Modeling and Control of Drug Delivery Systems [1 ed.] 0128211857, 9780128211854

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Modeling and Control of Drug Delivery Systems

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Modeling and Control of Drug Delivery Systems EDITED BY

PROF. AHMAD TAHER AZAR Faculty of Computers and Artificial Intelligence, Benha University, Benha, Egypt College of Computer & Information Sciences (CCIS), Prince Sultan University, Riyadh, Saudi Arabia Editor-in-Chief, International Journal of System Dynamics Applications (IJSDA) and International Journal of Service Science, Management, Engineering, and Technology (IJSSMET), IGI Global, Hershey, PA, United States Editor-in-Chief, International Journal of Intelligent Engineering Informatics (IJIEI), Inderscience Publishers, Olney, United Kingdom Associate EiC, IEEE Systems Journal

Academic Press is an imprint of Elsevier 125 London Wall, London EC2Y 5AS, United Kingdom 525 B Street, Suite 1650, San Diego, CA 92101, United States 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States The Boulevard, Langford Lane, Kidlington, Oxford OX5 1GB, United Kingdom © 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. Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library ISBN 978-0-12-821185-4 For information on all Academic Press publications visit our website at https://www.elsevier.com/books-and-journals

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Contents 4.4 Sensitivity Analysis, 17 4.5 Possible Antiviral Therapy Strategies, 21 5 Conclusion Remarks, 23 References, 27

CONTRIBUTORS, XI PREFACE, XV

1 Hepatitis C Virus Epidemic Control Using a Nonlinear Adaptive Strategy, 1 Javad K. Mehr, Samaneh Tangestanizadeh, Mojtaba Sharifi, Ramin Vatankhah, Mohammad Eghtesad 1 Introduction, 1 2 Related Research Work, 2 3 Dynamic Model of Hepatitis C Virus Epidemic, 3 4 Nonlinear Adaptive Controller Formulation for Epidemiology of HCV, 4 4.1 Nonlinear Adaptive Control Laws, 4 4.2 Stability Proof and Adaptation Laws, 5 5 Results and Discussion, 6 5.1 System Response to Different Uncertainty Levels, 8 6 Conclusion, 9 Appendix Barbalat’s Lemma, 10 References, 10

2 Integral Sliding Mode Control of Immune Response for Kidney Transplantation, 13 Pouria Faridi, Ramin Vatankhah, Mojtaba Sharifi 1 Introduction, 13 2 Mathematical Model of Transplant Recipients, 15 3 Control Scheme, 15 3.1 Design of an Integral Sliding Mode Controller, 16 3.2 Lyapunov Stability Proof, 16 4 Simulation Results, 17 4.1 Primary Infection Case, 17 4.2 Reactivation Case, 17 4.3 Control Signals, 17

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Smart Drug Delivery Systems, 29 Fatemeh Salahpour-Anarjan, Parinaz Nezhad-Mokhtari, Abolfazl Akbarzadeh 1 Introduction, 29 2 pH-Responsive Drug Delivery Systems, 29 3 Redox-Responsive Drug Delivery Systems, 32 4 Thermoresponsive Drug Delivery Systems, 34 5 Hypoxia-Responsive Drug Delivery Systems, 37 6 Other External Stimuli-Responsive Systems, 39 7 Conclusions, 41 References, 41

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Polymeric Transdermal Drug Delivery Systems, 45 Mohammad Shahrousvand, Nadereh Golshan Ebrahimi, Hadi Oliaie, Mahsa Heydari, Mohammad Mir, Mohsen Shahrousvand 1 Introduction, 45 2 Skin Structure and Wounds, 45 3 Wound Dressings, 47 4 Different Drug Types, 47 4.1 Hydrophilic Drugs, 49 4.2 Hydrophobic Drugs, 50 4.3 Metal-Based Nanoparticles, 50 5 Thermodynamic/Kinetic of Drug Delivery Systems, 51 5.1 Parameters Affecting the Drug Release, 51 5.2 Principles of Drug-Polymer Solubility Based on Thermodynamics of Mixing, 54 5.3 The Relationship Between the Thermodynamic and Kinetic Principles of Drug Release, 58

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CONTENTS 6 Polymeric Transdermal Drug Delivery Systems, 58 6.1 Film-Based Dressings, 59 6.2 Hydrogels, 59 6.3 Hydrocolloid-Based Dressings, 60 6.4 Electrospun Nanofibers, 60 6.5 Micro/Nanoparticles, 60 6.6 Biofoams, 61 6.7 Microneedles, 62 7 Concluding Remarks, 63 References, 64

5 Stimuli-Responsive Polymers as Smart Drug Delivery Systems, 67 Mehdi Jahanbakhshi, Mohsen Shahrousvand 1 Introduction, 67 2 Polymers as Responsive Drug Delivery Systems, 67 2.1 Thermoresponsive Polymeric Drug Delivery Systems, 68 2.2 pH-Responsive Polymeric Drug Delivery Systems, 71 2.3 Biological-Responsive Polymeric Drug Delivery Systems, 71 2.4 Ultrasound-Responsive Polymeric Drug Delivery Systems, 74 2.5 Electro-Responsive Polymeric Drug Delivery Systems, 74 2.6 Other Responsive Polymeric Drug Delivery Systems, 74 3 Future Trend and Conclusion, 75 References, 75

6 Efficacy of Polymer-Based Wound Dressings in Chronic Wounds, 79 Blessing A. Aderibigbe 1 Introduction, 79 2 Types of Wounds and Healing Phase, 80 3 Wound Dressings, 81 3.1 Hydrogels, 82 3.2 Hydrocolloids, 87 3.3 Foams, 89 3.4 Films, 90 3.5 Dermal Patches, 94 3.6 Nanofibers, 99 3.7 Membranes, 100 3.8 Polymer-Drug Conjugates, 102

4 Conclusion, 104 Acknowledgments, 105 References, 105 Website References, 110

7 Recent Progress of Transdermal Drug Delivery Systems for Biomedical Applications, 111 Jobin Jose, Iola Sandria Rodrigues, H.S. Preetha, Kiran Konkody 1 Introduction, 111 2 Skin Morphology, 111 2.1 Epidermis, 112 2.2 Dermis, 112 2.3 Hypodermis, 112 3 Skin Penetration, 113 3.1 Ideal Properties of Permeation Enhancers, 114 4 Components of Transdermal Drug Delivery Systems, 114 4.1 Polymer Matrix/Drug Reservoir, 114 4.2 Drug, 114 4.3 Pressure Sensitive Adhesives (PSA), 114 4.4 Penetration Enhancer, 114 4.5 Release Liner, 115 4.6 Backing Layer, 115 4.7 Other Excipients, 115 5 Approaches for Developing TDDS, 115 5.1 Membrane Permeation Controlled TDDS, 115 5.2 Adhesive Dispersion Type, 115 5.3 Matrix-Diffusion Controlled TDDS, 116 5.4 Micro Sealed Dissolution Controlled System or Micro Reservoir Type, 116 5.5 Micro Structured Transdermal System, 116 6 Evaluation of Transdermal DDS, 116 6.1 Physicochemical Properties, 116 6.2 Evaluation of Adhesives Used in TDDS, 117 6.3 In Vitro Evaluation, 117 6.4 In Vivo Evaluation, 117 6.5 Stability Studies, 118 7 Transdermal Drug Delivery Systems in the Management of Diseases, 118 8 Miscellaneous Bio Medical Applications of TDDS, 120 9 Conclusion, 121 References, 121

