Cy2. See Hybrid phenothiazine platform-based cyanine dye (MPT-Cy2) MRI. See Magnetic resonance imaging (MRI) mRNAs, 149
151 MRSA. See Methicillin-resistant S. aureus (MRSA) MS. See Mass spectrometry (MS) MSAG. See Mercaptosuccinic-acidcapped gold (MSAG) MSN. See Mesoporous-SiO2 nanoparticles (MSN) Mucin 1 protein (MUC1), 91t, 95 aptamers, 300
301 Mucins, 76
77 Multidrug-resistant bacteria, 286
287 Multiparameter assays, 76 Multiple disease detection, 221
222, 223f Multiplex biosensors, 26
29, 27f Multiplexed electrochemical immunosensor, 96 Multiwalled carbon nanotubes (MWCNTs), 10
12, 93
95, 114, 159
160, 280, 295
296 Murine double minute 2 (MDM-2), 163
164 Mutated genes, 149
151 MWCNTs. See Multiwalled carbon nanotubes (MWCNTs) MWI. See Microwave irradiation (MWI) MWPE-CVD method. See Microwave plasma enhanced chemical vapor deposition method (MWPE-CVD method) Mycotoxin, 24
25, 286 Myoglobin, 114

N 1-N-butyl-3-methylimidazolium hexafluorophosphate (BMIPF6), 118 N-cinnamoylanthranilic acids, 115

INDEX

N-doped graphene biosensor (N-GR biosensor), 151
156 NA. See Nucleic acids (NA) Na-TFPB. See Sodium tetra kis[3,5-bis (trifluoromethyl)phenyl]borate (Na-TFPB) NAAs. See Nucleic acid aptamers (NAAs) NAD1, 364
365 NADH. See Nicotinamide adenine dinucleotide (NADH) Nafion (Nf), 300 membranes, 113 NaNO2. See Sodium nitrite (NaNO2) Nanoadvanced sensing technology biological materials, 393 carbon-based nanomaterials, 386
388 challenges and future prospects of, 376 metal nanomaterials, 383
385 metal oxide nanomaterials, 385 nanofibers, 382
383 nanomaterial-based biosensors and chemical sensors, 376
382, 379t nanoparticles, 382
383 nanoprobes, 382
383 nanostructures, 382
383 nanowires, 382
383 polymer nanomaterials, 388
393 tubular and porous nanostructures, 383 Nanobiosensing, 232 Nanobiosensors, 232
244, 378
379 label-free, 233
238 labeled, 238
244 Nanobodies, 266 application in medical imaging, 267
268 against bacteria, 272
274 binding to transmembrane proteins, 266
267 characteristics for therapeutic application, 265
266 generation, 264
265, 265f technology, 268 against viruses, 269
272 Nanocomposites, 178 electrode, 326 Nanocontrolled release systems, 383 Nanodrug fabrications, 176

Nanofibers, 252, 382
383 Nanoimmunomagnetic separation (NIMS), 18
19 Nanomaterial-based biosensors, 114, 375 and chemical sensors, 376
382, 379t Nanomaterials, 382 aptasensors efficacy in combination with, 232 nanomaterial-based devices, 376 Nanoparticle-based biosensor performance, 29
33, 30t Nanoparticle-decorated Au nanoparticles (Ni
AuNPs), 164
165 Nanoparticles (NPs), 4
5, 24, 238
239, 280, 382
384, 384f nanoparticle-shaped film formation, 188 NP-based amperometric biosensors, 10 Nanoprobes, 382
383 Nanosensing, 376 Nanosensors, 376. See also Electrochemical biosensors platforms for biomedical diagnosis, 377f Nanosized carbon-based materials, 114 Nanostructure, 382
383 carbons, 357, 358f formation, 176 materials, 4 Nanosynthesis, 389, 393 Nanotechnology, 188, 232, 375 Nanotubes, 379
382 Nanowires, 379
383 NanoZyme activity, 241, 243f NanoZyme-based turn-off/turn-on approach, 241 NC. See Nitrocellulose (NC) NC membrane. See Nitrocellulose membrane (NC membrane) Near-infrared (NIR), 267, 287 Nernst relationship, 12 Neuron-based biosensors, 366 Nf. See Nafion (Nf) Ni
AuNPs. See Nanoparticledecorated Au nanoparticles (Ni
AuNPs) Nickel oxide (NiO), 331
332 NiCo2O4 labeling, 298
299

