Electrochemistry for Bioanalysis [1 ed.] 0128212039, 9780128212035

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Electrochemistry for Bioanalysis

Electrochemistry for Bioanalysis

Bhavik A. Patel

Professor of Clinical and Bioanalytical Chemistry, School of Pharmacy and Biomolecular Sciences, Centre for Stress and Age-Related Disease, University of Brighton, Brighton, UK

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 © 2021 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: 978-0-12-821203-5 For Information on all Elsevier publications visit our website at https://www.elsevier.com/books-and-journals Publisher: Gayathri S Acquisition Editor: Kathryn Eryilmaz Editorial Project Manager: Billie Jean Fernandez Production Project Manager: Bharatwaj Varatharajan Cover Designer: Christian J. Bilbow Typeset by Aptara, New Delhi, India

Contents

Contributors ix 1

Introduction to electrochemistry for bioanalysis 1 Bhavik A. Patel 1.1 Introduction 1 1.2 Bioanalysis 1 1.3  Principles of electrochemistry 4 Summary 8 Further reading 8

2

Amperometry and potential step techniques 9 Bhavik A. Patel 2.1 Introduction 9 2.2 Principles 10 2.3  Strengths and limitations 14 2.4 Applications 16 2.5 Summary 25 References 26

3 Voltammetry 27 B. Jill. Venton, Dana J. DiScenza 3.1 Introduction 27 3.2 Principles 28 3.3  Strengths and limitations 30 3.4 Applications 33 3.5 Summary 46 References 46 4

Microelectrodes and nanoelectrodes 51 Bhavik A. Patel 4.1 Introduction 51 4.2  Carbon fiber microelectrodes 53 4.3  Microelectrode arrays 63 4.4 Nanoelectrodes 67 4.5 Summary 70 References 70

vi Contents

5

Novel sensing materials and manufacturing approaches 73 Bhavik A. Patel 5.1 Introduction 74 5.2  Novel carbon materials for generation of electrodes 74 5.3  Carbon composite electrodes 88 5.4  3D printing for development of electrodes 92 5.5 Summary 97 References 97

6

Experimental design – challenges in conducting electrochemical measurements for bioanalysis 99 Bhavik A. Patel 6.1  Key factors that influence bioanalytical measurements 99 6.2  Electrode and instrumentation variables 102 6.3  Experimental variables 114 6.4  Biological environment 118 6.5 Summary 122 References 122

7

Electrochemistry at and in single cells 125 Alex S. Lima, Chaoyi Gu, Keke Hu, Andrew G. Ewing 7.1 Introduction 125 7.2  General introduction of exocytosis 126 7.3  Basic history at electrochemistry at/in cells 127 7.4  Electrodes for single cell and subcellular analysis 128 7.5  Cellular techniques to study exocytotic neurotransmitter release 129 7.6  Dynamics of exocytotic release revealed through interpretation of single-cell amperometric data 132 7.7  Modeling exocytosis and closing of the fusion pore 134 7.8  Applications of amperometry in neuroscience research 136 7.9  Intracellular electrochemistry 137 7.10  Measurements of reactive oxygen and nitrogen species (ROS/RNS) at/in single cells 143 7.11  Enzyme-based electrodes for single cell analysis 150 7.12  Scanning electrochemical microcopy (SECM) at single cells 152 7.13 Summary 154 References 154

8

Measurement from ex vivo tissues Bhavik A. Patel 8.1 Introduction 8.2  Ex vivo tissues – what are they? 8.3  Experimental considerations for measuring ex vivo tissues 8.4  Studies conducted using ex vivo tissues 8.5  Measurements from ex vivo organs from simple biological models

161 161 162 164 168 185

Contents

vii

8.6  Future directions 190 8.7 Summary 190 References 191 9

In vivo electrochemistry 195 Aya Abdalla 9.1 Introduction 195 9.2  What are in vivo measurements? 196 9.3  Strengths and limitations of in vivo experimentation 196 9.4  Criteria for ideal in vivo measurements 197 9.5  Electrochemical techniques 198 9.6  Experimental optimization for acute and chronic in vivo measurement 201 9.7 Measurements in vivo 208 9.8  Measurements in different regions of the body 216 9.9  Summary and future directions 216 References 217

10 Measurement in biological fluids 223 Bhavik A. Patel 10.1 Introduction 223 10.2  Different biological fluids 224 10.3 Blood 224 10.4 Urine 234 10.5 Saliva 237 10.6 Sweat 238 10.7  Interstitial fluid 240 10.8  Tear fluid 240 10.9  Future directions 242 10.10 Summary 243 References 243 11 Measurement of reactive chemical species 247 Bhavik A. Patel 11.1 Introduction 247 11.2  Reactive oxygen species (ROS) 248 11.3  Reactive nitrogen species (RNS) 249 11.4  Role of ROS/RNS in biology 251 11.5  Electrochemistry of ROS/RNS 252 11.6  Electrode modifications to measure ROS/RNS 254 11.7  Measurement of reactive species from biological environments 260 11.8 Summary 264 References 264

viii Contents

12 Electrochemical biosensors 267 Bhavik A. Patel 12.1 Introduction 267 12.2  Types of enzymatic biosensors 268 12.3  Immobilization of enzymes on electrode surfaces 272 12.4  Factors that influence the performance of biosensor measurements 274 12.5  Application of biosensors 277 12.6 Summary 282 Further reading 283 References 283 13 Electrogenerated chemiluminescence (ECL) 285 Andrew Danis, Janine Mauzeroll 13.1  Electrogenerated chemiluminescence introduction 285 13.2  Electrochemistry and ECL 288 13.3  Electron transfer theory and ECL 291 13.4  ECL history 294 13.5  ECL instrumentation 299 13.6  ECL simulation 301 13.7  ECL materials development 303 13.8  Conclusions and perspectives 306 References 306 Index 315

Contributors

Bhavik A. Patel School of Pharmacy and Biomolecular Sciences, Centre for Stress and Age-Related Disease, University of Brighton, Brighton, UK B. Jill. Venton Department of Chemistry, University of Virginia, Charlottesville, Virginia, United States Dana J. DiScenza Department of Chemistry, University of Virginia, Charlottesville, Virginia, United States Alex S. Lima Department of Chemistry and Molecular Biology, University of Gothenburg, Gothenburg, Sweden Chaoyi Gu Department of Chemistry and Molecular Biology, University of Gothenburg, Gothenburg, Sweden Keke Hu Department of Chemistry and Molecular Biology, University of Gothenburg, Gothenburg, Sweden Andrew G. Ewing Department of Chemistry and Molecular Biology, University of Gothenburg, Gothenburg, Sweden Aya Abdalla School of Pharmacy and Biomolecular Sciences, Centre for Stress and Age-Related Disease and Surface Analysis Laboratory, School of Environment and Technology, University of Brighton, Brighton, East Sussex, UK Andrew Danis Department of Chemistry, McGill University, Montreal, Québec, Canada Janine Mauzeroll Department of Chemistry, McGill University, Montreal, Québec, Canada

Introduction to electrochemistry for bioanalysis

1

Bhavik A. Patel School of Pharmacy and Biomolecular Sciences, Centre for Stress and Age-Related Disease, University of Brighton, Brighton, UK

Keypoints Principles Electrochemistry is the study of electron movement in an oxidation or reduction reaction at a polarized electrode surface. Each analyte is oxidized or reduced at a specific potential and the current measured is proportional to concentration. This technique is a powerful methodology towards bioanalysis.

Applications in bioanalysis Electrochemistry is widely used for measurement of a wide range of analytes (neurotransmitters, large biomolecules and biological gases) in a host of different biological areas. ■ Due to the ability to use a wide range of techniques and sensors, measurements can be made in vivo, ex vivo or in vitro. ■

1.1 Introduction In bioanalysis, measurement of biomolecules is an incredible challenge given that dynamic changes in small concentrations of chemicals occur in specific yet complex environments. Therefore, in order to meet these challenges any techniques must have high levels of sensitivity, selectivity, temporal resolution and spatial resolution. To date, electrochemical techniques have provided the answer to all these requirements using an array of different sensing platforms that can measure in the smallest of locations at real-time. With enhancement of materials and processes in which electrode reactions can be explored, the ability to monitor far and wide has opened new applications and knowledge through the application of electrochemical techniques in bioanalysis.

1.2 Bioanalysis Bioanalysis is a discipline of analytical chemistry, where measurement is focused towards the detection and determination of chemicals that play significant roles in biology. Various chemicals, from small reactive gaseous species to large biomolecules such as proteins and DNA are often widely considered for determination. The Electrochemistry for Bioanalysis. DOI: 10.1016/C2019-0-03108-1 Copyright © 2021 Elsevier Inc. All rights reserved.

2

Electrochemistry for Bioanalysis

Non-invasive

Invasive

Sali va

Non i n vas ive Inva sive

th ea Br

Tr de ansrm al

Blo od

eat Sw

sive nva n-i No

Urine &

Sto ol

sive nva i n No

Other organs

ve vasi n-in No

ain Br

rs Tea

Noninva siv e

Tissue

Non -inv asi ve

Fig. 1.1.  Schematic indicating the different bodily samples. Source: Booth, M. A.; Gowers, S. A. N.; Leong, C. L.; Rogers, M. L.; Samper, I. C.; Wickham, A. P.; Boutelle, M. G., Chemical Monitoring in Clinical Settings: Recent Developments toward Real-Time Chemical Monitoring of Patients. Analytical Chemistry 2018, 90 (1), 2–18.

measurement of such an array of different biological chemicals can provide significant insight into our understanding of the role of important chemicals in the body and how these may be altered in disease. To conduct bioanalysis, measurements need to be performed in specific biological environments. Fig. 1.1 shows the varying biological environments that chemical analysis can be conducted in and how they vary with regards to the invasiveness or non-invasiveness of the environments. Invasive measurements are where the recordings are conducted in the biological environment, which is also known as an in vivo measurement. This will be explored at depth in Chapter 9. Alternatively, non-invasive measurements can be used to measure chemicals in samples that can be obtained from the body such as tears, saliva, sweat, exhaled breath, urine, and stool samples. However, some biological environments such as tissues and blood have both been utilized for invasive and non-invasive measurements. Tissues can be extracted from the body and preserved for measurement. This is known as ex vivo measurements and will be explored at depth in Chapter 8.

Introduction to electrochemistry for bioanalysis

3

Invasive measurement poses more significant challenges than non-invasive measurement, as often there is a dynamic change in the biological environment and the measurement requires an approach in which the degree of invasiveness is limited. Invasive measurements provide the ability to conduct measurements and explore real physiological changes in the biological environment over time. For non-invasive measurements, often samples can be collected and measured offline. In these circumstances, the sample is static and only provides a snapshot of the biological environment at the time the sample is collected. Often these measurements are considered useful from a diagnostic point of view as biological samples are easily accessible and can yield vital information about levels of specific chemicals. However, knowing the right sample to take for measurements requires understanding on the precise role of the biomolecules that you intend to monitor, as this has profound implications on the measurement approaches that can be utilized.

1.2.1  Where is my biomolecule? Location and role of the biochemical of interest is critical in order to fully understand how bioanalysis can be conducted. Most biomolecules are synthesized and/or stored within cells and are unidentifiable in all the samples except tissues from Fig. 1.1. However, to obtain measurements from cells, requires the need to rupture the cell membrane to obtain access to the cell cytoplasm for measurement. These are considered intracellular measurements and are highly complex due to the small microenvironments and extremely low concentrations of biomolecules present. Therefore, the majority of bioanalysis is focused on measurement of biomolecules which exit the cell, which commonly occurs to drive communication. A biomolecule can leave the cell by multiple mechanisms of which the most common processes are autocrine, paracrine and endocrine. As shown in Fig. 1.2, these modes all require the release of a biomolecule, but the vicinity in which they play a key role in communication varies. Autocrine Signaling cell

Paracrine Signaling cell

Target cell

Endocrine Signaling cell

CVS

Target cell

Fig. 1.2.  Processes in which biomolecules communicate and therefore are required to leave the cell. Where CVS is the cardiovascular system.