CONTENTS

8 Towards the Development of Delivery Systems of Bioactive Compounds With Eyes Set on Pharmacokinetics, 125 João S. Silva, Dorinda Marques-da-Silva, Ricardo Lagoa 1 Introduction, 125 2 Drug Delivery Systems, 126 2.1 Release Mechanisms and Classes of Delivery Systems, 126 2.2 From the Release Dynamics to the Therapeutic Performance, 128 3 Issues on the Pharmacology of Natural Compounds, 130 3.1 Challenges in the Therapeutic Application, 130 3.2 Multifactorial Actions of Natural Compounds, 130 4 Pharmacokinetic Analysis, 132 4.1 Essential Concepts, 132 4.2 Pharmacokinetic Compartmental Models, 133 4.3 Application of Pharmacokinetic Models to a Prototypical Polyphenol, 134 5 Study Models and Application in Dermal Delivery, 135 5.1 Studies in Franz Cell, 135 5.2 Studies in Transwell System, 138 6 Conclusions, 140 Acknowledgments, 140 References, 140

9 Nanofiber: An Immerging Novel Drug Delivery System, 145 Dipak Kumar Sahu, Goutam Ghosh, Goutam Rath 1 Introduction, 145 2 Nanofiber as Prolonged-Drug Delivery System, 146 2.1 Wound Healing, 146 2.2 Cancer, 148 2.3 Microbial Diseases, 148 2.4 Cardiovascular Diseases, 148 2.5 Macromolecules for Miscellaneous Applications, 149 3 Clinical Applicability Challenges, 150 4 Future Perspective, 150 5 Conclusion, 151 References, 151

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10 Molecular Dynamics Simulations on Drug Delivery Systems, 153 Zahra Shariatinia 1 Introduction, 153 2 Polymer Composites/Nanocomposites as Drug Delivery Systems, 154 3 Graphene and Its Derivatives as Drug Delivery Systems, 156 4 Carbon Nanotubes and Their Derivatives as Drug Delivery Systems, 160 5 Fullerenes as Drug Delivery Systems, 164 6 DNAs as Drug Delivery Systems, 167 7 Peptides and Cell Penetrating Peptides as Drug Delivery Systems, 168 8 Proteins as Drug Delivery Systems, 170 9 Nanoparticles as Drug Delivery Systems, 171 10 Liposomes as Drug Delivery Systems, 172 11 Micelles as Drug Delivery Systems, 174 12 Conclusion, 175 Acknowledgments, 177 References, 177

11 Nanoparticle Drug Delivery: An Advanced Approach for Highly Competent and Multifunctional Therapeutic Treatment, 183 Saima Amjad, M Serajuddin 1 Introduction, 183 2 Background of Nanoparticles in Human History and Drug Development, 184 3 Types of Nanoparticles Used for Therapeutic Treatment, 185 3.1 Metal and Metal Oxide Nanoparticles, 185 3.2 Chitosan Nanoparticles, 186 3.3 Solid Lipid Nanoparticles, 186 3.4 Mesoporous Silica Nanoparticle, 187 3.5 Liposome Nanocarrier, 187 3.6 Polymeric Nanocarriers, 187 3.7 Dendrimer, 187 3.8 Polymeric Micelles Nanoparticles, 187 4 Toxicological Profile of Nanoparticles, 188 5 Conclusion and Future Development, 188 Acknowledgments, 188 References, 188

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12 Targeted Drug Delivery: Advancements, Applications, and Challenges, 195 Hossein Rahimi, Soodabeh Davaran, Hamed Nosrati, Hossein Danafar 1 Introduction, 195 2 Active Targeting, 196 2.1 Receptor-Mediated Active Targeting, 196 2.2 Peptides, 196 2.3 Folic Acid, 199 2.4 Aptamer, 201 3 Passive Targeting, 203 4 Comparison of Active and Passive Targeting, 205 5 Conclusion, 209 References, 209

13 Strategies-Based Intrathecal Targeted Drug Delivery System for Effective Therapy, Modeling, and Controlled Release Action, 213 Pravin Shende, Sharayu Govardhane 1 Introduction, 213 1.1 Outline of the Chapter, 213 2 Strategies for Intrathecal Drug Delivery, 214 2.1 Blood-Brain Barrier Disruption by Ultrasound, 214 2.2 BBB Disruption by Osmotic Mechanism, 214 2.3 Overcoming Active Efflux at the BBB, 214 2.4 Passive Diffusion of Drugs, 215 3 Emerging Trends in Intrathecal Drug Delivery, 216 3.1 Nanoparticulate Drug Carrier System, 216 3.2 Hydrogels-Mediated Drug Delivery, 220 3.3 Microbubble-Assisted Ultrasound-Based Drug Delivery, 221 3.4 Intranasal Drug Delivery, 222 3.5 Receptor-Mediated Opening, 222 3.6 Carbon Nanotubes, 223 4 Conclusion, 223 References, 224

14 Biopolymer-Based Hydrogel Wound Dressing, 227 Mona Alibolandi, Elnaz Bagheri, Marzieh Mohammadi, Elham Sameiyan, Mohammad Ramezani 1 Introduction, 227

2 Wound Dressing and Its Ideal Properties, 228 3 Wound Dressing Based on Biopolymers, 228 3.1 Dextran, 228 3.2 Collagen, 229 3.3 Chitosan, 230 3.4 Cellulose, 232 3.5 Alginic Acid, 235 3.6 Starch, 237 3.7 Gelatin, 237 3.8 Hyaluronan, 239 3.9 Keratin, 239 3.10 Silk, 240 4 Clinical Application, 241 5 Future Perspective, 243 Acknowledgments, 243 References, 243

15 Novel Controlled Release Pulmonary Drug Delivery Systems: Current updates and Challenges, 253 Daljeet S. Dhanjal, Meenu Mehta, Chirag Chopra, Reena Singh, Parvarish Sharma, Dinesh K. Chellappan, Murtaza M. Tambuwala, Hamid A. Bakshi, Alaa A.A. Aljabali, Gaurav Gupta, Srinivas Nammi, Parteek Prasher, Kamal Dua, Saurabh Satija 1 Introduction, 253 2 Background, 254 2.1 Global Scenario, 254 2.2 Anatomy and Physiology of the Lungs, 255 3 Methods, 255 3.1 Mechanism of Drug Administration, 255 4 Nanocarrier Drug Delivery Systems, 256 4.1 Advantages of Nanocarrier Drug Delivery System, 256 5 Drug Delivery Approaches for Pulmonary Respiratory Disease, 257 5.1 Liposomes, 257 5.2 Niosomes, 259 5.3 Nanoparticles, 259 5.4 Polymeric Nanoparticles, 259 5.5 Solid Lipid Nanoparticles, 260 5.6 Dendrimers, 260 5.7 Micelles, 261 5.8 Micro-emulsions, 261 5.9 Carbon Nanotubes, 262 5.10 Quantum Dots, 262 6 Future Directions, 264 7 Conclusion, 265 References, 265

CONTENTS

16 Nanoparticle Formulations and Delivery Strategies for Sustained Drug Release in the Lungs, 273 María L. Cuestas, Tomás Brito Devoto, María A. Toscanini, María J. Limeres, Germán A. Islán, Guillermo R. Castro 1 Introduction, 273 2 Benefits and Drawbacks of the Pulmonary Route Over Other Administration Routes, 275 3 Marketed Inhalable Products and Patient Compliance, 280 3.1 Inhalable Drugs Commercially Available in US and UE Markets, 280 3.2 Patient Compliance to Aerosol Therapy, 281 4 The Role of Formulation for Controlled PDD, 284 5 The Role of Inhaler Devices for Controlled PDD, 287 5.1 pMDIs, 287 5.2 Accessories of the pMDIs, 289 5.3 Breath-Actuated MDI Devices (BA-MDIs), 289 5.4 DPIs, 289 5.5 Nebulizers, 290 5.6 SMIs, 290 6 Nanobiotechnology Solutions Against Asthma and COPD, 290 7 Nanobiotechnology Solutions Against Pulmonary Infections and Cancer, 294 8 Conclusions, 295 References, 295