409

INDEX

Nicotinamide adenine dinucleotide (NADH), 364
365 NIMS. See Nanoimmunomagnetic separation (NIMS) NiO. See Nickel oxide (NiO) NIR. See Near-infrared (NIR) Nitric oxide (NO), 220
221 gas detection for asthma diagnosis, 220
221 Nitrocellulose (NC), 22 Nitrocellulose membrane (NC membrane), 23 2-Nitrophenyl octyl ether (o-NPOE), 214
216 NJRNG1157PCD-TE1, 349
350 NO. See Nitric oxide (NO) Nonenzymatic glucose biosensor, 332
333 proteins, 9
10 sensors, 331
336 Nonflavanoids, 121 Nonintrusive sensors, 209
210 Noninvasive carcinoma, 73
74 Noninvasive diagnostic biosensors for healthcare future perspectives, 224
225 gas sensors for healthcare, 217
224 wearable sweat sensors, 210
217 Noninvasive method, 222
224 Nonscalable cost-effective synthesis, 388 Nontyphoidal illness, 46
49 Norovirus (NoV), 46
50 C. coli O157:H7, 49 C. jejuni, 49 S. aureus, 49
50 Salmonella spp., 46
49 Vibrio spp., 50 NoV. See Norovirus (NoV) NPs. See Nanoparticles (NPs) Nucleic acid aptamers (NAAs), 82 Nucleic acids (NA), 7
8, 33, 82
83 microfluidics chips in NA detection, 283
284 Nucleotide base, 133 Nylon, 22
23 Nyquist plots, 135

O O-aminophenol (OAP), 289
290 o-dianisidine strategy, 326
327 o-NPOE. See 2-Nitrophenyl octyl ether (o-NPOE)

OAP. See O-aminophenol (OAP) Obesity, 72 Ochratoxin A (OTA), 24
25, 62
63 OD. See Optical density (OD) Okadaic acid, 344
345 BSA, 344 Omalizumab, 269 OmpD, 14 Oncotype DX, 79 One-dimensional nanomaterials, 379
382 One-step hydrothermal method, 367 OPs. See Organophosphorus pesticides (OPs) Optical biosensing applications, 192
193 Optical biosensors, 15
19, 16f, 57
62, 229, 233. See also Electrochemical biosensors mass-based biosensors, 20
22 optical fibers, 17
18 Raman and Fourier transform infrared spectroscopy, 18
19 SERS aptasensors, 58
62 SPR aptasensors, 57
58 surface plasmon resonance, 15
17 Optical density (OD), 24 Optical detection biosensors, 190
191 Optical fibers, 17
18 Optical multifunctional advanced hybrid materials, 190
196 chemiluminescent and electrochemiluminescent materials, 191
192 fluorescent materials, 192
195 Optical nanobiosensor, 232
235, 235t Optical sensors, 240, 349
350 Optical-based multiplex biosensors, 27
29, 28f ORAC. See Oxygen radical absorbance capacity (ORAC) Oral passive immunization, 272 Organic characteristics, 171 composites, 179
182 compounds, 212
214 detection of alcohols, 213
214 UA, 213 conductive polymers, 189 dye molecules, 3 membranes, 22 Organophosphorus pesticides (OPs), 302
303

Oseltamivir, 269
270 Osmium bipyridine, 114 OTA. See Ochratoxin A (OTA) Output signal, 281 Oxidase enzymes, 116
117 Oxidative stress, 106
109 Oxygen (O2), 288
289 Oxygen oxidoreductase. See Cholesterol oxidase (ChOx) Oxygen radical absorbance capacity (ORAC), 115
116, 121

P P-1,5DAN. See Poly-1,5diaminonaphthalene (P1,5DAN) p-aminothiophenol (ATP), 299 P-GO @CdS QDs, 289 P-peptide. See Phosphorylated peptide (P-peptide) p-polyphenol, 126
127 p-toluene sulfonic acid (PTS), 326 p53 tumor suppressor protein (p53), 72
73, 163
164 PA. See Paracetamol (PA) PAB system. See Potentiometric alternating biosensing system (PAB system) PAD. See Paper analytical device (PAD) Paget’s disease of nipple, 74 PAI-1. See Plasminogen activator inhibitor 1 (PAI-1) Palladium (Pd), 383
384 Palladium nanoparticles (PdNPs), 93, 379t PAMAM. See Polyamidoamine (PAMAM) PANI. See Polyaniline (PANI) Paper analytical device (PAD), 283
284 Paper working electrode (PWE), 300
301 Paracetamol (PA), 286 Parasites, 2 PASA. See Poly(aniline(sulfonic acid)) (PASA) Pathogen-free food and beverages, 1 Pathogens, 1 PBMC. See Peripheral blood mononuclear cells (PBMC) pCBAA. See Poly(carboxybetaine acrylamide) (pCBAA)