4

Electrochemistry for Bioanalysis

In autocrine communication, a biomolecule is released from a cell and directly binds to receptors on the same cell leading to changes in the function of the cell. Therefore, autocrine communication is highly localized to the vicinity of the cell. Hence, to conduct bioanalysis measurements where autocrine communication is explored, single cell studies need to be conducted. Paracrine communication occurs when a biomolecule is released and binds to receptors on a cell in close vicinity of the releasing cell. In this case, there is a direct communication between two cells though the biomolecule that is released. This form of communication is most widely used by cells to influence key biological functions such as neuronal communication at neuromuscular junctions. In bioanalysis, paracrine signaling is the most widely explored mode of communication, through the development of devices and methodologies. Finally, the last mode of communication is endocrine, where the biomolecule released utilizes the cardiovascular system in order to be transported to reach distant cells in specific organs and therefore provide communication over large distances. Hormones are widely released by this process. Given that this mode of communication utilizes the blood stream to transport the signaling biomolecules, bioanalysis either invasively or non-invasively in blood is the is most common method to explore endocrine signaling other than directly exploring the releasing cell. Of the different processes that govern biomolecule release, paracrine and endocrine signaling are the most widely explored. In both signaling mechanisms, biomolecules can be released either as individual chemicals into the extracellular space or released in small packages. These small packages are collectively known as extracellular vesicles, which are small nanometer sized lipid bilayer packets that are released from all cells within our body and contain a wide range of biomolecule cargo. The challenges of bioanalysis are significant, given the wide array of biomolecules that are released either individually or as a collective in significantly small concentrations over dynamic timescales. Additionally, these biomolecules can be released in incredibly short (nanometers) to large distances (meters) and are found in different biological regions where the environment is varied. Although there are various detection modes that can be utilized to meet these challenges, such as fluorescence, by far the most successful and widely used mode of detection is electrochemical detection. This is namely due to the ability to generate electrodes in any size that can be chemically or biologically modified to meet the needs for measurement of molecules in a host of biological environments. Electrochemical detection has provided significant enhancement in our knowledge of biomolecules and their precise role in biology thus making electrochemical analysis the most important approach to bioanalysis.

1.3  Principles of electrochemistry 1.3.1  The electrochemical reaction The fundamental principle governing electrochemical detection is the movement of electrons between an electronic conductor (usually metal or carbon), and a redox analyte species at the electrode (conductor) surface. Oxidation involves the loss of

Introduction to electrochemistry for bioanalysis

5

electrons from the highest occupied molecular orbital, whereas reduction involves electrons being injected into the lowest unoccupied molecular orbital of the analyte. For some arbitrary pair of compounds where R represents the reduced form and O represents the oxidized form, the electrochemical reaction can be written as:



O + ne- ⇔ R

To either oxidise or reduce an analyte of interest, an electrode potential is applied. The electron potential must be sufficiently positive in order to oxidise an analyte or sufficiently negative to reduce an analyte, therefore the potential at the electrode has significant impact on the form of the analyte at the electrode surface. As electron transfer is a key aspect of the reaction, the flow of electrons (current) from a reaction is the measure of the rate of the analyte’s oxidation or reduction. However, there are a series of processes which can influence this electrochemical reaction. These processes are: (i) mass transfer of analyte from bulk solution to the electrode surface, (ii) electron transfer at the electrode surface, (iii) reactions at the electrode surface such as adsorption and finally (iv) chemical reaction that may precede or occur following the electrochemical reaction. The process of this electrochemical reaction can be observed in Fig. 1.3. Within the process shown in Fig. 1.3, mass transfer and electron transfer are the simplest form of electrochemical reactions and often observed with outer sphere redox species such as ruthenium hexamine. In such process the mass transfer of the species to the electrode surface is via diffusion (movement of a species under the influence of a gradient of chemical concentration) and electron transfer occurs at the electrode surface. The rate of mass transfer to the electrode can be increased by the use of convection or migration (movement of a charged species under the influence of an electric field – i.e. a gradient of electrical potential), which will enhance the current response as more of the analyte reaches the electrode surface during the course of a measurement.

Fig. 1.3.  Process of an electrochemical reaction.

6

Electrochemistry for Bioanalysis

In the case of convection, this is usually through stirring or hydrodynamic transport and is most related to bioanalysis. Generally fluid flow is the cause of natural convection through changes in density gradients. However fluid flow can also be through forced convection, where fluidic movement is caused using stirring or pumps. In bioanalysis of the circulatory system, in which nutrients are pumped throughout the body by the heart, fluid movement is by forced convection and therefore measurements within this biological region are with increased mass transfer. Mass transfer is also increased if the behavior of the tissue is dynamic in nature. This is often the case for most muscular tissues, which convey relaxation and contractions, altering the resulting fluid microenvironment. Therefore, in many circumstances of bioanalysis convection is present and thus enhances the ability to conduct electrochemical measurements. Mass transfer can also be reduced if the electrode surface is modified with polymers or film coatings, as this will reduce the diffusional coefficient of the analyte whilst passing though this medium. Mass transfer is highly varied in biological measurement, due to the environment that the measurement is conducted within, as often geometric locality of the fluid environment defines the rate of analyte diffusion. In bioanalysis, simple diffusion driven electron transfer reactions are not common; more often complex electrochemical reactions which have additional processes occur. These are mainly due to the type of analytes measured, which have preferences in the process through which electron transfer occurs or they generate reactive by-products, which undergo additional chemical reactions. An example of this is the oxidation of serotonin, a biogenic amine which is an important transmitter in the brain and gastrointestinal tract. Serotonin will undergo oxidation following adsorption to the electrode surface and following a two-electron oxidation generate a reactive by-product which further reacts with other molecules of serotonin to generate dimers or trimers. In this case, this electrochemical reaction is not fully reversible. Other analytes such as hydrogen peroxide, which is a major biological reactive oxygen species which in excess can cause damage to cells and tissue, will undergo oxidation on specific electrode surfaces, namely those that facilitate catalytic reactions. The mechanism by which an analyte undergoes a redox reaction at the electrode surface, is an important consideration in development of a robust measurement approach using electrochemical techniques. As such knowledge will help define the type of electrode and measurement waveform that will be utilized, many examples of this careful planning on the measurement conditions are provided throughout the textbook.

1.3.2  The electrochemical cell In order to explore the electrochemical reaction, we need to have an electrochemical cell. A potentiostat is the electronic hardware that is required to control the electrochemical cell and conduct electrochemical measurements. At present, two types of electrochemical cells are employed for bioanalysis. This is either the three-electrode or two-electrode system as shown in Fig. 1.3. Common to both cells are the presence

Introduction to electrochemistry for bioanalysis

7

of a working electrode, in which both the voltage is applied, and the resultant current is measured. This is coupled with a reference electrode that has a known and fixed voltage that is as close to a non-polarized voltage as possible. The voltage applied is not often the voltage observed in the electrochemical cell as shown in equation below:



Eapplied = Ecell – iRs

Where Eapplied is the voltage applied on the potentiostat, of which the voltage is distributed into two parts. Firstly, the Ecell is the voltage applied to the working electrode in order to drive the redox reaction and the remainder of the voltage represent the ohmic drop (iRs) caused by the current passing through a resistive solution. This ohmic drop is characteristic of the bulk solution and thus will vary with electrolyte concentration. For bioanalysis, the ohmic drop is an essential factor in choice of electrochemical cell set-up. In measurements where very small electrodes (nano or microelectrodes) are utilized, and where the current is often in nanoamperes and the background electrolyte concentration is high, there is a minimal contribution of the ohmic drop and therefore a two-electrode electrochemical cell can be utilized. However, in circumstances where the electrolyte concentration within the solution where the bioanalysis is conducted is low and a large electrode (greater than hundreds of micrometers diameter electrode) is utilized, the contribution of the ohmic drop will be significant and therefore a three-electrode system is preferred. In this orientation, the third electrode as shown in Fig. 1.4 is a counter (or auxiliary). It is mainly used to avoid the current passing between the working and reference electrode, through instead having it flow between the counter and working electrodes.

Fig. 1.4.  Three-electrode and two-electrode electrochemical cell. Where Ewk is the potential at the working electrode and Eapplied is voltage applied on the potentiostat.

8

Electrochemistry for Bioanalysis

Summary Bioanalysis is complex due to the wide range of environments that can be explored. There are significant challenges in measurement due to the nature of analytes being present in small concentrations in distinct locations where often multiple other interfering substances are present. Additionally, the timescale that the analytes of interest are present are often short. Only a limited the number of analytical techniques can meet the complex challenges of bioanalysis. Electrochemistry is by far the best of these analytical techniques as an approach to monitor biomolecules. This approach utilized the redox reaction, which is specific for every analyte.

Further reading [1] A. Bard, L.R. Faulkner, Electrochemical Methods: fundamentals and Applications, John Wiley & Sons, New York, 2001.

Amperometry and potential step techniques

2

Bhavik A. Patel School of Pharmacy and Biomolecular Sciences, Centre for Stress and Age-Related Disease, University of Brighton, Brighton, UK

Keypoints Principles In amperometry, a fixed voltage is applied over time. This can be modified by applying more than one fixed voltage, where multiple voltages are applied at varying durations of time. In all cases the resultant current is monitored over time.

Applications in bioanalysis Widely used as the electrochemical technique of choice for the detection of vesicular release ■ Provides the ability to explore signalling processes, such as clearance of transmitter molecules ■ Used in conjunction with enzyme biosensors and separation science techniques ■

Strengths Provides by far the best approach for monitoring dynamic changes in biological signalling, due to the ability to record in real-time ■ Offers excellent sensitivity ■ Provides the ability to conduct rapid monitoring of analytes in an environment outside of the laboratory such as a clinical setting ■

Limitations This detection mode provides no selectivity unless used with a separation science approach ■ Can be influenced by changes in the diffusional layer, thus convection can alter current response ■

2.1 Introduction Amperometry is one of the most widely used techniques in bioanalysis. To-date, it has been utilized in three core areas. These include the detection of chemical transmitters released from cells (highlighted in Chapter 7) and as the mode of detection for various types of biosensors (highlighted in Chapter 12).

Electrochemistry for Bioanalysis. DOI: 10.1016/C2019-0-03108-1 Copyright © 2021 Elsevier Inc. All rights reserved.

10

Electrochemistry for Bioanalysis

Application of amperometry has led to a greater understanding of signalling mechanisms of key chemical transmitters. It has also provided the ability to monitor a wide array of analytes from biological samples when using separation sciences such as chromatography or capillary electrophoresis.

2.2 Principles 2.2.1 Amperometry Amperometry is one of the most widely used detection modes for bioanalysis. In this technique the working electrode is held at a fixed potential (E1), which is sufficient to oxidise or reduce the analyte of interest at the electrode surface and the resultant current is monitored over time. The waveform and expected output are shown in Fig. 2.1. Since a single voltage is applied, the choice of voltage is important. Often this is determined through a voltammetric response where the oxidation or reduction peak potential for the analyte of choice is obtained. As the voltage is applied, initially there is a charging current. Let us assume the charge on an electrode is zero and we suddenly change its potential so that the electrode’s surface acquires a positive charge. Cations near the electrode’s surface will respond to this positive charge by migrating

Fig. 2.1.  Amperometry. (A) shows a voltammetric response of an analytes, where the EpA is the oxidation peak potential and E1 would be the appropriate choice of applied voltage for the (B) waveform applied. The resultant (C) current output initially has a charging current which can hinder the measurement of the faradaic response.

Amperometry and potential step techniques

11

away from the electrode; anions, on the other hand, will migrate toward the electrode. This migration of ions occurs until the electrode’s positive surface charge and the negative charge of the solution near the electrode are equal. Because the movement of ions and the movement of electrons are indistinguishable, the result is a small, shortlived nonfaradaic current that we call the charging current. For amperometry detection, the voltage applied must be slightly positive of the oxidation peak potential or slightly negative of the reduction peak potential of a given analyte in order to ensure that any drift in the response of the reference electrode does not hinder the accuracy of measurement.