17 Current Perspectives on Mycosynthesis of Nanoparticles and Their Biomedical application, 301 Suriya Rehman, Mohammad Azam Ansari, Hanan A. Al-Dossary, Zeeshan Fatima, Saif Hameed, Wasim Ahmad, Abuzar Ali 1 Introduction, 301 2 Microbial Green synthesis: A Novel and Eco-friendly Approach, 301 2.1 Fungi, An Efficient System for the Biosynthesis of NPs, 301 2.2 The Probable Mechanism of Myconanoparticles Synthesis, 302 2.3 Achieving Different Sizes of Myconanoparticles, 302 2.4 Role of NPs in the Treatment of Infectious Diseases, 303

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3 Mycosynthesis of Various NPS and Their Biomedical Applications, 304 3.1 Silver Myconanoparticles Synthesis, 304 3.2 Application of Myco-synthesized Silver NPs, 304 3.3 Gold Myconanoparticles (AuNPs), 304 3.4 Application of Myco-synthesized Gold NPs, 304 3.5 Other Metal and Metal Oxide Myconanoparticles and Their Applications, 305 4 Conclusion and Future Prospects, 308 Acknowledgment, 308 References, 308

18 Solid Oral Controlled-Release Formulations, 313 Mitesh Bhansali, Neha Dabholkar, P. Swetha, Sunil Kumar Dubey, Gautam Singhvi 1 Introduction, 313 2 Need for Controlled-Release Dosage Forms, 314 3 Terminologies Used for Describing Controlled-Release Formulations, 315 3.1 Conventional Release Dosage Form, 316 3.2 Modified Drug-Release Dosage Form, 316 3.3 Prolonged-Release Dosage Form, 316 3.4 Controlled-Drug Release Dosage Form, 316 3.5 Delayed-Drug Release Dosage Form, 316 4 Polymers used in Controlled-Release Systems, 316 4.1 Hydrophilic Polymers, 318 4.2 Hydrophobic Polymers, 319 5 Types of Controlled-Release Drug Delivery Systems, 320 5.1 Diffusion-Controlled Systems, 320 5.2 Dissolution-controlled systems, 320 5.3 Dissolution-diffusion-controlled systems or hybrid systems, 321 6 Drug Release Characterization From Controlled-Drug Delivery Systems, 322 6.1 Evaluation of Drug Release Characteristics From Delivery Systems, 322 6.2 Swelling and erosion characterization for controlled-release dosage form, 326 6.3 Pharmacokinetic Evaluation of Solid Oral-Controlled Dosage Forms, 327

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CONTENTS 7 Conclusion, 328 8 Future Prospects, 328 References, 328

19 Advanced Solid Oral ControlledRelease Formulations, 333 Neha Dabholkar, P. Swetha, Mitesh Bhansali, Sunil Kumar Dubey, Gautam Singhvi 1 Introduction, 333 2 Gastro-Retentive Drug Delivery Systems, 333 2.1 Floating Systems, 335 2.2 Bioadhesive or Mucoadhesive Systems, 336 3 Colon-Targeted Drug Delivery Systems, 337 3.1 pH-Sensitive Drug Delivery, 338 3.2 Delayed or Time-Controlled Release Drug Delivery Systems, 339 3.3 Microbially Targeted Colonic Delivery, 339 3.4 Integrated Approaches for Colon-Targeted Delivery Systems, 339 4 Feedback-Regulated Systems, 340 4.1 Bioresponsive Systems, 340 4.2 Self-Regulating Systems, 340 5 Enteric Drug Delivery Systems, 340 5.1 Reasons for Enteric Coating, 340 5.2 Polymers Used for Enteric Coating and Its Mechanism, 341 6 Osmotic Drug Delivery Systems, 341 6.1 Mechanism Involved in Drug Release From Osmotic Drug Delivery Systems, 342 6.2 Factors Affecting the Osmotic Drug Delivery Systems, 342 6.3 Advantages of Osmotic Drug Delivery Systems, 344 7 3D Printing-Based Controlled-Release Formulations, 344 8 Ultra-Long Acting Formulations, 345 9 Patented Technologies, 345 10 Future Perspective, 346 11 Conclusion, 347 References, 347

20 Mucoadhesive Polymers: Gateway to Innovative Drug Delivery, 351 Muhammad Yaqoob, Aamir Jalil, Andreas Bernkop-Schnu€rch 1 Introduction, 351 2 Mechanism of Mucoadhesion, 351 3 Mucus Gel Layer, 352 3.1 Mucins, 352 3.2 Production, 352 4 Evaluation of Mucoadhesive Properties, 355 4.1 Force Determination Methods, 355 4.2 Molecular Interaction Methods, 356 4.3 Rinse Methods, 358 4.4 Cellular Methods, 359 4.5 Optical Methods, 359 5 Mucoadhesive Polymers, 359 5.1 Classification of Mucoadhesive Polymers, 360 5.2 Classification Based on the Source of Polymer, 361 6 Factors Affecting Mucoadhesion, 369 6.1 Polymer Backbone, 369 6.2 Physiological Factors, 370 7 Mucoadhesive Drug Delivery Systems, 371 7.1 In Situ Gelling Formulations, 371 7.2 Electrospun Nanofibers, 371 7.3 Mucoadhesive Nanoparticles, 371 7.4 Films, 372 7.5 Tablets, 372 7.6 Beads, 372 7.7 Mucoadhesive Microsphere, 373 7.8 Polymeric Micelles, 373 7.9 Polymer-Coated Liposomes, 373 7.10 Self-Emulsifying Drug Delivery System, 373 8 Marketed Mucoadhesive Products, 373 9 Future Perspectives of Mucoadhesive Drug Delivery Systems, 373 10 Conclusion, 375 References, 375 INDEX, 385

Contributors Blessing A. Aderibigbe Department of Chemistry, University of Fort Hare, Eastern Cape, South Africa Wasim Ahmad Department of Pharmacy, Mohammad Al-Mana College for Medical Sciences, Dammam, Saudi Arabia Abolfazl Akbarzadeh Department of Medical Nanotechnology, Faculty of Advanced Medical Sciences; Stem Cell Research Center, Tabriz University of Medical Sciences, Tabriz, Iran Hanan A. Al-Dossary Department of Epidemic Disease Research, Institute for Research and Medical Consultations (IRMC), Imam Abdulrahman Bin Faisal University, Dammam, Saudi Arabia Abuzar Ali College of Pharmacy, Taif University, Taif, Saudi Arabia Mona Alibolandi Pharmaceutical Research Center, Pharmaceutical Technology Institute; Department of Pharmaceutical Biotechnology, School of Pharmacy, Mashhad University of Medical Sciences, Mashhad, Iran Alaa A.A. Aljabali Faculty of Pharmacy, Department of Pharmaceutics and Pharmaceutical Technology, Yarmouk University, Irbid, Jordan Saima Amjad Department of Zoology, University of Lucknow, Lucknow, India Mohammad Azam Ansari Department of Epidemic Disease Research, Institute for Research and Medical Consultations (IRMC), Imam Abdulrahman Bin Faisal University, Dammam, Saudi Arabia