410 PCR. See Polymerase chain reaction (PCR) PCSK9. See Proprotein convertase subtilisin/kexin type 9 (PCSK9) PCTE membranes. See Polycarbonate track-etched membranes (PCTE membranes) Pd wormlike nanochains/graphitic carbon nitride (Pd WLNCs/ g-C3N4), 302
303 PDA. See Polydiacetylene (PDA) PDDA. See Poly-diallyl dimethylammonium chloride (PDDA) PDGF. See Platelet-derived growth factor (PDGF) PDGF-BB. See Platelet derived growth factor BB (PDGF-BB) PDMS. See Polydimethylsiloxane (PDMS) PdNPs. See Palladium nanoparticles (PdNPs) Pdots. See Polymer dots (Pdots) PEDOT. See Poly(3,4ethylenedioxythiophene) (PEDOT) PEG. See Polyethylene glycol (PEG) PEI. See Polyethyleneimine (PEI) Pencil graphite electrode (PGE), 10
12, 299 Peptide nucleic acid (PNA), 8, 82
83 Peripheral blood mononuclear cells (PBMC), 264
265 Peroxidase, 105
106, 127, 127f Peroxide-based biosensors, 366 Peroxiredoxin, 113 PERVs. See Porcine endogenous retrovirus (PERVs) Pesticide, 241 PET. See Positron emission tomography (PET) PF-4. See Platelet factor-4 (PF-4) PG. See Protein G (PG) PGE. See Pencil graphite electrode (PGE) Phage display method, 266 Pharmaceutical detection, 286 pHEMA. See Poly (2-hydroxyethyl methacrylate) (pHEMA) Phenolic compounds, 115 Phosphatase and TENsin homolog (PTEN), 72
73

INDEX

Phosphorylated peptide (P-peptide), 305 Photoelectrochemical (PEC) biosensor, 298
299, 298f dual-mode, 330 enzyme
sensing method, 288
289 gene biosensor, 289
290 materials, 193
194 sensors, 279, 305
307 Photogenerated holes, 305
307 Photoluminescence biotransducers, 148
149 Physiological fluids, 209 Phytochemicals, 115 Piezoelectric biosensor, 20
21 crystals, 63 detection, 238 quartz crystals, 26 PIP, 286
287 PKA. See Protein kinase A (PKA) PLA. See Polylactic acid (PLA) Plants antioxidant capacity assessment, 115
120 Plasminogen activator inhibitor 1 (PAI-1), 76 Platelet derived growth factor BB (PDGF-BB), 303
304 Platelet factor-4 (PF-4), 281 Platelet-derived growth factor (PDGF), 244 Platinized Pt electrode, 328
329 Platinum (Pt), 10, 112
113, 174, 383
384 nanoparticles, 331
332, 379t Platinum chromium alloy (PtCr), 177 PLLA NP. See Poly-L-lactide nanoparticles (PLLA NP) 2PMAB. See (Z)-2-((Pyridin-2-yl) methyleneamino) benzenethiol (2PMAB) PNA. See Peptide nucleic acid (PNA) POCT. See Point-of-care testing (POCT) Point-of-care based smartphone platform, 343
344 Point-of-care testing (POCT), 280
281, 346
349 Poliovirus, 269 Poly (2-hydroxyethyl methacrylate) (pHEMA), 258
259 Poly (N [3-(trimethoxysilyl)propyl] aniline) (PTMSPA), 327
328