2.2.2 Chronoamperometry In chronoamperometry, the voltage is initially held at a resting potential (E1), where neither oxidation nor reduction of any of the analytes of interest can occur. The voltage is then stepped to a value where clear oxidation or reduction of the analyte of interest can occur (E2). The waveform can be observed in Fig. 2.2. When E1 is applied, the charging current contribution can be clearly observed as this is effectively the background current during electrochemical measurements. As this means there is always a non-zero baseline response in electrochemical measurements, this can significantly influence the faradaic measurements conducted. When the electrode is stepped to E2, which allows for complete oxidation or reduction of the analyte of interest, the resultant faradaic current rises instantaneously as the electrode surface concentration of the analyte falls to zero. As the concentration of analyte depletes with increasing time, the current decays with the square root of time (t1/2). For large electrodes, the current is given by the Cottrell equation: i=

nFAc D πt

where i is the current, n is the number of electrons transferred, F is Faraday’s constant, A is the area of the electrode, c is the initial concentration of the analyte, D is the diffusion coefficient of the analyte and t is time.

Fig. 2.2.  Chronoamperometry. (A) waveform applied and (B) expected output, where the jump in the current is observed due to the application of E2.

12

Electrochemistry for Bioanalysis

Fig. 2.3.  Relationship between current (A) and the square root of time ( t ) for macro and microelectrodes.

For microelectrodes (an electrode which has an electrode surface area  200 µm) for the electrochemical determination of catecholamines, using normal pulse polarography [7]. Within the same year, Julian Millar created a microelectrode with a shorter tip (10 −30 µm) for measurements within the central nervous system [8]. From those initial studies to present, carbon fiber microelectrodes are most widely used in bioanalytical measurements using electrochemistry and unraveled new and exciting information about the role of chemicals in biological systems. To-date, studies using carbon fiber electrodes have used fibers with diameters between 5 and 10  µm. Carbon fibers have excellent electrochemical properties, have been widely characterized over the years and have been shown to be biologically compatible (not toxic to cells), making them ideal materials for electrochemical bioanalysis. Additionally, due to their small diameter, they are amenable to implantation and cause significantly less tissue damage than other larger probes used. Carbon fibers are made predominately from organic polymers. About 90%percent of the carbon fibers produced are made from polyacrylonitrile (PAN). The remaining 10 percent are made from rayon or petroleum pitch. The exact composition of the organic polymers varies and thus there is a vast variation between carbon fibers from different manufacturers based on the precursor material and manufacturing process. The process for making carbon fibers is part chemical and part mechanical. The strands and fibers of the organic polymer precursor are heated to a very high temperature in anaerobic conditions. Without the presence of oxygen, the atoms vibrate violently until most of the non-carbon atoms are expelled. This process is called carbonization and leaves a fiber composed of long, tightly interlocked chains of carbon atoms with only a few non-carbon atoms remaining.

4.2.1  Making carbon fiber microelectrodes There are many methods of manufacture of carbon fiber microelectrodes [8–12]. The varying steps in the fabrication of carbon fiber microelectrodes are: a) Preparing the carbon fiber: Carbon fibers can be obtained from varying manufacturers, and all different types have been shown to provide good electrochemical performance. Sometimes the fibers are provided with a thin layer coating of resin or binder that helps to keep the fibers together. This coating can be removed by soaking a bundle of fibers in either acetone or dimethylsulfoxide. Some methods suggest heating the solvent in order

54

Electrochemistry for Bioanalysis Carbon fibre

Electrode blank

Test tube Acetone Plasticine

Forceps

Fig. 4.2.  Inserting the carbon fiber microelectrode into an empty glass capillary. Source: Millar, J.; Pelling, C. W. A., Improved methods for construction of carbon fibre electrodes for extracellular spike recording. Journal of Neuroscience Methods 2001, 110 (1–2), 1–8.

to enhance the efficiency of this process. Removal of this coating is key to enhancing the electrical connectivity of carbon fiber. b) Inserting the carbon fiber into the glass capillary: The first stage of manufacturing, and often the step that is most frustrating, is to extract a single strand of carbon fiber and to insert it into an empty glass capillary. Due to electrostatic forces, multiple fibres are often held together which can make single fiber isolation quite difficult. The carbon fiber needs to be completely inserted into the capillary, which again can be difficult to achieve due to the electrostatic forces. Either using your hands or a fine pair of forceps the carbon fiber can be inserted slowly into the capillary. If this is difficult, then insertion can also be aided using a capillary that is placed in a test tube containing acetone (Fig. 4.2) or by using vacuum suction whilst ensuring the fiber is held in place at one end. The later approach is the most widely utilized. c) Pulling the glass capillary and sealing the carbon fiber: The sealing of the glass capillary around the carbon fiber is achieved using a conventional micropipette puller. Different models of pullers exist which utilize either a heated filament or laser which melt the glass whilst weights gently pull the capillary in two. The settings of the puller need to be optimized to determine the length of the pulled shaft that seals the carbon fiber as well as the length of the seal. After pulling the glass capillary, the carbon fiber should be visibly protruding several centimeters from both ends of the pulled capillary. The fiber should be cut in two carefully using a fine pair of scissors. The seal between the carbon fiber and the glass capillary must be watertight otherwise fluid will leak into the gap (Fig. 4.3). In this case, the fluid would increase the effective surface area of the electrode and will thus cause unpredictable fluctuations in the background signal, leading to greater noise. To enhance the sealing of the carbon fiber with the glass capillary a thin layer of epoxy resin can be applied. d) Tailoring the microelectrode: Once the connection is achieved, the electrode can be modified to make any shape or geometry through a variety of processes, such as cutting to define a specific length of the electrode, etching to create a conical shaped electrode or bevelling to generate a disc electrode.

Microelectrodes and nanoelectrodes Region of tight bonding between carbon and glass

55

Thin glass “shoulder”

Carbon fibre

Fig. 4.3.  Appearance of the carbon fiber after pulling the glass capillary to create a tight seal. Source: Millar, J.; Pelling, C. W. A., Improved methods for construction of carbon fibre electrodes for extracellular spike recording. Journal of Neuroscience Methods 2001, 110 (1–2), 1–8.

e) Making electrical connection with the carbon fiber: Of all the stages in electrode manufacture, poor electric contact is the most likely cause of failing electrodes. There are various approaches to connect the electrode to a metal wire, commonly silver or copper. The simplest approach and potentially the most robust is to back-fill the glass capillary with a highly conductive electrolyte such as 3  M KCl. This conductive electrolyte forms a connection between the carbon fiber and the metal wire. Another widely used approach is to use a conductive silver paint to form a connection between the carbon fiber and the metal wire. Often during measurements, the electrode may become noisy for no clear apparent reason. This is most likely due to contact between the carbon fiber and metal wire becoming disturbed or due to cracks in the seal as mentioned above.

4.2.2  Types of carbon fiber microelectrodes After the microelectrode has been fabricated, often the length and shape of the electrode is altered. This is mainly to tailor the electrode towards the bioanalysis study to be conducted. Fig. 4.4 shows a variety of different geometries that have been fabricated for measurements. Applications of these sensors for measurement from single cell to in vivo will be highlighted in future chapters. The first set of electrodes are achieved through beveling the exposed carbon fiber to the glass, to create an inlaid disc or ellipsoid electrode. The electrode is held on a micromanipulator and gently grinded against a spinning diamond abrasive plate. This is often done in the presence of some oil or alumina, to prevent any friction on

Fig. 4.4.  Varying geometries of carbon fiber microelectrodes utilized for bioanalysis.

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the fragile tip of the electrode. If the electrode is beveled at 90° then a disc electrode will be generated, but for an ellipsoid often angles between 45 to 60° are utilized. This type of electrode has a reduced electrode surface area; however it provides a significant advantage for penetrating into tissues or holding onto the surface of single cells when compared to electrodes where the fiber is exposed. The disc or ellipsoid electrode can be pushed against the surface of a single cell and be used to form an artificial synapse and thus can be used for monitoring release from cells (see Chapter 7 for further details). Given that this electrode has the smallest surface area of all the different geometries of carbon fiber electrodes used, it is not useful where there are low analyte concentrations present or where the electrode surface is used for adsorption of the analyte prior to detection. Protruding tip or conical carbon fiber electrodes are the most used forms of microelectrode and widely used for in vivo measurements. In this type of electrodes, the protruding carbon fiber is cut to a defined length (commonly between 10 – 150 µm) under a microscope using a sharp scalpel. In order to achieve the defined length, a graticule is used. The choice of length depends on the application and the requirement for measurement of chemicals that may be present in significantly lower concentrations. Increased area allows for greater adsorption of analytes for faradaic measurement but also gives increased capacitive signal and therefore an increase in the background noise is observed. Therefore, the right choice of length is critical for enhancing the signal-to-noise ratio. Etched conical electrodes have sharp tips so are ideal for penetration into cells or between cells, thus providing the means for conducting intracellular measurements. Etching can be achieved by either sharp etching or flame etching. In sharp etching, high dc voltage pulses are applied between a tungsten electrode and the carbon fiber electrode, generating micro-sparks that progressively erode the carbon fiber. This approach preserves the diameter of the carbon fiber whilst creating the conical tip shape. In flame etching the carbon fiber is placed within a butane flame for a duration of a few seconds resulting in a conical tip with a significantly reduced electrode diameter. There are other less common types of microelectrodes such as recessed disc electrode. The type of electrode geometry chosen is fundamentally linked to the bioanalysis to be conducted, the sensitivity needed and the spatial location where measurements will be conducted.

4.2.3  Electrochemical behavior of carbon fiber microelectrodes Given that carbon fibers are widely used in bioanalysis, it is important to understand how the carbon fiber structure and surface chemistry affect the analytical response. There are varying approaches that can be widely used to modify the electrode for specific applications so understanding the role of different surface chemistry on electrochemical measurement is important. Within carbon fibers, the carbon is the sp2 hybridized form. When carbon fibers are made, the curing process is designed to orient the graphitic layers within the carbon fibers along the fiber axis. This orientation is highly desirable to improve tensile strength. However, there is large variation between fibers, and they differ significantly

Microelectrodes and nanoelectrodes

A

57

B OH

hydroxyl

O

carbonyl O lactone

O O

carboxylic acid OH

O quinone O

Fig. 4.5.  Surface of carbon fibers (A) commonly encountered cross-sections of carbon fiber, where the lines represent the graphitic planes. (B) shows representative functional groups that may be present on the surface.

in the degree of orientation on the fiber axis and the size of the crystallite. Therefore, the cross section of the fiber can be significantly varied depending on the preparation procedure. These variables have significant importance to the electrochemical performance of the carbon fiber. Fig. 4.5A shows examples of commonly encountered carbon fiber cross-sections. Carbon fibers are quite disordered and exhibit a capacitance of around ∼20–50 µF/cm2 and fast electron-transfer rates. Unfortunately, there is still no clear relationship between the structure of the carbon fiber and its electrochemical activity due to the large number of fiber types and manufacturers. As a result, most users of carbon fibers have a fixed source which they characterize and use for bioanalysis applications [13]. Another major feature of the carbon fiber microelectrode is the surface oxide groups which terminate the carbon group. Fig. 4.5B shows the variety of, difficult to remove, surface oxides that could be present on the surface of carbon fibers. These surface oxides can have a variety of effects on electron-transfer, molecular interaction, and adsorption of analytes for electron transfer [14]. The presence of oxide functional groups raises the possibility of specific chemical interactions between the surface and a biomolecule in solution. This is mainly the case for inner sphere redox couples. Such interactions can catalyze redox reactions, promote or inhibit adsorption and affect the background current. Often, the quinone redox peak that can be observed on carbon electrodes (around 0.3 – 0.6 V) can greatly influence the voltammetry of redox species measured. This quinone peak can also be

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∆i (nA)