Elnaz Bagheri Pharmaceutical Research Center, Pharmaceutical Technology Institute; Department of Pharmaceutical Biotechnology, School of Pharmacy, Mashhad University of Medical Sciences, Mashhad, Iran Hamid A. Bakshi School of Pharmacy and Pharmaceutical Sciences, Ulster University, Coleraine, Northern Ireland, United Kingdom € rch Andreas Bernkop-Schnu Center for Chemistry and Biomedicine, Department of Pharmaceutical Technology, Institute of Pharmacy, University of Innsbruck, Innsbruck, Austria Mitesh Bhansali Department of Pharmacy, Birla Institute of Technology and Science, Pilani, Rajasthan, India Guillermo R. Castro Laboratory of Nanobiomaterials, CINDEFI, Department of Chemistry, Faculty of Exact Sciences, National University of La Plata-CONICET (CCT La Plata), La Plata; Max Planck Laboratory for Structural Biology, Chemistry and Molecular Biophysics of Rosario (MPLbioR, UNR-MPIbpC), Partner Laboratory of the Max Planck Institute for Biophysical Chemistry (MPIbpC, MPG), Center for Interdisciplinary Studies (CEI), National University of Rosario, Rosario, Santa Fe, Argentina Dinesh K. Chellappan Department of Life Sciences, School of Pharmacy, International Medical University, Bukit Jalil, Kuala Lumpur, Malaysia Chirag Chopra School of Bioengineering and Biosciences, Lovely Professional University, Phagwara, Punjab, India María L. Cuestas University of Buenos Aires, CONICET, Institute for Research in Microbiology and Medical Parasitology (IMPaM), Buenos Aires, Argentina xi

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CONTRIBUTORS

Neha Dabholkar Department of Pharmacy, Birla Institute of Technology and Science, Pilani, Rajasthan, India Hossein Danafar Department of Pharmaceutical Biomaterials, School of Pharmacy, Zanjan University of Medical Sciences, Zanjan, Iran Soodabeh Davaran Drug Applied Research Center, Tabriz University of Medical Sciences, Tabriz, Iran Tomás Brito Devoto University of Buenos Aires, CONICET, Institute for Research in Microbiology and Medical Parasitology (IMPaM), Buenos Aires, Argentina Daljeet S. Dhanjal School of Bioengineering and Biosciences, Lovely Professional University, Phagwara, Punjab, India Kamal Dua Discipline of Pharmacy, Graduate School of Health, University of Technology Sydney, Ultimo; Priority Research Centre for Healthy Lungs, Hunter Medical Research Institute (HMRI) & School of Biomedical Sciences and Pharmacy, University of Newcastle, Callaghan, NSW, Australia Sunil Kumar Dubey Department of Pharmacy, Birla Institute of Technology and Science, Pilani, Rajasthan, India Nadereh Golshan Ebrahimi Polymer Engineering Department, Chemistry Engineering Faculty, Tarbiat Modares University, Tehran, Iran Mohammad Eghtesad Department of Mechanical Engineering, Shiraz University, Shiraz, Iran Pouria Faridi Department of Mechanical Engineering, Shiraz University, Shiraz, Iran; Department of Medicine and Dentistry, University of Alberta, Edmonton, AB, Canada Zeeshan Fatima Amity Institute of Biotechnology, Amity University Haryana, Gurugram, India

Goutam Ghosh School of Pharmaceutical Sciences, Siksha 'O' Anusandhan (Deemed to be University), Bhubaneswar, Odisha, India Sharayu Govardhane Shobhaben Pratapbhai Patel School of Pharmacy and Technology Management, SVKM’S NMIMS, Mumbai, India Gaurav Gupta School of Phamacy, Suresh Gyan Vihar University, Jaipur, India Saif Hameed Amity Institute of Biotechnology, Amity University Haryana, Gurugram, India Mahsa Heydari Polymer Engineering Department, Chemistry Engineering Faculty, Tarbiat Modares University, Tehran, Iran Germán A. Islán Laboratory of Nanobiomaterials, CINDEFI, Department of Chemistry, Faculty of Exact Sciences, National University of La Plata-CONICET (CCT La Plata), La Plata, Argentina Mehdi Jahanbakhshi Caspian Faculty of Engineering, College of Engineering, University of Tehran, Rezvanshahr, Iran Aamir Jalil Center for Chemistry and Biomedicine, Department of Pharmaceutical Technology, Institute of Pharmacy, University of Innsbruck, Innsbruck, Austria; Department of Pharmaceutics, Faculty of Pharmacy, University of Lahore, Lahore, Pakistana Jobin Jose Nitte (Deemed to be University), NGSM Institute of Pharmaceutical Sciences (NGSMIPS), Department of Pharmaceutics, Mangalore, Karnataka, India Kiran Konkody Nitte (Deemed to be University), NGSM Institute of Pharmaceutical Sciences (NGSMIPS), Department of Pharmaceutics, Mangalore, Karnataka, India

CONTRIBUTORS Ricardo Lagoa School of Technology and Management, Polytechnic Institute of Leiria, Leiria, Portugal María J. Limeres University of Buenos Aires, CONICET, Institute for Research in Microbiology and Medical Parasitology (IMPaM), Buenos Aires, Argentina Dorinda Marques-da-Silva School of Technology and Management, Polytechnic Institute of Leiria, Leiria, Portugal Javad K. Mehr Department of Electrical and Computer Engineering; Department of Medicine, University of Alberta, Edmonton, AB, Canada

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Hadi Oliaie Department of Polymer Engineering and Color Technology, Amirkabir University of Technology (Tehran Polytechnic), Tehran, Iran Parteek Prasher Department of Chemistry, University of Petroleum and Energy Studies, Dehradun, India H.S. Preetha Nitte (Deemed to be University), NGSM Institute of Pharmaceutical Sciences (NGSMIPS), Department of Pharmaceutics, Mangalore, Karnataka, India Hossein Rahimi Department of Medical Biotechnology, School of Medicine, Zanjan University of Medical Science, Zanjan, Iran

Meenu Mehta School of Pharmaceutical Sciences, Lovely Professional University, Phagwara, Punjab, India; Discipline of Pharmacy, Graduate School of Health, University of Technology Sydney, Ultimo, NSW, Australia

Mohammad Ramezani Pharmaceutical Research Center, Pharmaceutical Technology Institute; Department of Pharmaceutical Biotechnology, School of Pharmacy, Mashhad University of Medical Sciences, Mashhad, Iran

Mohammad Mir Polymer Engineering Department, Chemistry Engineering Faculty, Tarbiat Modares University, Tehran, Iran

Goutam Rath School of Pharmaceutical Sciences, Siksha 'O' Anusandhan (Deemed to be University), Bhubaneswar, Odisha, India

Marzieh Mohammadi Department of Pharmaceutics, School of Pharmacy, Mashhad University of Medical Sciences, Mashhad, Iran

Suriya Rehman Department of Epidemic Disease Research, Institute for Research and Medical Consultations (IRMC), Imam Abdulrahman Bin Faisal University, Dammam, Saudi Arabia

Srinivas Nammi School of Science and Health, Western Sydney University, Penrith, NSW, Australia Parinaz Nezhad-Mokhtari Department of Medical Nanotechnology, Faculty of Advanced Medical Sciences; Student Research Committee, Tabriz University of Medical Sciences, Tabriz, Iran Hamed Nosrati Department of Pharmaceutical Biomaterials, School of Pharmacy, Zanjan University of Medical Sciences, Zanjan, Iran