Poly-1,5-diaminonaphthalene (P1,5DAN), 151
156 Poly-diallyl dimethylammonium chloride (PDDA), 156
157, 303
304 Poly-L-lactide nanoparticles (PLLA NP), 299 Poly-vinylpyrrolidone (PVPP), 162 Poly(2-hydroxyethyl methacrylate-coN,N-dimethylaminoethyl methacrylate) (HEMA-coDMAEMA), 257 Poly(2-hydroxyethyl methacrylate) [poly(HEMA)], 15
17 Poly(3,4-ethylenedioxythiophene) (PEDOT), 88
91 Poly(aniline(sulfonic acid)) (PASA), 112 Poly(carboxybetaine acrylamide) (pCBAA), 15
17, 27
29 Poly(p-aminothiophenol)-linked gold nanoparticles, 182 Poly(vinyl chloride), 214
216 Poly(vinylidene fluoride-cohexafluoropropylene) (PVDF-HFP), 177 Polyamidoamine (PAMAM), 304
305 Polyaniline (PANI), 110, 190, 217, 326
327 Polycarbonate, 22 Polycarbonate track-etched membranes (PCTE membranes), 22 Polydiacetylene (PDA), 9 Polydimethylsiloxane (PDMS), 22, 257, 280
281 Polyethylene glycol (PEG), 29, 251 Polyethyleneimine (PEI), 190, 364
365 Polyethylenimine, 26
27 Polyglutamic acid/carbon nanotubes (PGA/CNTs/anti-CA15
3/ Ab), 151
156 Polylactic acid (PLA), 22 Polymer dots (Pdots), 290 Polymer nanomaterials, 388
393 Polymer-based nanocomposites, 186
187 Polymerase chain reaction (PCR), 3, 52, 79, 148
149, 228, 284 Polymerization, 179 packed bed, 389
390 Polymers, 151

411

INDEX

Polyphenol oxidase (PPO), 194, 285
286 Polyphenols, 105
106, 135 analysis in beverage, 120
135, 122t electrochemical biosensing of, 125
135 Polypropylene microfiber membranes, 25 Polypyrrole (PPy), 10
12, 112
113, 151
156, 189
190 Polyvinyl alcohol (PVA), 19 Polyvinyl butyral (PVB), 214
216 Polyvinyl chloride (PVC), 216
217 Polyvinylidene fluoride (PVDF), 22 Porcine endogenous retrovirus (PERVs), 269 Porous nanostructures, 383 Positron emission tomography (PET), 267 Postexposure prophylaxis, 271 Potassium ferricyanide (K3[Fe (CN)6]32), 327
328 Potassium peroxydisulfate (K2S2O8), 303 Potentiometric aptamer-based bioreceptor, 13 biosensor, 79, 85 detection, 62 nanobiosensors, 236 potentiometric-based biosensors, 12
13 sensor, 214
216, 216f Potentiometric alternating biosensing system (PAB system), 12 PPO. See Polyphenol oxidase (PPO) PPy. See Polypyrrole (PPy) PPY
MWCNT/ITO nanocomposite electrode, 326 PR. See Progesterone (PR) Precision agriculture. See Smart agriculture Predictive biomarkers, 148 prim. See Primuletin (prim) Primary antibody, 149 Primuletin (prim), 110
111 Progesterone (PR), 75, 89t, 92
93, 149
151 Prognostic biomarkers, 148 Proprotein convertase subtilisin/kexin type 9 (PCSK9), 175 Prostate-specific antigen (PSA), 195
196, 281 Protease system, 77

Protein G (PG), 18 Protein kinase A (PKA), 305 Protein-based electrochemical biosensors, 256 Protein-based insulin delivery systems, 257
258 Proteins, 83, 232, 268 PSA. See Prostate-specific antigen (PSA) Pseudomonas aeruginosa, 273
274 Pt nanowires (PtNWs), 160
162 Pt/G-quadruplex/hemoglobin, 289
290 PtCr. See Platinum chromium alloy (PtCr) PTEN. See Phosphatase and TENsin homolog (PTEN) PTMSPA. See Poly (N [3(trimethoxysilyl)propyl]aniline) (PTMSPA) PtNWs. See Pt nanowires (PtNWs) PTS. See p-toluene sulfonic acid (PTS) PVA. See Polyvinyl alcohol (PVA) PVB. See Polyvinyl butyral (PVB) PVC. See Polyvinyl chloride (PVC) PVDF. See Polyvinylidene fluoride (PVDF) PVDF-HFP. See Poly(vinylidene fluoride-cohexafluoropropylene) (PVDFHFP) PVPP. See Poly-vinylpyrrolidone (PVPP) PWE. See Paper working electrode (PWE) (Z)-2-((Pyridin-2-yl) methyleneamino) benzenethiol (2PMAB), 331
332