50

50

A

25

B

40 nA

25 1.2 0.6

0 –0.6 V

20 nA

0

1.2 0.6

0 –0.6 V

0

Fig. 4.6.  Calibration records for 5 mM dopamine (DA) calibration records in artificial CSF (ACSF) and in phosphate-buffered saline (PBS). Calibration in: (A) ACSF with a spark-etched, uncoated electrode; (B) PBS with a spark-etched, uncoated electrode. All calibrations were obtained at 32 °C in the recording chamber. FCV scan rate was 900 V s1, applied at 100-ms intervals; scan range was 0.5 to 1.1 V vs. Ag:AgCl. Each data point represents oxidation current averaged over 100 mV around the DA oxidation peak potential. Insets are DA subtraction voltammograms which are the difference between 10 averaged voltammograms at the calibration maximum and 10 averaged baseline records. Source: Kume-Kick, J.; Rice, M. E., Dependence of dopamine calibration factors on media Ca2 + and Mg2 + at carbon-fiber microelectrodes used with fast-scan cyclic voltammetry. Journal of Neuroscience Methods 1998, 84 (1), 55–62.

used for measuring the pH of the solution , but is very hard to remove. Ensuring that the subsequent positive excursions do not exceed ∼ + 1.0 V at neutral pH is assumed to keep only oxygen moieties such as carboxylates or aliphatic carbonyls, which are not electroactive and thus do not alter the capacitive current in between scans [15]. The influence of the background electrolyte on the behavior of faradaic responses on carbon fibers is also varied. Fig. 4.6 shows the response of dopamine prepared in different physiological medias on a carbon fiber microelectrode. Dopamine calibration factors and thus electrode sensitivities, were 2–3-fold higher in phosphate- or HEPES-buffered saline than in a bicarbonate-based artificial cerebrospinal fluid (ACSF). Removal of Ca2+ and Mg2+ from ACSF eliminated this difference, suggesting that divalent cations altered the response of dopamine on the electrode [16]. These findings highlight that when using carbon fibers for bioanalytical measurement, any form of calibration should be carried out in as close to physiological conditions as possible if robust estimates of true concentrations are to be made. As mentioned in Chapter 3, fast-scan cyclic voltammetry is frequently used for measurement of biomolecules, where a large background current is generated and subtracted to detect the analyte. Surface oxides can greatly increase background current by a factor of 10 or more [10,17,18], owing in part to interactions between the charged surface groups and electrolyte ions, which is important in the context of bioanalysis. Due to these factors, different modifications of the carbon fiber using chemical and electrochemical treatments have been used to enhance electrochemical activity for bioanalysis. The next sections will explore some examples of modifications that can alter the carbon fiber electrode surface in order to enhance sensitivity and selectivity for bioanalysis.

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4.2.4  Modification of carbon fiber microelectrodes There are various strategies that can be used to improve sensitivity and selectivity towards the detection of a specific analyte within a biological environment. Modifications include using electrochemical or chemical pretreatments to alter the surface chemistry of the carbon fiber microelectrode or alternatively covering the surface of the carbon fiber microelectrode using film coating.

4.2.4.1  Electrochemical pretreatment Often, prior to bioanalytical measurement, electrochemical pretreatment (the use of voltage to modify the chemistry at the surface of the electrode) or conditioning is used as an approach to enhance sensitivity and selectivity. Pretreatment can not only enhance the current response for a given analyte, but also shift the oxidation potentials. The effect of electrochemical pretreatment is highly dependent on various factors such as the type of carbon electrode surface, the composition and pH of the background solution, and most importantly the pre-treatment protocol executed [19,20]. Electrochemical pretreatments have been shown to cause a general enhancement of the response of most carbon-based electrodes, even though the actual mechanism of enhancement remains largely unknown [21–23]. Previous research has hypothesized that different mechanisms are dependent on the type of pre-treatment. Anodization, or holding at positive potentials, is believed to activate the functional groups present, which improves the electroactive surface area of an electrode. In addition, it results in the formation of a layer on the surface that is most likely an oxide layer as well as causes the oxidative removal of a layer of impurities from the surface, thus priming the surface for enhanced electrochemistry. Cathodization, or holding at negative potentials, on the other hand, is thought to be responsible for the stripping or the reduction of any oxide layers found on the surface, thus clearing the surface for optimal measurements. Finally, the process of cycling is thought to generate a surface similar to that post cathodization, due to the formation and subsequent stripping of an oxide layer in addition to the functionalization of the surface groups, resulting in an anionic surface [19]. Electrochemical pretreatment may also alter the microstructure of the carbon material, generating mesopores or breaks in the material exposing additional edges for enhanced activity. Therefore, the format of the electrochemical pretreatment needs to be specifically tailored to the application. Fig. 4.7 shows the untreated and pretreated (0 to + 2.6 V for 30 s) responses from two redox species (dopamine and ferrocyanide). The pretreated electrode has approximately a 10-fold increase in current for dopamine, whereas its response to ferrocyanide is, if anything, slightly decreased [21]. These findings indicate that electrochemical activity of different redox probes can be influenced by electrochemical pretreatment and how they interact with the surface.

4.2.4.2  Chemical pretreatment Another strategy for pretreatment of carbon fiber electrodes is chemical pretreatment. This approach can be used to block the activity of certain oxygen functional groups

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10

1.0

0.1

25 1.0

10 Conc [µM}

100

1000

Fig. 4.7.  Anodic peak current sensitivity of untreated and pretreated carbon fibers to ferrocyanide and dopamine oxidations by cyclic voltammetry: DA oxidation, (open circles) untreated fibers, (closed circles) pretreated fibers; ferrocyanide oxidation, (open squares) untreated fibers, (closed squares) pretreated fibers. Source: Feng, J. X.; Brazell, M.; Renner, K.; Kasser, R.; Adams, R. N., Electrochemical pretreatment of carbon fibers for in vivo electrochemistry: effects on sensitivity and response time. Analytical Chemistry 1987, 59 (14), 1863–1867.

which may reduce the electrochemical activity of the analyte of interest. Therefore, this could be used to provide selectivity for bioanalytical measurement. Alternatively, chemical pretreatment can involve attaching a chemical agent to the surface of the carbon fiber through covalent bonding, which can serve to enhance the sensitivity for the measurement of a specific analyte. Diazonium salt reduction has been demonstrated to be a versatile method to functionalize carbon−electrode surfaces [24]. Although strong bonding and dense packing of monolayers is produced by the diazonium reduction, some aryl diazonium salts tend to form multilayer structures on the electrode surfaces. This therefore does not provide a uniform chemical treatment over the electrode surface. Another approach, as shown in Fig. 4.8, is to modify carbon microelectrodes via reduction of 4-sulfobenzenediazonium tetrafluoroborate (4-SBD) [25]. This is a negatively charged molecule, which potentially acts to restrict the access of other negatively charged analytes such as ascorbic acid to the surface of the electrode, thus providing specificity for the oxidation of positively charged analytes. With this layer attached, adsorption of positively charged dopamine, noradrenaline, and serotonin to carbon fibers is increased

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B. 1 mM Ascorbic acid

A. 5 µM Norepinephrine 0.5 nA

7.5

1V

–0.4 V

1V

–0.4 V –0.5 nA

–2.5

C. 100 µM DOPAC

D. 5 µM Serotonin

1.00 nA

7.5 nA

1V

–0.4 V –2.5 nA

1V

–0.4 V –2.5 nA

Fig. 4.8.  Background-subtracted cyclic voltammograms for various compounds. Response at bare P-55 elliptical electrodes is shown with dashed lines, whereas the response at the 4-SBD modified electrode is shown with solid lines. The cyclic voltammograms were recorded in a flow injection cell 300 ms following injection of the analyte. Source: Hermans, A.; Seipel, A. T.; Miller, C. E.; Wightman, R. M., Carbon-Fiber Microelectrodes Modified with 4-Sulfobenzene Have Increased Sensitivity and Selectivity for Catecholamines. Langmuir 2006, 22 (5), 1964–1969.

leading to enhanced sensitivity. However, no clear presence of negatively charged ascorbic acid is observed, which is considered a major interferant. This example highlights how the clever use of chemical agent modifications can change the surface of the carbon fiber and provide clear sensitivity with enhanced selectivity. This approach is highly attractive when detection is needed in complex environments, where multiple analytes of similar structure are present.

4.2.4.3  Film coatings Film coating are essentially chemical membranes that are utilized to coat the surface of the carbon fiber electrode. Coatings are mainly used to enhance the selectivity of the electrode for detection of the analyte of interest, against the wide range of analytes present in bioanalytical environments. Additionally, coatings may also be of benefit to the sensitivity of the electrode, but this is less common than enhancing selectivity. The selectivity achieved by coatings can be through two main processes: (i) ion-selectivity by prevention of either negative or positively charged ions from reaching the surface of the electrode (charge exclusion) or (ii) through the prevention of molecules of certain sizes from reaching the electrode surface (size exclusion - this

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commonly used for detection of gaseous molecules, which will be discussed in Chapter 11, as well as to prevent biofouling by large biomolecules). Often a single coating of the electrode surface will be used, but in certain circumstances multiple film coatings would be applied to provide complete selectivity in complex biological environments. However, there are key limitations to film coatings. These are mainly associated with the film coating reducing the mass transfer of the analyte to the electrode surface and therefore reducing the response time of the electrode (temporal resolution). This is not necessarily the case for all types of film coatings, dependent on the type and thickness, but it is best to explore the response time of the electrode coated and uncoated to understand the capabilities of the sensor for bioanalytical measurement. There are two common methods used to apply film coatings on microelectrodes; dip coating and electropolymerization. In dip coating, the electrode is immersed within a solution of the film coating chemical to coat the surface of the electrode. This can then be held for a certain period to allow the coating to generally drip to create a smooth layer. Following this process, the film coating will need to cure, which can be done at room temperature over a period of a few days or at a higher temperature to speed up the process. This process can be repeated if more than one layer of film coating is desired. It is very fast and simple to conduct but can result in varying thickness of the film coating between electrodes which limits reproducibility. Electropolymerization is achieved by applying a potential and placing the tip of the electrode in a polymer solution which thus results in the generation of a polymeric film which grows on the surface of the electrode through electrochemical activity. This process only works if the polymer can be formed through electrochemical reaction and can be simple to conduct. Its main advantage over dip coating is that it, allows for thin-layer films that are uniform and utilizes the potential as well as the time of potential application to regulate the thickness of the film. One of the most widely used polymers on microelectrodes to date is, Nafion® which is a perfluorosulfonated negatively charged derivative of Teflon. It acts as a perm-selective film (this is a film that allow the preferential movement of certain ions through the ion-exchange membrane) on the surface of the electrode, preventing large biological macromolecules from blocking the electrode surface. In addition, Nafion® is highly sensitive to cationic amines such as dopamine and serotonin and has minimal sensitivity to anions such as ascorbic acid and uric acid at physiological concentrations [26]. Other examples of film coatings which have been utilized includes base hydrolyzed cellulose acetate (HCA), and fibronectin. HCA was found to be relatively fouling resistant to dopamine whilst fibronectin coating is associated with moderate losses in sensitivity after coating and fouling [27]. Other membranes are utilized for the measurement of gaseous transmitters, where the role of the membrane is to prevent large molecules to reach the electrode surface. Membranes have been electropolymerized onto the carbon fiber electrodes, where poly(o-phenylenediamine) [28–30] and poly(eugenol) [31] have successful been used to monitor important gaseous transmitters like nitric oxide [32].