Iola Sandria Rodrigues Nitte (Deemed to be University), NGSM Institute of Pharmaceutical Sciences (NGSMIPS), Department of Pharmaceutics, Mangalore, Karnataka, India Dipak Kumar Sahu School of Pharmaceutical Sciences, Siksha 'O' Anusandhan (Deemed to be University), Bhubaneswar, Odisha, India Fatemeh Salahpour-Anarjan Department of Medical Nanotechnology, Faculty of Advanced Medical Sciences, Tabriz University of Medical Sciences, Tabriz, Iran

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CONTRIBUTORS

Elham Sameiyan Department of Pharmaceutical Biotechnology, School of Pharmacy, Mashhad University of Medical Sciences, Mashhad, Iran Saurabh Satija School of Pharmaceutical Sciences, Lovely Professional University, Phagwara, Punjab, India; Discipline of Pharmacy, Graduate School of Health, University of Technology Sydney, Ultimo, NSW, Australia M. Serajuddin Department of Zoology, University of Lucknow, Lucknow, India Mohammad Shahrousvand Polymer Engineering Department, Chemistry Engineering Faculty, Tarbiat Modares University, Tehran, Iran Mohsen Shahrousvand Caspian Faculty of Engineering, College of Engineering, University of Tehran, Rezvanshahr, Iran Zahra Shariatinia Department of Chemistry, Amirkabir University of Technology (Tehran Polytechnic), Tehran, Iran Mojtaba Sharifi Department of Electrical and Computer Engineering; Department of Medicine; Department of Medicine and Dentistry, University of Alberta, Edmonton, AB, Canada Parvarish Sharma School of Pharmaceutical Sciences, Lovely Professional University, Phagwara, Punjab, India Pravin Shende Shobhaben Pratapbhai Patel School of Pharmacy and Technology Management, SVKM’S NMIMS, Mumbai, India

João S. Silva School of Technology and Management, Polytechnic Institute of Leiria, Leiria, Portugal Reena Singh School of Bioengineering and Biosciences, Lovely Professional University, Phagwara, Punjab, India Gautam Singhvi Department of Pharmacy, Birla Institute of Technology and Science, Pilani, Rajasthan, India P. Swetha Department of Pharmacy, Birla Institute of Technology and Science, Pilani, Rajasthan, India Murtaza M. Tambuwala School of Pharmacy and Pharmaceutical Sciences, Ulster University, Coleraine, Northern Ireland, United Kingdom Samaneh Tangestanizadeh Department of Mechanical Engineering, Shiraz University, Shiraz, Iran María A. Toscanini University Buenos Aires, Faculty of Pharmacy and Biochemistry, Institute of Nanobiotechnology (NANOBIOTEC), Buenos Aires, Argentina Ramin Vatankhah Department of Mechanical Engineering, Shiraz University, Shiraz, Iran Muhammad Yaqoob Center for Chemistry and Biomedicine, Department of Pharmaceutical Technology, Institute of Pharmacy, University of Innsbruck, Innsbruck, Austria

Preface Drug delivery is the method or process of administering a pharmaceutical compound to achieve a therapeutic effect in humans or animals. Delivering drugs at controlled rate, slow delivery, and targeted delivery are other very attractive methods and have been pursued enthusiastically. Various drug delivery and targeting systems have been developed to minimize drug degradation and adverse effect, and to increase drug bioavailability. Site-specific drug delivery may be either an active and/or passive process. In the past few years, researchers have appreciated the potential benefits of nanotechnologies in providing vast improvements to drug delivery and targeting. Improving delivery techniques that minimize toxicity and increase efficacy offer great potential benefits to patients and also open up new markets for pharmaceutical companies. The book is intended to provide a comprehensive coverage of various drug delivery and targeting systems and their state-of-the-art related works, ranging from theory to the real-world deployment. Also, it discusses the future perspectives of DDS.

ABOUT THE BOOK The new Elsevier book, Modeling and Control of Drug Delivery Systems, consists of 20 contributed chapters by subject experts who are specialized in the various topics addressed in this book. The special chapters have been brought out in this book after a rigorous review process in the broad areas of modeling, simulation, control, and drug delivery systems. Special importance was given to chapters offering practical solutions and novel methods for the recent research problems in drug delivery systems.

OBJECTIVES OF THE BOOK This book will attract many researchers working in the area of drug delivery Systems. This book presents some of the latest innovative of approaches to DDS from a point of view of dynamic modeling, system analysis, optimization, control, and monitoring, and so on. This book will be an important source of information for pharmaceutical scientists and pharmacologists working

in the academia as well as in the industry. It has useful information for pharmaceutical physicians and scientists in many disciplines involved in developing DDS such as chemical engineering, biomedical engineering, protein engineering, gene therapy, and so on. This will be an important reference for executives in charge of research and development at several hundred companies that are developing drug delivery technologies. Thus the purpose of this book is to provide the community with a unique, recent, and comprehensive reference on DDS with the focus on cutting edge technologies and the recent research trends in the area.

ORGANIZATION OF THE BOOK This well-structured book consists of 20 full chapters.

BOOK FEATURES • The book chapters deal with the recent research problems in the areas of modeling, simulation, control, and drug delivery systems. • The book chapters present various techniques of drug delivery systems such as Hepatitis C Virus Epidemic Control Using a Nonlinear Adaptive Strategy, Integral Sliding Mode Control of Immune Response for Kidney Transplantation, Smart Drug Delivery Systems, Polymeric Transdermal Drug Delivery Systems, Nanoparticle Drug Delivery, Targeted Drug Delivery, and so on. • The book chapters contain a good literature survey with a long list of references. • The book chapters are well-written with a good exposition of the research problem, methodology, block diagrams, and mathematical techniques. • The book chapters are lucidly illustrated with numerical examples and simulations. • The book chapters discuss details of different applications and future research areas.

AUDIENCE • The book is primarily meant for researchers from academia and industry, who are working on modeling, xv

xvi

PREFACE

simulation, control and drug delivery research areas—biomedical engineering, control engineering, computer engineering, computer science, nanotechnology, applied sciences, pharmacy, and medicine. The book can also be used at the graduate or advanced undergraduate level as a text book or major reference for courses such as Concepts of Drug Delivery, Advanced Drug Delivery systems, Controlled Drug delivery systems, and many others.

ACKNOWLEDGMENTS • As the editors, we hope that the chapters in this wellstructured book will stimulate further research in

modeling, simulation, control and drug delivery systems, and utilize them in real-world applications. We hope sincerely that this book, covering so many different topics, will be very useful for all readers. We thank all the reviewers for their diligence in reviewing the chapters. Special thanks to Elsevier, especially the book Editorial team. Ahmad Taher Azar Faculty of Computers and Artificial Intelligence, Benha University, Benha, Egypt College of Computer & Information Sciences (CCIS) Prince Sultan University, Riyadh, Saudi Arabia

CHAPTER 1

Hepatitis C Virus Epidemic Control Using a Nonlinear Adaptive Strategy

JAVAD K. MEHRa,b • SAMANEH TANGESTANIZADEHc • MOJTABA SHARIFIa,b • RAMIN VATANKHAHc • MOHAMMAD EGHTESADc a

Department of Electrical and Computer Engineering, University of Alberta, Edmonton, AB, Canada, Department of Medicine, University of Alberta, Edmonton, AB, Canada, cDepartment of Mechanical Engineering, Shiraz University, Shiraz, Iran b