Q QCM. See Quartz crystal microbalance (QCM) QCMA-1 device, fully automated, 21f Quantum confinement effect, 287 Quantum dots (QDs), 4, 5t, 24
25, 87, 194
195, 287
294, 382 electrochemical enzyme biosensing, 287
289, 288f electrochemical gene biosensing, 289
292 electrochemical immunosensing, 293
294 QD
based FRET, 148
149

Quartz crystal microbalance (QCM), 63, 238 sensor, 20 Quartz crystals, 238 Quercitin (quer), 110
111, 122t

R rAbs. See Recombinant antibodies (rAbs) Radioimmunoassay, 148
149 Radiolabeled mAbs, 267 Raman spectroscopy, 18
19, 295 Randles
Sevcik equation, 368 Reactive nitrogen species (RNS), 108 Reactive oxygen and nitrogen species (RONS), 108 Reactive oxygen species (ROS), 105
106, 115, 120
121, 132 biosensors for, 108
115 biosensors for H2O2, 113
115 hydroxyl radical, 109
111 superoxide anion radical, 112
113 Reader device detection, 24
25 Ready-to-eat products, 1 Real-time electrochemical profiling (REP), 286 Receptor, 172 Recombinant antibodies (rAbs), 6 Recombinase polymerase amplification (RPA), 18 Reduced graphene oxide (rGO), 174, 219
220 nanocomposites, 293 Refractive index (RI), 15 REP. See Real-time electrochemical profiling (REP) Repressor proteins, 286
287 Reproducibility, 79
80 Resonance energy transfer (RET), 304
305 Respiration monitoring, 217
218 Respiratory syncytial virus (RSV), 269
270 Responsive hydrogels, 254, 255f Resveratrol, 128
131 RET. See Resonance energy transfer (RET) Reverse transcriptasepolymerasechain reaction (RTPCR), 3, 264
265 rGO. See Reduced graphene oxide (rGO)

412 RI. See Refractive index (RI) Ribose-based biosensor, 366 RNS. See Reactive nitrogen species (RNS) RONS. See Reactive oxygen and nitrogen species (RONS) ROS. See Reactive oxygen species (ROS) Rotavirus, 269, 271 RPA. See Recombinase polymerase amplification (RPA) RSV. See Respiratory syncytial virus (RSV) RT-PCR. See Reverse transcriptasepolymerasechain reaction (RTPCR) Ruggedness, 388
389

S SA. See Stearic acid (SA) Saccharides, 210
212 glucose, 210
212 lactose, 212 saccharides-based biosensor, 364 Salmonella sp., 2, 45
46 S. bongori, 2 S. enterica, 2 S. enteritidis, 2 S. typhimurium, 2 S. typhimurium, 45
46 Salt-induced aggregation, 238
239, 239f Salycilc acid, 122t sAMPs. See Synthetic antimicrobial peptides (sAMPs) SAMs. See Self-assembled monolayers (SAMs) Sandwich immunoassay using magnetic beads, 151, 152f “Sandwich-like” biosensor, 88
91 “Sandwich” sensing system, 244, 244f scFv. See Single chain fragment variable Ab fragments (scFv) SCO. See Sterptomyces hygroscopicus (SCO) Screen printed carbon electrode (SPCE), 158
159 Screen printed graphite electrode (SPGE), 151
156 Screen-printed electrodes (SPEs), 93
95, 300, 330 Seafood, 1

INDEX

SEB. See Staphylococcal enterotoxin B (SEB) SEER. See Surveillance, and Epidemiology and End Results (SEER) Selective electrochemical biosensors, 112 Selectivity, 252
254, 389 SELEX. See Systematic evolution of ligands by exponential enrichment (SELEX) Self-assembled DNA concatamer, 96
97 Self-assembled inorganic nanorods, 175
176 Self-assembled monolayers (SAMs), 6, 322 Self-assembled nanoparticle superlattices, 175 Self-assembled nanorods, 172
173 Self-assembly method, 389 Semiconducting nanoclusters, 376
378 Semiconductor QDs, 194 Sensitive disease-relevant biomarkers detection, 281 Sensitivity, 80, 252 Sensors, 209, 376 for health and agricultural applications future prospects, 352
353 global market for smart sensors, 351
352 smart agriculture sensors, 349
350 smart health sensors, 343
349 Sequencing analysis, 3 Serine protease, 364 SERS. See Surface-enhanced Raman scattering (SERS) SERS-LF. See Surface-enhanced Raman scattering
based lateral-flow (SERS-LF) SEs. See Staphylococcal enterotoxins (SEs) SET. See Single electron transfer (SET) SFP. See Staphylococcal food poisoning (SFP) Shear deformation, 384
385 Shiga toxin, 45
46, 49 Shiga toxin-producing Escherichia coli (STEC), 272 Shiga toxins (Stxs), 1
2