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4.3  Microelectrode arrays Within bioanalysis, the quest to obtain greater in-depth knowledge has led to development of sensing devices where multiple measurements can be conducted. Microelectrode arrays are devices where more than one detection electrode, which has geometry in the micrometer scale, is used for measurement. Microelectrode arrays can be fabricated using a variety of different methods [33]. The simplest approach is the assembly of electrode materials within insulators such as epoxy and elastomer. This approach has many advantages such as the ability to use a range of different materials (metal: gold, platinum, silver, nickel, etc.; carbon: carbon fiber, carbon paste, graphite, glassy carbon, etc.), relative ease of fabrication, robustness to polishing, and no special laboratory equipment needed other than a clean room. However, it is difficult to create a perfect regular microelectrode array due to the physical nature of the assembly process. Therefore, batch to batch variability exists. A popular and widely used approach for generation of microelectrode arrays is photolithography [33]. Photolithography refers to a process used in microfabrication to selectively remove parts of a thin film (or the bulk of a substrate). It is based on photoresists (light‐sensitive chemicals) and exposure tools equipped with mercury arc lamp illumination sources producing ultraviolet wavelengths. The detailed process is highlighted in Fig. 4.9. The developed microelectrode array is highly precise in geometry but are usually made using metal films, and thus their narrow potential window in the negative region may limit their application to reductive electrochemical reactions. Other approaches to fabricating arrays include screen-printing and electrodeposition, however, these approaches are not very common for the development of microelectrode arrays. For screen-printed electrodes, they lack the resolution for the generation of electrode with small diameters and are often used for millimeter sized electrodes. For electrodeposition, this approach can be difficult to accurately regulate to generate monolayer films. Microelectrode arrays provide the means to measure several different analytes within the biological environment or to monitor the same analyte over different regions of the biological cell/tissue where measurements are being conducted. These benefits do not come without limitations. Firstly, given that radial diffusion occurs at microelectrodes, there is potential for “crosstalk” between the two electrodes. If the electrodes are close together then radial pathways will overlap and result in reduced current response. However, by increasing the distance between adjacent electrodes to prevent any “crosstalk”, it may result in the measurements conducted not being in the same location within the cell/tissue. The size and geometry of microelectrode arrays pose challenges for conducting non-invasive measurement, particularly in the case of in vivo measurements. Lastly, if the microelectrode array is being used to monitor the release of different analytes within a cell/tissue, the differences in the locality of the sensors, even over micrometer distances, may result in measuring different biological activity which may not be comparative. Arrays can provide a depth and wealth of bioanalytical measurements, but the usefulness can be hindered by geometric constraints.

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A 1 Silicon/glass/quartz plate

2 Silicon oxide or silicon nitride

3 Positive or negative photoresist

Adhesive layer Electrode materials layer

4 Desired mask 1

Wet-etching

5 Formation of insulator layer

B

6 Mask 2

Dry-etching

C 1 2 Positive PR

Negative PR

3 Silicon/glass/quartz plate

Photoresist

SiO2/Si3N4/SiO2

Silicon oxide or silicon nitride

Desired mask

UV light

Adhesive layer

Electrode materials layer

Fig. 4.9.  (A) Fabrication procedure using standard photolithographic techniques to construct microelectrode arrays: (1) thermal oxidation of silicon in order to obtain a SiO2 layer for electrical isolation between the substrate and the surface (this step is not necessary if substrates of glass plates are used); (2) an adhesive layer followed by electrode materials was then deposited onto the oxide; (3) after spin‐coating the photoresist (PR), the desired mask 1 containing the pattern for the electrodes and their interconnections covered the surface for the next etching; (4) the wafer was then dipped in successive etching solutions to dissolve metallic electrode materials and adhesive layer; (5) after removal of the PR, an insulating layer was evaporated to the whole surface of the substrate, and the insulating layer was subsequently patterned using a desired mask 2; (6) etching process to create openings of the microelectrode recording sites and the connector pads. (B) Alternative way to create a microelectrode array, corresponding to the flow marked by dotted line frame. (C) Difference between positive and negative PR under UV light. Source: Huang, X.-J.; O’Mahony, A. M.; Compton, R. G., Microelectrode Arrays for Electrochemistry: Approaches to Fabrication. Small 2009, 5 (7), 776–788.

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5µm

65

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carbon fiber 20 µm

1

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2

6

3

7

20 µm i / pA

600 400

4

5

4 6

7

3 1

2

200 20 µm

0 –0.3 0 0.3 0.6 V vs Ag//AgCI

Fig. 4.10.  Carbon fiber microelectrode arrays. (A) schematic drawing and (B) scanning electron microscopy of carbon fiber microelectrodes containing two, three and seven disc electrodes. (C) Steady-state voltammetric response at 20 mV/s of a seven-fiber MEA (1–7) in 1 mM FcCH2OH and 0.2 M KCl, where. 1—7 a schematic of the microelectrode assemble is shown in the bottom right. Source: Zhang, B.; Adams, K. L.; Luber, S. J.; Eves, D. J.; Heien, M. L.; Ewing, A. G., Spatially and Temporally Resolved Single-Cell Exocytosis Utilizing Individually Addressable Carbon Microelectrode Arrays. Anal. Chem. 2008, 80 (5), 1394–1400.

Fig. 4.10 shows a novel approach for the development of carbon microelectrode array. Fabrication involves pulling a multibarrel glass capillary, containing a single carbon fiber in each barrel, into a sharp tip. The electrode tip is then beveled to form an array (10−20 μm) of carbon microdiscs. This approach provides the means to make microelectrode arrays ranging from 2 to 7 sensing elements. This simple fabrication procedure eliminates the need for complicated wiring of the independent electrodes, thus allowing preparation of high-density individually addressable microelectrodes. This approach has been enhanced to utilize even greater number of electrodes where carbon ring microelectrodes were deposited by pyrolysis of acetylene in the lumen of quartz capillary arrays [34]. However, with all electrodes in proximity of each other, the risk of crosstalk increases. Crosstalk is where the diffusional pathways of electrodes adjacent to one another cross over. For microelectrodes this would be the overlapping of radial diffusional pathways. This in the concept of the measurement means that two electrodes are competing for the oxidation of a single analyte and therefore the current will be lower than expected on an electrode due to this competition. This reduction in the limiting current on certain electrodes can be observed in the multi-barrel electrode due to geometrically hindered diffusion combined with depletion of the analyte by the surrounding electrodes. This is more evident for the center electrode, as the limiting current at this electrode is somewhat shielded by the collection of microelectrodes surrounding it [35].

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2

3

1 3, 4 - ACh Recording

Pt

Pt

ChOX

ACh 3

4

ACh

ACh

ACh

Ch

DA

ACh AChE

1, 2 - Ch Sentinel

AA

B 0.2 nA

4

ChOX

A

Ch

C DA

mPD

AA

1

2

DA

mPD

AA

0

200

400

600

800

Seconds

Fig. 4.11.  Ceramic microelectrode array. (A) Photograph of the microelectrode tip showing a schematic diagram highlighting the detection strategy for acetylcholine and control (choline sentinel electrode). The microelectrode array consists of four 333 × 15 µm Pt recording sites. Note the blunt tip, which penetrates brain tissue with minimal damage. (B) Time plots showing microelectrode current responses to additions of stock solutions during calibration. Upper trace is the signal from the ACh ± Ch electrode and the lower trace is the Ch sentinel electrode (control electrode). Additions of analytes and interferents occurred at the arrows: AA (ascorbic acid), Ach (acetylcholine), Ch (choline) and DA (dopamine). (C) The Ch signal can be removed by self-referencing or subtraction. Source: Burmeister, J. J.; Pomerleau, F.; Huettl, P.; Gash, C. R.; Werner, C. E.; Bruno, J. P.; Gerhardt, G. A., Ceramic-based multisite microelectrode arrays for simultaneous measures of choline and acetylcholine in CNS. Biosensors and Bioelectronics 2008, 23 (9), 1382-1389.

Another example of an array used for bioanalytical measurements is a ceramicbased 4-site microelectrode array consisting of 2 pairs of 333  μm × 15  μm Pt recording sites with 30 μm spacing between adjacent pairs and a 100 μm distance between the paired ends of the microelectrode array. This device has been used for in vivo measurement of acetylcholine [36]. Fig. 4.11 shows the microelectrode device and strategy for determination of acetylcholine using biosensing, which will be explained in greater detail in Chapter 12. Within this electrode device, two electrodes (number 1 and 2) are used as the control (named sentinel) and will monitor any substances present within the biological environment that can be oxidized at the voltage applied. The other two electrodes (number 3 and 4) are modified to contain a biological entity, that provides the ability to monitor acetylcholine as well as other oxidizable substances. Therefore, accurate detection of acetylcholine can be achieved as the background response can be subtracted. This approach is popular when using biosensors, however the major limitation is that the control and detection electrodes are in different locations, therefore potentially making subtraction unreliable as the microenvironment varies around different biological environments and is often dynamic and evolving. Accurate detection of acetylcholine within this type of microelectrode array assumes that the background analytes detected on all electrodes are equal.

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Microelectrode arrays have great benefits, as they can provide spatial information on cells/tissues as well as the potential to measure multiple analytes simultaneously. However, when measuring signals, you will need to account for any crosstalk and also consider the fact that measurement of different geometric locations in biological cells/ tissues can have varied responses, potentially resulting in control measurements being different to the true control values.

4.4 Nanoelectrodes Nanoelectrodes, with dimensions below 100  nm, are becoming highly popular in bioanalysis due to the significant advantages of their high sensitivity and high spatial resolution. These electrodes have attracted increasing attention in various fields such as single cell analysis (namely for intracellular measurements) and for conducting high-resolution imaging on cell membrane surfaces [37]. One of the most widely used approaches for fabrication of nanoelectrodes, with good control of the electrode geometry and high reproducibility, is to pull nanopipettes together with an incorporated metal wire using a laser-assisted pipette puller. The success of this approach is based on having a high-quality pipette puller however, excess glass from the tip needs to be removed by either chemical etching in HF or electrode polishing [37,38]. An alternative approach for fabrication of nanoelectrodes, which produces a carbon electrode, is the pyrolytic decomposition of carbon precursor gases inside pulled quartz glass nanopipettes, which gives rise to nanometric carbon electrodes. The major advantage of this fabrication method is the ease of fabrication and handling of these electrodes. Electrodes can be generated with ease in rapidly and require no polishing or tedious connecting of the electrodes. Except for a reliable high-quality capillary puller, only inexpensive materials and equipment are required for this manufacturing method. Working with nanoelectrodes brings different challenges which can impact bioanalysis. Even after using elaborate fabrication approaches and thorough polishing procedures, nanoelectrodes are very likely to change size, shape and state of the electrode surface during handling, storage, and electrochemical experiments. Such alterations can drastically vary the accuracy of measurements conducted if electrodes are used for multiple bioanalytical measurements. At the surface of nanoelectrodes, fouling and adsorption of contaminants from the air will rapidly deteriorate the properties of the electrode surface when compared to carbon fiber microelectrodes. This is a major issue for bioanalytical measurements, given the nature of complex biomolecules present within the environment that are prone to stick to the electrode surface and reduce the ability for long-term robust measurements. A more simplistic approach to the development of a nanoelectrode, is the fabrication of a nanotip microelectrode. This is fabricated from a 5 µm cylindrical carbon fiber microelectrode, which is flame etched (in this process the carbon layers within the electrode are eroded away using a high heat flame source). The resultant flameetched carbon-fiber microelectrode has a sharp tip and somewhat larger effective surface to enable insertion into the cell as shown in Fig. 4.12. These electrodes

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B

C

2.0 C)

I / nA

1.5 1.0 0.5 0.0 –0.2

0.0 0.2 0.4 0.6 E / V (vs. Ag/AgCI)

0.8

F

DOQ t

2e−

G

2e−

20 pA

E

I

DA

D

20 s

Fig. 4.12.  Nanotip electrode for intracellular measurements. (A) Global view (scanning electron microscopy) of a nanotip conical carbon-fiber microelectrode (scale bar: 5 mm). (B) Amplified view of the tip of another nanotip conical carbon-fiber microelectrode (scale bar: 400 nm). (C) Representative cyclic voltammogram (CV) of 0.10 mm dopamine at a nanotip conical carbon-fiber microelectrode (scan rate: 0.10 V/s). (D) Schematic illustration (small orange circles represent vesicles) and (E) bright-field photomicrograph (electrode approaching a cell from the lower right) of a nanotip conical carbon-fiber microelectrode placed in the cytoplasm of a single PC12 cell (scale bar: 20 mm). (F) Mechanism of the adsorption and opening of vesicles on the in situ electrode. DA = dopamine, DOQ = dopamine orthoquinone. (G) Amperometric traces for a nanotip conical carbon-fiber microelectrode placed inside a PC12 cell. Source: Li, X.; Majdi, S.; Dunevall, J.; Fathali, H.; Ewing, AG, Quantitative Measurement of Transmitters in Individual Vesicles in the Cytoplasm of Single Cells with Nanotip Electrodes. Angewandte Chemie International Edition 2015, 54 (41), 11,978–11,982.