1. INTRODUCTION The hepatitis C virus (HCV) is a blood-borne virus identified as the main cause of liver diseases [1–3]. Globally, about 3% of the world population (170 million) are dealing with HCV and 71 million people have chronic hepatitis C infection [1, 4–6]. Several studies showed that the chronic stage of HCV will develop cirrhosis and liver cancer in the case of no treatment and approximately 339,000 people die every year due to these diseases [1, 7]. Despite previously mentioned statistics which makes HCV infection one of the important health threats, this disease received little attention especially in the regions with a higher rate of infectiousness [4]. Although fatigue and jaundice were mentioned as symptoms of the HCV, this disease often has no considerable symptom, even in the advanced stages. This is the reason that the HCV outbreak is called “the silent epidemic” [4, 8]. Several different ways were reported for HCV prevalence, which includes sharing injection equipment, unsafe sexual contacts, inadequately sterilization of syringes and needles especially for health-care personnel, and transfusion of polluted blood [1, 9]. Even though these are the main causes of the HCV epidemic, but some other reasons may also be critical in some societies based on special conditions. For instance in the developed countries, since there is precise control on the blood transfusion, the importance of injecting drug use in transmission of the disease has increased compared to the transfusion of polluted blood and its products [2, 9]. Natural cure at the chronic stage of HCV is not common, but it can happen for about 10%–15% of patients that the RNA of HCV is indistinguishable in their serum [5, 6]. For the rest of the patients (80%–85%) that the HCV could not be healed by their immune system response, some drug therapy regimes should be employed. Hepatitis C drugs have recently had some

developments. Available safe, highly effective, and endurable combinations of oral antivirals that act directly have currently developed for this disease [4, 10]. Although vaccination is the most vital way of controlling different viral diseases, but unfortunately there is no vaccine for the HCV yet [5]. Therefore, preventing this disease has an important role in stopping the extension of its epidemic. In the present study, a nonlinear adaptive method is developed for treatment and control of the HCV epidemic. For this purpose, the recently published nonlinear HCV epidemiological model in [4] is employed and different parametric uncertainties are taken into account, despite the previous optimal strategies [4]. The main goal of the proposed control scheme is the population decrease in the unaware susceptible and chronically infected compartments in the existence of parametric uncertainties. Accordingly, two control inputs (efforts to inform susceptible individuals and treatment rate) are employed to track descending desired populations of the previously mentioned compartments. The asymptotic stability and tracking convergence of the closed-loop system having uncertainties are proven using the Lyapunov stability theorem and Barbalat’s lemma. Innovations of this research are as follows: • For the first time, a nonlinear adaptive method is developed to control the HCV epidemic by defining a novel Lyapunov function candidate that provides the tracking convergence proof. • Due to the lack of accurate information about HCV model parameters in each society, parametric uncertainty is taken into account in this research for the first time, and the defined control objectives are achieved in the presence of these inaccuracies. • In all of the previous studies that have been conducted on the control of the HCV outbreak,

Modeling and Control of Drug Delivery Systems. https://doi.org/10.1016/B978-0-12-821185-4.00016-6 © 2021 Elsevier Inc. All rights reserved.

1

2

Modeling and Control of Drug Delivery Systems

populations of some undesired compartments at the end of the investigation period were considered as the criterion for designing control inputs [4, 11–14]. However, in this study for the first time, the populations of two unaware susceptible and chronically infected classes during the entire treatment period are considered as the criterion and control inputs are designed in a way to track desired values instead of focusing on their final populations at the end of process. The rest of this chapter is organized as follows. In Section 2, related research work will be explained. Description of the dynamic model and the proposed control scheme will be presented in Sections 3 and 4, respectively. The simulation results will be depicted and discussed in Section 5, and the concluding remarks will be mentioned in Section 6.

2. RELATED RESEARCH WORK Previous related studies are presented in this section and are divided into three parts, including mathematical modeling, optimal control for HCV, and adaptive control strategies for different biological systems. Several analytical analyses were conducted on the dynamic modeling of the HCV epidemic, which are described here. Martcheva and Castillo-Chavez [15] presented a simple mathematical model with three compartmental variables including susceptible, acutely infected, and chronically infected. They considered different epidemiological observations in the model. Yuan and Yang [8] added the exposed class to the previous model [15]. They considered that the susceptible individuals transmit to the exposed compartment in the case of having contact with the infected compartments. Zhang and Zhou [5] added a new term in the model of Yuan and Yang [8], which denotes the death rate due to the HCV. Hu and Sun [16] proposed another epidemiological model for the HCV with four classes in which the recovered compartment was taken into account for the first time. Naturally, the recovered people transmit to this class from the acutely infected and chronically infected compartments and become immune against this. Ainea et al. [17] extended the previous model [16] by adding the exposed class. Both these models [16, 17] considered the HCV diseaseinduced death rate for both acutely infected and chronically infected classes. Shen et al. [18] proposed a dynamical model with six classes including susceptible, exposed, acutely infected, chronically infected, treated, and recovered populations. They propounded treatment influence for the first time and classified treated people in a distinct class. Shi and Cui [19] improved

the model in [18] and divided the treated class into two different classes by defining the treatment for chronic infection and aware reinfection. Some researches have been conducted for optimal control of the HCV outbreak. Okosun [11] employed a SITV (susceptible, acutely infected, treated, and chronically infected) model for the HCV that was an extended form of the dynamics presented in [8]. This model [11] included the treatment compartment and considered movement for susceptible, treated, and acutely and chronically infected people among their compartments. Some time-dependent optimal control strategies are proposed, in order to control the HCV disease. A cost function is calculated for these strategies in order to evaluate the effectiveness of the control methods and select the most efficient one. Okosun and Makinde [12] employed a SEITV (susceptible, exposed, acutely infected, treated, and chronically infected) dynamical model for the HCV outbreak considering the screening rate and drug efficacy as control inputs for acutely and chronically infected populations and used the Pontryagin’s principle to solve the optimal control problem. Another epidemiological model was investigated in [4] for the HCV outbreak in which the susceptible class was divided into aware and unaware classes. Moreover, they considered two control inputs including screening and treatment rates for the HCV epidemic model, which was determined by an optimal control law. In [4], the dynamics was formulated with the susceptibility reduction due to the publicity and the treatment process to identify the feasible effect of public concerns and treatment on the HCV. An optimal neurofuzzy strategy was also introduced in [13] in order to control the HCV epidemic. They [13] employed the mathematical model proposed in [12] as a deterministic model and utilized the genetic algorithm to obtain optimal control inputs. As described, all of previous studies on the control of HCV epidemic were conducted on the optimal strategies. On the other hand, some other research works were performed on the adaptive control of different diseases as presented here. Moradi et al. [20] suggested a Lyapunov-based adaptive method to control three different hypothetical models of the cancer chemotherapy inside the human body and compared results among these models. In the next step of this research [21], a composite adaptive strategy has been developed for online identification of cancer parameters during the chemotherapy process. Boiroux et al. [22] employed a model predictive controller for the type 1 diabetes model and used an adaptive controller to balance the blood glucose. They determined the model parameters

CHAPTER 1

Hepatitis C Virus Epidemic Control Using a Nonlinear Adaptive Strategy

based on the clinical information of past patients. Aghajanzadeh et al. [23] suggested an adaptive control strategy for hepatitis B virus infection inside the human body by antiviral drugs. They considered model parameters uncertainties on model parameters and employed adaptive controller to control the dynamic despite uncertainties of the system. Sharifi and Moradi [24] designed a robust scheme with adaptive gains to control the influenza epidemic, considering its dynamic model’s uncertainties. Padmanabhan et al. [25] proposed an optimal adaptive method to control the sedative drug in anesthesia administration. They employed an integral reinforcement learning method in order to overcome the uncertainty of parameter values.