Shigella boydii, 2 Shigella dysenteriae, 2 Shigella flexneri, 2 Shigella sonnei, 2 SiC. See Silicon carbide (SiC) Signal amplification, 193 by multistep redox reactions, 163
164, 164f strategies of electrochemical immunosensors, 151, 157f strategy, 182 Signal transducers, 315 Silica, 176 matrix, 187 silica-PANI bienzyme cholesterol biosensor, 327
328 Silicon, 351
352 atom, 295 nanocrystals, 217 Silicon carbide (SiC), 295 Silicon nanowires (SiNWs), 382
383 Silver (Ag), 10 Silver nanoparticles (AgNPs), 4, 110
111, 173, 379t, 383
384 Single chain fragment variable Ab fragments (scFv), 158
159 Single electron transfer (SET), 115
116 Single-stranded DNA (ssDNA), 7
8, 190, 284 Single-walled carbon nanotubes (SWCNTs), 13, 26
27, 114, 210
212, 280, 357 SiNWs. See Silicon nanowires (SiNWs) SiO2-coated magnetic Fe3O4 dispersed multiwalled carbon nanotubes (Fe3O4@SiO2/MWNT), 332
333 Size-exclusion membranes, 253 Small molecule detection, microfluidics chips for, 285
287 Smart agriculture, 349
350 sensors, 349
350 Smart farming, 349
350 Smart health sensors, 343
349 Smart sensors, 343 global market for, 351
352 Smartphone-based spectrometer, 350, 352f SOD. See Superoxide dismutase (SOD) Sodium ion, 216
217 Sodium nitrite (NaNO2), 302

413

INDEX

Sodium tetra kis[3,5-bis (trifluoromethyl)phenyl]borate (Na-TFPB), 216
217 Sol-gel method, 179, 184 coating method, 21
22 Solar panels, 342
343 Sol
gel matrix, 322
323 Solid smartphone-based spectral analyzer, 350 Solid-state potentiometric sensing technology, 214
216 sensors, 214
216 Soluble SWCNT derivatives, 361 SPCE. See Screen printed carbon electrode (SPCE) Spectrophotometric methods, 125 SPEs. See Screen-printed electrodes (SPEs) SPGE. See Screen printed graphite electrode (SPGE) SPR. See Surface plasmon resonance (SPR) SPRi. See Surface plasmon resonance imaging (SPRi) SPW. See Surface plasmon wave (SPW) Square wave voltammetry (SWV), 149 Square-wave anodic stripping voltammetry (SWASV), 26
27 ssDNA. See Single-stranded DNA (ssDNA) Staphylococcal enterotoxin B (SEB), 52 Staphylococcal enterotoxins (SEs), 49
50 Staphylococcal food poisoning (SFP), 49
50 Staphylococcus aureus, 49
50 Stearic acid (SA), 326
327 STEC. See Shiga toxin-producing Escherichia coli (STEC) Steric hindrance, 287 Sterptomyces hygroscopicus (SCO), 316 STP-ALP. See Streptavidin
alkaline phosphatase (STP-ALP) Streptavidin horseradish peroxidase (Strep-HRP), 191
192 Streptavidin
alkaline phosphatase (STP-ALP), 159
160 Streptococcus pneumoniae, 25
26 Stxs. See Shiga toxins (Stxs) Subcutaneous immunization, 264
265 Subgroup J of avian leukosis viruses (ALVs-J), 301
302