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also have increased sensitivity, kinetics, and signal-to-noise ratio and a faster time response for many neurotransmitters, as compared to cylindrical carbon-fiber microelectrodes and electrochemically etched carbon-fiber microelectrodes. The nanoscale tips make these electrodes nearly ideal single-cell surgical tools for measurements with high spatial and temporal resolution and minimal disturbance to the cells. These advantageous properties make flame-etched carbon-fiber microelectrodes particularly attractive for effective detection of the contents of individual vesicles in the cellcytoplasm environment [39]. The ability to create a nanoelectrode that has the potential to be inserted within a single cell provides the means for intracellular measurements to sample internalized content, which cannot be achieved with microelectrodes that are to invasive for single cell measurements. The majority of studies using electrochemistry in bioanalysis are focused on monitoring chemicals that are released, namely due to the limitation in the tools to monitor intracellularly. Fig. 4.12 shows the ability to monitor vesicles within a single cell, which provides for the first time the ability to accurately measure the content present within the vesicle. This can be a powerful tool to determine what proportion of the intracellular content can be released. The novel development of a nanotip electrode highlights how advances in small-scale sensing devices can offer new insight into chemical levels within biological environments that have previously been anatomically difficult to explore. Nanoelectrodes can be combined with other technologies, such as microscopy, to create multifunctional devices. Fig. 4.13 shows the topography of differentiated PC12 cells as investigated using double‐barrel carbon nanoprobes (DBCNPs) as a nanoelectrode capable of conducting both scanning electrochemical microscopy (SECM) with scanning ion conductance microscopy (SICM). This work shows that the nanoprobe has the ability to provide real enhanced resolution for imaging, as the ability to

18.0 µ

m

m

95.3 µ

95.3 µm

15.0 µm

18.0 µm

3.5 µm

Fig. 4.13.  Nanoscale topography images of differentiated PC12 cells using the DBCNP nanopipette. The arrows showed the dendritic structures. Source: Takahashi, Y.; Shevchuk, A. I.; Novak, P.; Zhang, Y.; Ebejer, N.; Macpherson, J. V.; Unwin, P. R.; Pollard, A. J.; Roy, D.; Clifford, C. A.; Shiku, H.; Matsue, T.; Klenerman, D.; Korchev, Y. E., Multifunctional Nanoprobes for Nanoscale Chemical Imaging and Localized Chemical Delivery at Surfaces and Interfaces. Angewandte Chemie International Edition 2011, 50 (41), 9638–9642.

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observe neuronal features such as dendritic structures of the neuron [40]. The ability to conduct imaging as well as electrochemical measurement provides a powerful tool for bioanalytical measurement and can unravel new information about the locality of signaling molecules and their pattern of release from cells/tissues. Overall, the popularity of nanoelectrodes and their application in bioanalytical measurements is growing as their ability to provide enhanced spatial resolution within cells is seen as an opportunity to gain novel understanding of rapid, dynamic biological functions such as transmission and metabolism.

4.5 Summary Microelectrodes have been the dominant force in bioanalysis namely due to the appropriate geometric size to sample biological environments. Microelectrodes provide the appropriate spatial resolution and are able to be inserted into varying biological environments. Even though varying conductive materials can be used to fabricate microelectrodes, carbon fiber microelectrodes are most commonly used for bioanalytical measurement, as the versatile nature of this electrode allows for tailoring of the electrode surface. Microelectrode arrays provide the potential to explore and sample varying regions of tissues/cells to understand the changes in either one or multiple different analytes. Nanoelectrodes have provided a new dawn in bioanalytical measurement with the high-resolution imaging capabilities of the sensor providing the ability to understand changes in chemicals over small biological structures and development of sensors that can penetrate the cell for intracellular measurements.

References [1] M.I. Davies, C.E. Lunte, Microdialysis sampling coupled on-line to microseparation technique, Chem. Soc. Rev. 26 (1997) 215–222. [2] C.E. Lunte, D.O. Scott, P.T. Kissinger, Sampling living systems using microdialysis probes, Anal. Chem. 63 (15) (1991) 773A–780A. [3] A.S. Khan, A.C. Michael, Invasive consequences of using micro-electrodes and microdialysis probes in the brain, TrAC Trends in Anal. Chem. 22 (9) (2003) 503–508. [4] L.M. Borland, et al., Voltammetric study of extracellular dopamine near microdialysis probes acutely implanted in the striatum of the anesthetized rat, J. Neurosci. Methods 146 (2) (2005) 149–158. [5] A. Jaquins-Gerstl, A.C. Michael, Comparison of the brain penetration injury associated with microdialysis and voltammetry, J. Neurosci. Methods 183 (2) (2009) 127–135. [6] C.G. Zoski, Ultramicroelectrodes: design, fabrication, and characterization, Electroanalysis 14 (15-16) (2002) 1041–1051. [7] J.L. Ponchon, et al., Normal pulse polarography with carbon fiber electrodes for in vitro and in vivo determination of catecholamines, Anal. Chem. 51 (9) (1979) 1483–1486. [8] J. Millar, G.V. Williams, Ultra-low noise silver-plated carbon fibre microelectrodes, J. Neurosci. Methods 25 (1) (1988) 59–62. [9] R.S. Kelly, R.M. Wightman, Bevelled carbon-fiber ultramicroelectrodes, Anal. Chim. Acta 187 (1986) 79–87.

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[10] M.L. Huffman, B.J. Venton, Carbon-fiber microelectrodes for in vivo applications, Analyst 134 (1) (2009) 18–24. [11] J. Millar, T.G. Barnett, A low-noise optically isolcated preamplifier for use with extracellular microelectrodes, J. Neurosci. Methods 51 (1994) 119–122. [12] J. Millar, C.W.A. Pelling, Improved methods for construction of carbon fibre electrodes for extracellular spike recording, J. Neurosci. Methods 110 (1–2) (2001) 1–8. [13] R.L. McCreery, K.K. Cline, Carbon Electrodes, in: P.T. Kissinger, W.R. Heineiman (Eds.), in Laboratory Techniques in Electroanalytical Chemistry, Marcel Dekker, New York, 1996, pp. 293–332. [14] P. Chen, R.L. McCreery, Control of electron transfer kinetics at glassy carbon electrodes by specific surface modification, Anal. Chem. 68 (22) (1996) 3958–3965. [15] R.L. McCreery, Carbon electrodes: structural effects on electron transfer kinetics, Electroanal. Chem. 17 (1991) 221–374. [16] J. Kume-Kick, M.E. Rice, Dependence of dopamine calibration factors on media Ca2+ and Mg2+ at carbon-fiber microelectrodes used with fast-scan cyclic voltammetry, J. Neurosci. Methods 84 (1) (1998) 55–62. [17] P.L. Runnels, et al., Effect of pH and surface functionalities on the cyclic voltammetric responses of carbon-fiber microelectrodes, Anal. Chem. 71 (14) (1999) 2782–2789. [18] J.G. Roberts, et al., Specific oxygen-containing functional groups on the carbon surface underlie an enhanced sensitivity to dopamine at electrochemically pretreated carbon fiber microelectrodes, Langmuir 26 (11) (2010) 9116–9122. [19] R.C. Engstrom, Electrochemical pretreatment of glassy-carbon electrodes, Anal. Chem. 54 (13) (1982) 2310–2314. [20] R.L. McCreery, Advanced carbon electrode materials for molecular electrochemistry, Chem. Rev. 108 (7) (2008) 2646–2687. [21] J.X. Feng, et al., Electrochemical pretreatment of carbon fibers for in vivo electrochemistry: effects on sensitivity and response time, Anal. Chem. 59 (14) (1987) 1863–1867. [22] K. Ravichandran, R.P. Baldwin, Enhanced voltammetric response by electrochemical pretreatment of carbon paste electrodes, Anal. Chem. 56 (9) (1984) 1744–1747. [23] R.C. Engstrom, V.A. Strasser, Characterization of electrochemically pretreated glassy carbon electrodes, Anal. Chem. 56 (2) (1984) 136–141. [24] M. Delamar, et al., Covalent modification of carbon surfaces by grafting of functionalized aryl radicals produced from electrochemical reduction of diazonium salts, J. Am. Chem. Soc. 114 (14) (1992) 5883–5884. [25] A. Hermans, et al., Carbon-fiber microelectrodes modified with 4-sulfobenzene have increased sensitivity and selectivity for catecholamines, Langmuir 22 (5) (2006) 1964– 1969. [26] M.P. Brazell, et al., Electrocoating carbon fiber microelectrodes with Nafion improves selectivity for electroactive neurotransmitters, J. Neurosci. Methods 22 (2) (1987) 167–172. [27] Y.S. Singh, et al., Head-to-head comparisons of carbon fiber microelectrode coatings for sensitive and selective neurotransmitter detection by voltammetry, Anal. Chem. 17 (2011) (2011) 6658–6666, p.null-null. [28] M.N. Friedemann, S.W. Robinson, G.A. Gerhardt, o-Phenylenediamine-modified carbon fiber electrodes for the detection of nitric oxide, Anal. Chem. 68 (15) (1996) 2621–2628. [29] J.-.K. Park, et al., In vivo nitric oxide sensor using non-conducting polymer-modified carbon fiber, Biosens. Bioelectron. 13 (11) (1998) 1187–1195. [30] N.R. Ferreira, et  al., Electrochemical measurement of endogenously produced nitric oxide in brain slices using Nafion/o-phenylenediamine modified carbon fiber microelectrodes, Anal. Chim. Acta 535 (1–2) (2005) 1–7.

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[31] B.A. Patel, et al., Detection of Nitric Oxide release from single neurons in the pond snail, Lymnaea stagnalis, Anal. Chem. 78 (22) (2006) 7643–7648. [32] F. Bedioui, N. Villeneuve, Electrochemical nitric oxide sensors for biological samples principle, selected examples and applications, Electroanalysis 15 (1) (2003) 5–18. [33] X.-J. Huang, A.M. O’Mahony, R.G. Compton, Microelectrode arrays for electrochemistry: approaches to fabrication, Small 5 (7) (2009) 776–788. [34] Y. Lin, et al., Carbon-ring microelectrode arrays for electrochemical imaging of single cell exocytosis: fabrication and characterization, Anal. Chem. 84 (6) (2012) 2949–2954. [35] B. Zhang, et al., Spatially and temporally resolved single-cell exocytosis utilizing individually addressable carbon microelectrode arrays, Anal. Chem. 80 (5) (2008) 1394–1400. [36] J.J. Burmeister, et al., Ceramic-based multisite microelectrode arrays for simultaneous measures of choline and acetylcholine in CNS, Biosens. Bioelectron. 23 (9) (2008) 1382–1389. [37] J.T. Cox, B. Zhang, Nanoelectrodes: recent advances and new directions, Annu. Rev. Anal. Chem. 5 (1) (2012) 253–272. [38] D.W. Arrigan, Nanoelectrodes, nanoelectrode arrays and their applications, Analyst 129 (12) (2004) 1157–1165. [39] X. Li, et al., Quantitative measurement of transmitters in individual vesicles in the cytoplasm of single cells with nanotip electrodes, Angew. Chem. Int. Ed. 54 (41) (2015) 11978–11982. [40] Y. Takahashi, et al., Multifunctional nanoprobes for nanoscale chemical imaging and localized chemical delivery at surfaces and interfaces, Angew. Chem. Int. Ed. 50 (41) (2011) 9638–9642.