3. DYNAMIC MODEL OF HEPATITIS C VIRUS EPIDEMIC Mathematical modeling is a useful way of analyzing the epidemiology of a disease. These models have two important capabilities: (1) finding out mechanistic understanding of the disease and (2) exploring potential outcomes of the epidemic under different conditions [26]. For assessment of the proposed method for the HCV prevalence control in a population, a nonlinear compartmental model is used with five different classes including unaware susceptible (Su), aware susceptible (Sa), acutely infected (I), chronically infected (C), and the treated (T) humans [4]. The susceptible compartment is divided into two classes, including aware and unaware people. Note that aware people have information about the HCV transmission ways and preventing methods despite the unaware population. Since there is no available vaccine for the HCV, informing people about preventing methods is a very important way to reduce the risk of infection for susceptible people [1]. Therefore, the unaware susceptible individuals (Su) will be infected in contact with the infected population (I, C, and T) with a higher rate in comparison with the aware susceptible individuals (Sa) [4]. Thus, the transmission rate for unaware susceptible humans (Su) should be considered larger than this rate for aware susceptible humans (Sa) in the dynamic model [4]. The nonlinear mathematical model of HCV epidemic is as follows: Su S_u ¼ b  λSu  ðμ + u1 ðtÞÞSu + ð1  qÞγI N Sa _ Sa ¼ u1 ðtÞSu  λSa  μSa + ð1  pÞξT N S S u a I_ ¼ λSu + λSa  ðμ + γÞI N N _ C ¼ qγI  ðμ + u2 ðtÞ + θÞC + pξT T_ ¼ u2 ðtÞC  ðμ + ξÞT

3

where λSu ¼ βðI + K1 C + K2 TÞ and λSa ¼ αλSu . u1 and u2 are control inputs and defined respectively as the effort rate to inform unaware susceptible individuals and the treatment rate for chronically infected class. N denotes the total population and will be calculated as N ¼ Su + Sa + I + C + T

(2)

The population of unaware susceptible (Su) increases with the rate of b. Unaware and aware susceptible individuals are also infected in contact with acutely and chronically infected and treated individuals at the rates of λSu and λSa , respectively. Infectiousness rate for acutely infected people is higher than chronically infected individuals, and the treated people have the lowest rate; thus, it is assumed that K1 > K2 [4, 5]. The total population (N) decreases with two different rates μ and θ, where μ denotes the rate of natural death that decreases populations of all compartments. However, θ is the rate of HCV-induced death and decreases the population of the chronically infected compartment (C). During the acute stage (I), the HCV could have different behaviors for each patient based on his/her immune system response. For 15%–25% of cases in this stage, the RNA of HCV becomes indistinguishable in their blood serum and the ALT level returns to the normal range. This observation is defined by the term (1  q)γI in the proposed HCV dynamics [4, 6]. Approximately, the immune system in 75%–85% of the patients could not remove the hepatitis C virus in the acute stage and their disease becomes advanced to the chronic stage. Note that if the HCV RNA remains in the patient’s blood for at least 6 months after the onset of acute infection, the chronic level of the disease will appear which is defined by the term qγI in Eq. (1) [5, 6]. Finally, the defeat in the treatment process is defined by the term p. The treated population decreases by the rate of ξT and joins the chronic class by the rate of pξT in the case of treatment failure and the rest of this population (1  p)ξT will join the aware susceptible class if the treatment is successful. The schematic diagram of the proposed nonlinear dynamics of the HCV epidemic is depicted in Fig. 1 and descriptions of the parameters are presented in Table 1 [4].

(1)

FIG. 1 Schematic diagram of population transmission

among different classes of HCV epidemic.

4

Modeling and Control of Drug Delivery Systems TABLE 1

Parameters of the Mathematical Model of the HCV (1) [4]. Parameter

Description

b

Rate of birth

μ

Rate of death

β

Transmission coefficient

K1

Chronic stage infectiousness relative to acute stage

K2

Treated individuals’ infectiousness relative to acute ones

α

Rate of being infected for aware people relative to unaware ones

γ

Leaving rate of acutely infected class

q

Progressing proportion from acute stage to chronic one

ξ

Transferring rate from treated class to other ones

p

Moving back proportion from treated class to chronic one

θ

HCV-induced death rate

4. NONLINEAR ADAPTIVE CONTROLLER FORMULATION FOR EPIDEMIOLOGY OF HCV In this section, a new nonlinear adaptive controller is formulated for the uncertain hepatitis C virus epidemic. The main purpose of the control method is to minimize the populations of unaware susceptible (Su) and chronically infected (C) classes. Two control inputs u1(t) and u2(t) are considered in order to reach this objective. u1(t) denotes the effort rate to inform the susceptible individuals from the HCV by media publicity, educational campaigns, public service advertising, and so on, and u2(t) is employed to reflect the rate of treatment on chronically infected individuals [4]. Using the above-mentioned control inputs, the populations of unaware susceptible (Su) and chronically infected (C) classes will decrease by tracking some desired values. Moreover, due to the decrease of the mentioned components, the number of aware susceptible (Sa) and treated (T) individuals will increase and decrease, respectively. The Lyapunov theorem is employed to prove the stability of the closed-loop system. In addition, some adaptation laws are defined in order to update the estimated parameters of the system to guarantee the stability and robustness of the system against the uncertainties of the dynamic model. A

conceptual diagram of the proposed nonlinear feedback controller with the adaptive scheme is illustrated in Fig. 2.

4.1. Nonlinear Adaptive Control Laws Control inputs (u1(t), u2(t)) could be calculated using dynamics of the unaware susceptible and chronically infected compartments from Eq. (1) as u1 ¼ 

S_u b β I +  ðI + K1 C + K2 TÞ  μ + ð1  qÞγ Su Su Su N

(3)

C_ I T + qγ  ðμ + θÞ + pξ C C C

(4)

u2 ¼ 

Property. The right-hand sides of Eqs. (3), (4) can be linearly parameterized in terms of their available parameters. ϕ1 and ϕ2 are considered to be the arbitrary variables instead of _ S_u and C. Now, the right sides of the above equations can be rewritten as 

S_u b β I ϕ +  ðI + K1 C + K2 TÞ  μ + ð1  qÞγ ¼  1 + Y1 θ 1 Su Su Su N Su (5) 

C_ I T ϕ + qγ  ðμ + θÞ + pξ ¼  2 + Y2 θ2 C C C C

(6)

where Y1 and Y2 are the regressor matrices, contain known functions of HCV epidemic variables. θ1 and θ2 are the parameter vectors, which contain unknown parameters of the dynamic (Eqs. 7, 8). Accordingly, these matrices and vectors are defined as
Y1 ¼

 1 I C T I    1 ; Su N N N Su

θ1 ¼ ½ b β βK1 βK2 ð1  qÞγ μ T
Y2 ¼

 I T T 1 ; θ2 ¼ ½ qγ pξ ðμ + θÞ  C C

(7) (8)

This regressor presentation is used to summarize the equations and define the adaptation and control laws. In order to design nonlinear control laws, two new variables ϕ1 and ϕ2 are defined as follows: 

(9)



(10)

ϕ1 ¼ S_ ud  λ1 S u ϕ2 ¼ C_ d  λ2 C

where λ1 and λ2 are the controller gains and considered to be positive and constant. Now, the nonlinear adaptive control laws are defined as 

S_ud  λ1 Su

+ Y1 θ^1

(11)

_ u2 ¼  Cd Cλ2 C + Y2 θ^2

(12)

u1 ¼ 

Su 

where θ^1 and θ^2 are the vectors of estimated parameters.