Superoxide anion (O22 ), 112 biosensors for, 112
113 superoxide anion radical, 112
113 Superoxide dismutase (SOD), 112
113 Superoxide enzyme, 132
133 Surface functionalization, 195
196 Surface plasmon resonance (SPR), 15
17, 190
191, 195
196, 233
234, 345 aptasensors, 57
58 biotransducers, 148
149 determination, 345
346 Surface plasmon resonance imaging (SPRi), 195
196 Surface plasmon wave (SPW), 15 Surface poisoning, 252 Surface villous Au nanocage (SVAu nanocage), 300
301 Surface-enhanced Raman scattering (SERS), 27
29, 173, 195
196 aptasensors, 57
62 Surface-enhanced Raman scattering
based lateral-flow (SERS-LF), 18 Surfactant-assisted emulsion process, graphite exfoliation via, 359
360 Surveillance, and Epidemiology and End Results (SEER), 78 SVAu nanocage. See Surface villous Au nanocage (SVAu nanocage) SWASV. See Square-wave anodic stripping voltammetry (SWASV) SWCNTs. See Single-walled carbon nanotubes (SWCNTs) Sweat biomarkers, 224
225 and exhaled breath gas sensors, 209
210 SWV. See Square wave voltammetry (SWV) Synthetic antimicrobial peptides (sAMPs), 14 Synthetic material
based glucose sensor, 258 Synthetic methodologies, 376
378 Systematic evolution of ligands by exponential enrichment (SELEX), 8, 9f, 55
56, 56f, 58f, 92
93, 230, 230f whole-cell, 56, 57f

T T-cut quartz crystal sensor, 321
322 Tattoo-based wearable system, 214 TBBPA. See Tetrabromobisphenol-A (TBBPA) TCGA. See The Cancer Genome Atlas (TCGA) TCNQ. See Tetracyanoquinodimethane (TCNQ) TCVD process. See Thermal chemical vapor deposition process (TCVD process) TEA. See Triethanolamine (TEA) TEAC. See Trolox equivalent antioxidant capacity assay (TEAC) Teflon-LOX-POD biosensors, 253 Tenascin-C (TN-C), 163 TEPA. See Tetraethylene pentamine (TEPA) TET-binding aptamer. See Tetracycline-binding aptamer (TET-binding aptamer) TetR, 286
287 Tetrabromobisphenol-A (TBBPA), 301
302 Tetracyanoquinodimethane (TCNQ), 235
236 Tetracycline-binding aptamer (TETbinding aptamer), 305
307 Tetraethylene pentamine (TEPA), 162 Tetrahydrofuran (THF), 216
217 3,30 ,5,50 -Tetramethylbenzidine (TMB), 93
95, 241 Tetrathialfulvalene (TTF), 235
236 Tetrazine derivatives, 194 TGA. See Thioglycolic acid (TGA) TGFα. See Transforming growth factor (TGFα) Th/PVPP/GR modified GCE surface, 162 The Cancer Genome Atlas (TCGA), 78 Theaflavin, 122t Theaflavin-3-gallate, 122t Theaflavin-3,3’-digallate, 122t Thermal chemical vapor deposition process (TCVD process), 362 Thermal expansion reduction, 296 THF. See Tetrahydrofuran (THF) Thioglycolic acid (TGA), 322

414 Thionine (Th), 301
302 Th-mpg-C3N4, 301
302 Thiophene-3-boronic acid, 329
330 Thioridazine, 135 Three-dimension (3D) biomimetic electrochemical biosensor, 114 nanostacks, 330 polymeric materials, 254 3D structured graphene nanosheets catalyst-free production, 362f CVD, 362 exfoliation process, 357
361 microwave irradiation, 361 synthesis, 357
362 3D-structured photonic quartz carbon inverse opal rods, 333 3D graphene/carbon nitride/layered double hydroxide hybrid (Gr-LDH-C3N4), 365
366 Threshold Immunoassay System, 12 Thrombotic thrombocytopenic purpura (TTP), 268 Thyroid transcription factor 1 (TTF-1), 234
235 Ti-based metal
organic frameworks (Ti
MOFs), 173
174 TiO2 nanotubes arrays (TNA), 195 TiO2-NWs, 330 TIR. See Total internal reflection (TIR) Tissue plasminogen activator (tPA), 163, 165 Tissues, 83 Titanium-oxo clusters, 172
174 TMB. See 3,30 ,5,50 Tetramethylbenzidine (TMB) TN-C. See Tenascin-C (TN-C) TNA. See TiO2 nanotubes arrays (TNA) TNBC. See Triple-negative breast cancer (TNBC) Tobacco, 72 Top
down processes, 384 Total internal reflection (TIR), 17 Total oxidant scavenging capacity (TOSC), 115
116 Total radical-trapping antioxidant parameter (TRAP), 115
116 tPA. See Tissue plasminogen activator (tPA) Tracer antibody (Ab2), 149, 151 Tracking drug efficiency, 209