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Bhavik A. Patel School of Pharmacy and Biomolecular Sciences, Centre for Stress and Age-Related Disease, University of Brighton, Brighton, UK

Keypoints Principles Enhancing the performance of electrochemical measurement in bioanalysis can be achieved through novel and interesting sensor development. Many different types of novel carbon-based materials (carbon nanotubes and diamond) have been the basis of sensor fabrication, in order to enhance sensitivity and stability. The trends have been to create novel approaches to manufacture such electrodes to maintain the microelectrode geometry. However, for making sensors to monitor over wider areas, or in complicated shapes and sizes, 3D printing has emerged as an exciting platform to make electrodes for bioanalysis in ways that were not imaginable.

Applications in bioanalysis ■

■

Novel carbon-based materials such as carbon nanotubes and boron-doped diamond have allowed for more diverse measurement in vivo, ex vivo or in vitro measurement of a host of biomolecules. 3D printed electrodes can be made in any size and shape and therefore provides the ability to interface sensing into new biological environments.

Strengths ■

■

The use of new carbon-based materials has enhanced the sensitivity and stability for bioanalytical measurements, when compared to the gold standard carbon fibre microelectrode. Electrodes of any size and shape can be fabricated using 3D printing, providing new and novel approaches to interface sensors in biological environments.

Limitations ■

■

Often the ability to make sensors using novel materials require complex processes as they usually need to be attached or grown on conductive solid materials. 3D printable materials are not highly conductive and thus have limited capabilities for electrochemical measurements.

Electrochemistry for Bioanalysis. DOI: 10.1016/C2019-0-03108-1 Copyright © 2021 Elsevier Inc. All rights reserved.

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5.1 Introduction Carbon based materials have been widely used because of their low cost, good electron transfer kinetics, good chemical stability, and biocompatibility. However, to-date carbon fiber microelectrodes have been the most dominant form of electrode utilized in bioanalysis. To reach new boundaries in bioanalytical measurements, diversity in the materials utilized provide an attractive approach for the detection of low levels of biomolecules in new environments. Although most measurements are focused on small geometries, measurement from biological environments where tissue sizes are large, are posing new challenges. Therefore, new approaches in electrode manufacturing are needed to provide robust batch-to-batch manufacturing of electrodes with larger geometries. With the ability to use new materials and manufacturing processes that can challenge the size and scale of sensing measurements, a new dawn of bioanalytical measurement can be achieved.

5.2  Novel carbon materials for generation of electrodes The performance of the carbon fiber electrode is limited, and with the need to detect more varied biomolecules at significantly low concentrations in complex biological environments, the evolution of the electrode materials utilized is widely considered a means to exploring new boundaries in bioanalysis. Often the materials that are utilized are allotropes of carbon, of which the structures can be seen in Fig. 5.1. Carbon fiber, which to-date is the most used carbon material, is a graphitic structure of multiple graphene sheets. Other utilized forms include graphene itself, which is a single layer of graphite. If this single layer of graphite is then rolled in a helical fashion, a carbon nanotube can be formed, which can be again single or multiple layers. All these carbon-based structures are sp2 hybridized. The one exception is diamond, which has varying properties and is sp3 hybridized. Each of these varying allotropes have provided new enhancements into bioanalysis.

5.2.1  Carbon nanotubes Carbon nanotubes were discovered in 1991 by Iijima, who produced these nanotubes in a similar fashion to buckminsterfullerene and other fullerenes [1]. Carbon nanomaterials are widely been incorporated into sensors. The sizes of nanomaterials are 1 to 100 nm graphite

graphene

nanotube

diamond

Fig. 5.1.  Different allotropes of carbon that have been utilized to fabricate electrodes that are used for bioanalysis.

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Table 5.1  Comparison between SWCNTs and MWCNTs. SWCNT

MWCNT

Single layer of graphene Catalyst is required for synthesis Bulk synthesis is difficult as it requires sophisticated control over growth and atmospheric conditions Purity is poor A chance of defect is high during functionalization Less accumulation in the body Characterization and evaluation are easy Can be easily twisted and is more pliable

Multiple layers of graphene Can be produced without catalyst Bulk synthesis is easy

Purity is high A chance of defect is less but once it occurrs it becomes difficult to improve More accumulation in the body Has a very complex structure Cannot be easily twisted

and they are advantageous, because of their large surface-to-volume ratio and specific surface area. In addition, carbon nanomaterials have enhanced interfacial adsorption properties, better electrocatalytic activity, higher biocompatibility, and faster electron transfer kinetics compared to many traditional electrochemical sensor materials. Depending on the number of sheets rolled into concentric cylinders, there are two main categories of carbon nanotubes: single-walled carbon nanotubes (SWCNTs) and multiwalled carbon nanotubes (MWCNTs). Table 5.1 shows a comparison between the key features of both types of nanotubes. SWCNTs are made up of a single graphene sheet wrapped into a hexagonal close-packed cylindrical structure that is held together by van der Waals forces. This makes them easily twistable and more pliable. MWCNTs consist of several cylinders, each made of a single graphene sheet surrounding a hollow core. Given the sp2 hybridization structure within cylinders of nanotubes in MWCNTs, a delocalized electron cloud provides interactions between the adjacent cylindrical layers thus resulting in a less flexible and more defective structure. Both types of nanotubes, as well as more exotic types, have been utilized for the development of sensors for bioanalytical detection, however MWCNTs are more widely used at present. This is mainly due to the fact that they are widely available at mass and can be made to high purity without significant impurities from other forms of carbon or metals.

5.2.1.1  Preparation of carbon nanotubes The production of carbon nanotubes is mainly associated with the transformation of a carbon source into nanotubes, usually at high temperature and low pressure, wherein the synthesis conditions influence the characteristics of the final product. However, in all processes that are utilized for the preparation of carbon nanotubes, there is often the generation of carbon-based or metallic impurities. Therefore, post-synthesis, purification is an essential step to be considered to ensure that the electrochemical activity of the nanotubes is explored and not that of conductive metal and carbon impurities. Fig. 5.2 shows schematic representations of the core methods used for the synthesis of carbon nanotubes and the mechanism on how these nanotubes are synthesized.

Source: Rastogi, V.; Yadav, P.; Bhattacharya, S. S.; Mishra, A. K.; Verma, N.; Verma, A.; Pandit, J. K., Carbon nanotubes: an emerging drug carrier for targeting cancer cells. Journal of drug delivery 2014, 2014.

Fig. 5.2.  Schematic representation and mechanism of methods used for carbon nanotube synthesis: (A) Arc discharge method, (B) chemical vapour deposition method, (C) laser ablation method.

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Table 5.2  Summary and comparison of the three most common carbon nanotube synthesis methods.

Arc discharge

CVD

Laser ablation

Yield rate SWCNT or MWCNT Advantages

>75 percent Both

>75 percent Both

1. Simple 2. Inexpensive 3. High quality

1. Relatively high purity 2. Room-temperature synthesis

Disadvantages

1. High temperature 2. Purification required 3. Tangled nanotubes

1. Method limited in production 2. Crude product purification required

>75 percent Both (mainly SWCNT) 1. Simple 2. Low temperature high purity 3. Large-scale production 4. Aligned growth possible 1. Synthesized CNTs are usually SWCNTs 2. Defects present

Table 5.2 shows a summary of the key advantages and disadvantages of the various approaches utilized for carbon nanotube preparation.

5.2.1.1.1  Arc discharge method This was the first method utilized for the fabrication of carbon nanotubes, by Iijima. In this method, an arc is generated when a DC current of 200 A to 20 V is applied across two carbon electrodes which are placed in a vacuum chamber that is typically filled with inert gas (e.g. helium, argon) at low pressure (~50 –700  mbar). The positive electrode is gradually brought closer to the negative one, to generate the electric arc. The electrodes become red hot and such an extreme condition turns the gas into plasma. Once the arc stabilizes, the electrodes are kept about a millimeter apart while the carbon nanotubes deposit on the negative electrode. Arc discharge is the most practical and inexpensive approach, because the method yields highly graphitized tubes due to the high process temperature. However, many by-products besides carbon nanotubes are generated. Therefore, important purification steps to maintain the quality of nanotubes and eliminate species such as amorphous carbon and metallic nanoparticles are needed. Additionally, this process for producing nanotubes provides no control on alignment. Alignment is important to provide enhanced conductivity.

5.2.1.1.2  Chemical vapour deposition (CVD) While the arc discharge method can produce large quantities of impurified nanotubes, CVD provides a more controllable process for the selective production of nanotubes with predefined properties. In this method a mixture of hydrocarbon gas (ethylene, methane, or acetylene) and a process gas (ammonia, nitrogen, and hydrogen) are made

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to react in a reaction chamber on a heated metal substrate at temperatures of around 700 °C–900 °C, at atmospheric pressures. Residual gas diffuses away, whereas free carbon atoms dissolve into the nanoparticles and then segregate to the catalyst surface to form nanotubes. The key parameters include the nature of hydrocarbons, catalysts, and the growth temperature. The process using CVD is hard to scale up for production of large quantities but can produce aligned and ordered carbon nanotubes. The microstructure of the carbon nanotube tips synthesized by the CVD technique, exhibits well-formed caps compared to other techniques. A major drawback associated with the CVD technique is that there are high defect densities in the MWCNT structures grown by this process, due to the lack of sufficient thermal energy for annealing CNTs because of the relatively low growth temperature.

5.2.1.1.3  Laser ablation method In this approach a direct laser beam is focused on transition-metal/graphite composite rods to produce carbon nanotubes. A pulsed laser is made to strike at a graphite target in a high temperature reactor (furnace) in the presence of inert gas, such as helium which vaporizes the graphite target and forms a laser plume. The laser plume contains vaporized carbon and metallic nanoparticles that lead to the reassembling of carbon, in the form of carbon nanotubes. The nanotubes develop on the cooler surfaces of the reactor, as the vaporized carbon condenses. Nanotubes produced by laser ablation have higher purity (up to about 90 percent) and their structure is better graphitized, than those produced by the arc discharge method. The high cost of laser and small carbon amounts produced are the major limitations of this method. In addition, the method mainly favors the growth of SWCNTs, and special reaction conditions are required to generate MWCNTs.

5.2.1.2  Making carbon nanotube sensors There are varying strategies that can be utilized for incorporating carbon nanotubes into electrode fabrication. The simplest approach is dip coating or drop casting nanotubes onto a solid carbon electrode. Fig. 5.3 shows an example of the dip coating approach, where there have been two different strategies that have been utilized. The first involves the dispersion of nanotubes into an organic solvent and dip coating the electrode to allow nanotubes to adhere to the electrode surface. This is most effective method, if the electrode is carbon based as this allows for adherence of the nanotubes onto the electrode by π-π forces. This approach increases the surface area of the electrode with the additional presence of the nanotubes on the surface. However, in this strategy, the number of nanotubes that cover the electrode surface varies from electrode to electrode, creating poor reproducibility. In addition, it is not clear if the improved electrochemical performance is a result of the nanotubes or just the effect of the enhanced surface area. Lastly as the nanotubes are loosely held, they may not stay adhered on the electrode surface during the measurement meaning that the electrode surface may alter over time. This is a significant limitation of this strategy of fabricating nanotube electrodes.

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Nanotubes dispersed in organic solvent

Nanotubes dispersed in polymer

Fig. 5.3.  Strategies for dip coating nanotubes onto the electrode surface.

To enhance the stability of the electrode for robust measurements, dip coating approaches focused on the dispersion of nanotubes in polymeric coatings were developed (Fig. 5.3). This approach is simple, but still suffers from poor batch to batch reproducibility due to the number of nanotubes present on the electrode and the consistency in the coating thickness of the polymer film. The introduction of polymer has drawbacks including restricting diffusion, slowing temporal resolution, and decreasing conductivity. However, some polymeric films can be electropolymerized onto the electrode surface and therefore can provide a more uniform film which have nanotubes entrapped within. To overcome the poor reproducibility attained by dip coating approaches, other strategies for utilization of nanotubes in microelectrode fabrication was adopted. One such example is to directly grow the nanotubes onto the surface of a carbon electrode using chemical self-assembly [2]. To achieve this, the electrode is coated with a metal catalyst and then placed in functionalized carbon nanotubes to generate a uniform carbon nanotube forest on the electrode surface. The advantage of these electrodes is that they have a more regulated surface coating of nanotubes that are all aligned vertically on the electrode surface, thus providing exposure to the best surface sites of the nanotubes for electron transfer. Using this approach, significant enhancement in the sensitivity for the detection of biomolecules can be achieved. Fibers made from carbon nanotubes are a very attractive approach to make electrodes, as it solely uses the enhanced electrochemical performance of the nanotubes for electrochemical measurement. The Poulin group developed a method of making carbon nanotube fibers through polymer wet spinning [3], which is shown in Fig. 5.4. They separated carbon nanotube bundles in an aqueous surfactant solution, to overcome van der Waals forces of attraction and aggregation. The suspended nanotubes were then pushed into a streaming solution of poly(vinyl alcohol) (PVA), which displaced the surfactant and formed nanotube ribbons which subsequently collapsed in

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PVA solution

Nanotube ribbon

Rotating statge

Fig. 5.4.  Schematic of the experimental setup used to make nanotube ribbons. Flowinduced alignment of the nanotubes take place at the tip of the capillary. The ribbons formed a helical structure when the polymer solution was slowly pumped out from the bottom of the container.

air into fibers. Other polymer-CNT fibers have been developed using the wet spinning approach, such as those using poly(ethylene)imine (PEI). This polymer enhances the conductivity of the fiber, due to the intercalation of amine into bundles of SWCNTs. These approaches provide reproducible electrodes that allow for robust measurements in biological environments. The presence of the binder further provides mechanical integrity to the fiber. Another approach to make electrodes is to generate carbon yarn or fiber, through the spinning of carbon cotton. Carbon cotton is made using CVD, where ultralong individual carbon nanotubes are formed and the resultant material is analogous to conventional cotton [4]. Therefore, makes this carbon cotton very attractive as it can be spun to generate fibers. This material benefits from the formation of fibers or yarn, without the need for any other modifiers or binders to hold the nanotubes together. This therefore provides a true reflection on the electrochemical activity of nanotubes for detection of biomolecules. Fig. 5.5 shows an example of an electrode generated using carbon yarn for the measurement of neurotransmitters. Individual MWCNTs were synthesized via CVD, spun into yarns, and subsequently used in the fabrication of carbon nanotube yarn diskshaped (CNTy-D) electrodes [5]. The major benefit of this from of electrode is in the characterization of various neurochemicals. The voltammograms collected using these CNT yarn microelectrodes exhibited sharper peaks and faster apparent electron transfer kinetics, when compared to the conventional carbon-fiber microelectrodes. This is of significant benefit in bioanalysis, as being able to distinguish between chemical species provides significant advantage when resolving multiple compounds, in complex environments such as the brain. Additionally, the carbon nanotube yarn sensors exhibited greater sensitivity for the detection of neurotransmitters when compared to carbon fiber microelectrodes. The development of fibers and yarns have provided an effective way of developing reproducible microelectrodes that utilize the major properties and features of carbon

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A

B

10 µm

25 µm

C CNTy-D

1

−0.4V

CNTy-D

Conventional

1

−0.4V

1

1.4V −0.4V

−1

1

1

−0.4V

1.4V −0.4V

−1 1

−0.4V

1

1.4V −0.4V

−1

−1

H2O2

1

Adenosine

1.4V

−1

Ascorbic Acid

1

DOPAC

1.4V −0.4V

−1

Serotonin

1.4V

1.4V −0.4V

−1

Serotonin

Adenosine

1.4V −0.4V

−1

Dopamine

1

H2O2

1.4V −0.4V

−1

Conventional

1

Dopamine

Ascorbic Acid

1.4V

−1 1

1.4V −0.4V

−1

DOPAC

1.4V

−1

Fig. 5.5.  Carbon nanotube yarn microelectrodes. (A) Spun MWCNTs form a continuous yarn. (B) A single CNTy-D microelectrode. (C) Voltammograms for several analytes commonly encountered in brain tissue. Normalized data collected using CNTy-D electrodes are shown in red, and those collected using conventional carbon-fiber electrodes are shown in black. Potential was scanned from −0.4 to +1.4 V and back at 400 V/s and applied at 10 Hz to allow for the detection of many analytes. Source: Schmidt, A. C.; Wang, X.; Zhu, Y.; Sombers, L. A., Carbon Nanotube Yarn Electrodes for Enhanced Detection of Neurotransmitter Dynamics in Live Brain Tissue. ACS Nano 2013, 7 (9), 7864–7873.

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A

PEI/CNT Fiber 8 nA

8 nA

−0.4 V

1.3 V 10 Hz 100 Hz

−8 nA

−0.4 V −8 nA

CNT Yarn

1.3 V 10 Hz 100 Hz Carbon Fiber

10 nA

30 nA

−0.4 V

1.3 V 10 Hz 100 Hz

−10 nA

−0.4 V −30 nA

1.3 V 10 Hz 100 Hz

120 Normalized Ip,a/%

B

100

PEI/CNT Fiber CNT Yarn Carbon Fiber

50

CA/CNT Fiber 0

0

50 Frequency (Hz)

100

Fig. 5.6.  Exploring the effect of scan repetition frequency for 1 μM dopamine detection at different carbon nanotube electrode. (A) Example CVs of 1 μM dopamine at the scan repetition frequency of 10 Hz (blue line) and 100 Hz (orange line), with −0.4 to 1.3 V waveform. (B) Peak oxidation current at PEI/CNT fiber microelectrodes (blue circle, n = 4), CNT yarn microelectrodes (black triangle, n = 5), carbon fiber microelectrodes, (red circle, n = 5), and CA/CNT fiber microelectrodes (green triangle, n = 4), from top to bottom. Data is normalized to dopamine oxidation signal at different microelectrodes with scan repetition frequency of 10 Hz. Source: Yang, C.; Trikantzopoulos, E.; Jacobs, C. B.; Venton, B. J., Evaluation of carbon nanotube fiber microelectrodes for neurotransmitter detection: Correlation of electrochemical performance and surface properties. Analytica chimica acta 2017, 965, 1–8.

nanotubes. Fig. 5.6 shows a study which compares the performance of these varying types of carbon nanotube electrodes. In this study spun poly(vinyl alcohol)/carbon nanotube (PVA/CNT); poly(ethylene)imine/carbon nanotube (PEI/CNT), chlorosulfonic acid/carbon nanotubes, and carbon nanotube yarn microelectrodes were compared [6]. The carbon nanotube yarn microelectrodes provided the highest sensitivity and fastest electron transfer kinetics, because of the well-aligned, high purity CNTs, abundant oxygen containing functional groups at the surface, and small crevices that were critical in trapping dopamine for short times. CA/CNT fibers provided high

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selectivity to dopamine over uric acid and ascorbic acid, because of their negatively charged surface. In addition, CNT yarn and PEI/CNT fiber microelectrodes showed good stability. However, the larger crevices on the PEI/CNT fibers lead to a slower time response, illustrating that too much roughness can be detrimental to a fast temporal response. There is no doubt that the introduction and use of carbon nanotubes in the development of electrodes for bioanalysis has had considerable impact. Such electrodes have provided significant enhancements on sensitivity and selectivity due to the fast electron transfer kinetics observed at the carbon nanotube surface. Moreover, this significantly enhances the stability of the electrode, making them less prone to fouling. All these advantages have played a major role in augmenting the number of molecules and areas that electrochemical measurements can be used for in bioanalysis.

5.2.2  Boron-doped diamond Often BDD electrodes are fabricated by the deposition of a thin film of diamond on several substrates (e.g. silicon, quartz, platinum, tungsten, molybdenum, titanium, tantalum, niobium and carbon) by either hot filament or microwave plasma-assisted chemical vapor deposition (MPCVD). Regardless of the type of substrate used, the surface of substrate must be seeded first, usually by being polished with diamond powder or ultrasonicated in a suspension of diamond nanoparticles. This process is essential, as these particles of diamond that are on the substrate act as nucleation sites where the diamond film can grow. To date, there are two procedures for preparing nanocrystalline boron-doped diamond (BDD) thin films. One procedure, developed by the Naval Research Laboratory group, involves the use of a conventional CH4/H2 source gas mixture, to produce nanocrystalline diamond (NCD) [7]. Another method, developed by the Argonne National Laboratory group, involves the use of a CH4/H2/Ar source gas mixture, resulting in the formation of(ultra)nanocrystalline diamond (UNCD) [8,9]. These UNCD films are very different from the NCD films as the former possess a smaller grain size, more grain boundaries, and sp2-bonded carbon atoms within the grain boundaries. Fig. 5.7 shows a typical schematic of a MPCVD system. The diamond films are typically deposited using a CH4/H2/Ar source gas mixture under microwave power and system pressure. Boron doping is accomplished by adding B2H6 to the source gas.

5.2.2.1  Fabrication of BDD electrodes The first report on the preparation of a diamond microelectrode used sharpened tungsten wire as a substrate [10], and thus has been the basis for the development of BDD electrodes using MPCVD [11,12]. However another approach is to use a Pt micrometer wire to prepare BDD microelectrodes by MPCVD [13]. In this process, the Pt wire was first etched electrochemically to form a conical shape. The etched Pt wire is then ultrasonically seeded in a diamond powder suspension. During this seeding process, the surface of the Pt wire is scratched by the diamond particles causing

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Electrochemistry for Bioanalysis

Microwave generator Mass flow controllers

Quartz chamber

CH4 cylinder

Plasma

H2 cylinder

Substrate B2H6 cylinder

Substrate holder

Gas inlet

Pressure gauge

Exhaust value Rotary pump

Fig. 5.7.  Schematic of a microwave plasma chemical vapour deposition (MPCVD) system used for growing nanocrystalline boron-doped diamond thin films.

some particles to get embedded. Both the scratches and the embedded particles likely serve as the initial nucleation sites for film growth. A high instantaneous nucleation density is desired, as this leads to the formation of a continuous film in the shortest time with a low nominal thickness. The Pt wires were then placed in the center of the reactor's molybdenum substrate stage and the thin film of boron-doped diamond was deposited with 4 ~ 5 ppm of diborane. The resultant diamond microelectrode is then attached to a Cu wire with conductive epoxy and sealed in polypropylene. Fig. 5.8 shows an example of a conical BDD microelectrode. BDD microelectrodes provide unique characteristics compared to other carbonbased microelectrodes. These are due to the fact that diamond is sp3 hybridized rather than sp2 hybridized, which is the case for all the other forms of carbon allotropes. BDD

Fig. 5.8.  Conical BBD microelectrode.

Novel sensing materials and manufacturing approaches

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microelectrodes are attractive for electroanalytical measurements in complex media, like biological environments, because of its (i) hard and lubricious nature that enables easy penetration into tissue with minimal peripheral damage, (ii) low and stable background current over a wide potential range, (iii) good chemical and microstructural stability, (iv) low surface oxygen content (when H-terminated) – a property that leads to minimal change in the background current with variation in solution pH, (v) chemical inertness and (iv) a non-polar, hydrophobic surface which renders it resistant to molecular adsorption (i.e., deactivation and fouling). Another purported property of diamond is biocompatibility. Fig. 5.9 shows an example for the long-term stability of BDD microelectrodes for the detection of norepinephrine over a period of 8 h [13]. The current response of the carbon fiber was attenuated far greater than that of the BDD electrode, which did reduce from initial. This reduction is believed to be due to the imperfections in the BDD film, where small pockets of sp2 carbon are preferentially fouled during measurements. Although BBD microelectrodes have considerably strong benefits for long-term bioanalytical measurement, there is one limitation. Electrochemical studies of various neurotransmitters at slower scan rates (