INDEX

Traditional microbiological antibiotic tests, 286
287 Traditional microfluidics fabrication methods, 285 Traditional mycotoxin detection methods, 286 Transducer, 10
22, 172 in biosensors, 56
57 electrochemical biosensors, 10
15 elements, 279 optical biosensors, 15
19 technology. See also Aptamer— technology electrochemical biosensors, 85
86 magnetic biosensors, 86 optical biosensors, 84
85 piezoelectric biosensors, 86 thermometric biosensors, 86 Transducing mechanism, 4, 10 Transduction method, 172 principles, 229 Transferrable smartphone-based device, 344
345, 347f Transforming growth factor (TGFα), 157
158 Transmembrane proteins, 266
267 TRAP. See Total radical-trapping antioxidant parameter (TRAP) Triethanolamine (TEA), 293, 303 Triethylamine (Et3N), 303 Triple-negative breast cancer (TNBC), 73 Triton X-100, 318, 322 Trolox equivalent antioxidant capacity assay (TEAC), 115
116 Tryptophan-functionalized grapheme nanocomposites, 296
297 TTF. See Tetrathialfulvalene (TTF) TTF-1. See Thyroid transcription factor 1 (TTF-1) TTP. See Thrombotic thrombocytopenic purpura (TTP) Tubular nanostructures, 383 Tumor cells, 267 Tumor protein 53 (p53), 77 Tumor-to-blood ratio, 267 Tungsten oxide (WO3), 218
219, 219f, 221f Two-dimensional graphene, 177

Type I diabetes, 249
250 Typhoidal illness, 46
49 Tyrosinase (Tyr), 116
117, 127
128, 127f, 156, 285
286

U UA. See Uric acid (UA) UC. See Upconversion (UC) UCPs. See Upconverting nanoparticles (UCPs) Ultrafine particles (UFP), 382 Ultrasound, 86
87 United States Department of Agriculture and Food Safety and Inspection Service (USDA-FSIS), 2, 49 uPA. See Urokinase plasminogen activator (uPA) Upconversion (UC), 194
195 Upconverting nanoparticles (UCPs), 4, 25, 93 Uric acid (UA), 213, 296 Urokinase plasminogen activator (uPA), 76

V Valence band (VB), 293, 305 Vanillic acid, 122t Variable domain (VH), 263
264 Variable heavy chain (VHH), 263, 264f Vascular endothelial growth factor (VEGF), 251
252, 299 Vascular endothelial growth factor receptors (VEGFRs), 160 VB. See Valence band (VB) VEGF. See Vascular endothelial growth factor (VEGF) VEGFRs. See Vascular endothelial growth factor receptors (VEGFRs) VH. See Variable domain (VH) VHH. See Variable heavy chain (VHH) Viable but nonculturable microbial pathogenic strains (VBNC microbial pathogenic strains), 50 Vibrio spp., 2, 50 V. alginolyticus, 13 V. cholerae, 2, 50 V. parahaemolyticus, 2 V. vulnificus, 2 Viruses, 269 nanobodies against, 269
272

415

INDEX

Vishay photo IC sensor, 351f Visual detection, 23
24 Vitamins, 105
106 Volatile organic compounds (VOCs), 210 Voltammetric techniques, 135

W Warfarin sodium (WFS), 286 Waspmote Plug & Sense! unit, 342
343, 343f WBAN. See Wireless body area network (WBAN) Wearable biosensor applications, 210
212 Wearable gadgets, 341

Wearable sweat sensors, 210
217 detection of ions, 214
217 detection of organic compounds, 212
214 detection of saccharides, 210
212 Wearable wireless personal/body area networks (WWBANs), 342, 342f WFS. See Warfarin sodium (WFS) WHO. See World Health Organization (WHO) Whole-cell SELEX, 56, 57f Wine, 121 polyphenols in, 131 Wireless body area network (WBAN), 341
342

World Health Organization (WHO), 45
46, 106, 147
148, 227 WWBANs. See Wearable wireless personal/body area networks (WWBANs)

X Xanthine, 256 (Z)-2-((Xhexanedithiol (HDT), 159
160 XPS spectroscopy, 295

Z Zinc oxide nanoparticles (ZnONPs), 160
162 Zirconium oxo-cluster, 173
174