CHAPTER 1

Hepatitis C Virus Epidemic Control Using a Nonlinear Adaptive Strategy

5

FIG. 2 Conceptual diagram of the nonlinear adaptive method developed to control the HCV epidemic in the

existence of uncertainties on the model parameters.

In the following section, taking advantage of the Lyapunov stability theorem, it will be proven that the control laws (11), (12) together with some adaptation laws guarantee the tracking convergence, stability, and robustness for the treatment of HCV outbreak.

4.2. Stability Proof and Adaptation Laws The closed-loop dynamics of the system is achieved first by substituting the control laws (11), (12) into the dynamics of the HCV epidemic (1): _



 S u + λ1 S u ¼ Y1 θ 1 Su



  _  C + λ2 C ¼ Y2 θ 2 C

(13)

(15) and (16), the tracking convergence, stability and robustness for the aware susceptible and chronically infected classes will be proven. With this aim, a positive definite Lyapunov candidate function is selected as T   1   T 1 V ¼ ½S 2u + C 2 + θ 1 Γ 1 1 θ 1 + θ 2 Γ2 θ 2  2

The Lyapunov function’s time derivative is determined: T T  _     _ ^_ 1 V ¼ S u S u + C C + θ^_ 1 Γ 1 1 θ 1 + θ 2 Γ2 θ 2 

(15)

T  θ^_ 2 ¼ C C Y2 Γ 2

(16)

where Γ 1 and Γ 2 are the adaptation gain matrices and considered to be positive definite. Now, employing the Lyapunov stability theorem [27] and based on the previously derived closed-loop dynamics (13), (14) and the designed adaptation laws

2



V ¼ λ1 S u  λ2 C 

T  θ^_ 1 ¼ Su S u Y1 Γ 1

(18)

_ It should be mentioned that θ ¼ θ^_ because θ is constant (θ_ is zero). By substituting the adaptation laws (15) and (16) into Eq. (18), the time derivative of V is simplified to

(14)

where θ i (for i ¼ 1, 2) is defined as θ^i  θi . The adaptation laws are designed to update parameters’ estimation to keep the system’s robustness against uncertainties, as

(17)

2

(19)

As mentioned in the previous descriptions, λ1 and λ2 are considered to be positive; thus, the Lyapunov function’s time derivative is negative semidefinite. Thus, based on Barbalat’s lemma (described in Appendix) and the Lya punov stability theorem [27], it is proven that S and u  C converge to the zero. In other words, employing the suggested adaptive feedback control strategy ensures  the tracking convergence and stability (S ! 0 and u  C ! 0 as t !∞) in the presence of uncertainties. Thus, the numbers of unaware susceptible (Su) and chronically infected (C) people converge to the desired values (Su ! Sud and C ! Cd).

6

Modeling and Control of Drug Delivery Systems

For the effectiveness evaluation of the proposed method, some simulations are presented in this section. Note that computer simulations have proven to be useful for evaluating the spreading behavior of infectious diseases [28]. In the present study, the simulation process is performed in the Simulink-Matlab environment. The parameters’ values of the HCV epidemic model (1) are listed in Table 2. A small society with a total population of 1310 people at the beginning of the investigation is used. The following desired scenarios are considered for the reduction of unaware susceptible individuals (Sud ) and the treatment of chronically infected people (Cd): Sud ¼ ðSu0  Suf Þexpða1 tÞ + Suf

(20)

Cd ¼ ðC0  Cf Þexpða2 tÞ + Cf

(21)

where a1 and a2 are the desired population reduction rates. Su0 and Suf are the initial and final (steady-state) populations of the unaware susceptible class, respectively. C0 and Cf are the initial and final (steady-state) populations of the chronically infected compartment, respectively. The presented reduction and treatment scenarios (20), (21) are employed in these simulations as the desired decreasing behavior of the HCV epidemic control. However, other continuously decreasing fashion can be used as desired scenarios without loss of generality. The values of parameters in the desired HCV population reduction scenarios (20), (21) are listed in Table 3. These scenarios for unaware susceptible and chronically infected compartments are shown in Fig. 3.

TABLE 3

Values of Parameters in the Desired HCV Population Reduction Scenarios (20), (21). Parameter

Value

Su0

1000

C0

100

Suf

0

Cf

0

a1

0.4

a2

0.2

1000 Sud : Desired unaware susceptible population

900

C d : Desired chronically infected population

800

Desired populations

5. RESULTS AND DISCUSSION

700 600 500 400 300 200 100 0 0

5

10

15

20

Time (year) FIG. 3 Desired scenarios for the reduction of unaware

TABLE 2

Values of the HCV Parameters in Its Mathematical Model (1) [4]. Parameter

Value

b

0.012

μ

0.006

β

0.15

K1

0.5

K2

0.2

α

0.1

γ

4

q

0.2

ξ

0.8

p

0.5

θ

0.001

susceptible and chronically infected compartments in the HCV epidemic.

In the absence of control inputs, the HCV infection will extend in the society based on Eq. (1). Accordingly, the treated population will decrease and reach zero exponentially due to the lack of treatment process. In that case (no control input), unaware and aware susceptible individuals will get the hepatitis C virus in contact with the infected people in I and C compartments and will join the acutely infected class (I). Since there is no treatment for acutely infected individuals (as seen in Eq. 1), the HCV disease will progress and reach the chronic stage. Thus, the population of the chronically infected compartment (C) will increase and the populations of all other compartments will decrease. Fig. 4 depicts the above-mentioned points about the HCV outbreak in the case of no control input.

CHAPTER 1

Hepatitis C Virus Epidemic Control Using a Nonlinear Adaptive Strategy

1100

1400

1000

Su : Unaware susceptible population

1200

900

Population of compartments

Population of compartments

7

Su : Unaware susceptible population Sa: Aware susceptible population

800 700 600 500 400

Sa: Aware susceptible population

1000 800 600 400 200

300 200

0 0

5

10

(A)

15

20

0

5

10

(A)

Time (year)

15

20

Time (year) 100

140

I: Acutely infected population C: Chronically infected population T: Treated population

90

Population of compartments

Population of compartments

I: Acutely infected population C: Chronically infected population T: Treated population

120 100 80 60 40

80 70 60 50 40 30 20 10

20

0 0

0 0

(B)

5

10

15

20

Time (year)

FIG. 4 Populations of (A) unaware and aware susceptible,

(B)

5

10

15

20

Time (year)

FIG. 5 Populations of (A) unaware and aware susceptible,

and (B) acutely infected, chronically infected and treated classes in the absence of control inputs.

and (B) acutely infected, chronically infected and treated classes in the presence of control inputs (u1 and u2) based on the proposed laws (11), (12).

However, applying the proposed strategy based on the designed nonlinear control laws (11), (12) with the obtained adaptation laws (15), (16), the population changes in different compartments in the presence of 20% parametric uncertainty are depicted in Fig. 5. As seen, due to the employment of the first control input (u1), the population of unaware susceptible compartment (Su) reduces and they join the aware susceptible class (Sa). Since the rate of infection for the aware susceptible people is less than that of the unaware ones, and due to the effect of control input u1, the extension of the HCV infection decreases compared with the no-control-input case (shown in Fig. 4). Moreover, using the treatment as the second control input (u2), the population of the chronically infected compartment (C)

decreases (Fig. 5) based on the described scenarios (Cd in Fig. 3). Thus, the populations of unaware susceptible and chronically infected classes reduce and the population of aware susceptible increases in Fig. 5, which are in accordance with the HCV dynamics (1). Although 20% parametric uncertainty is taken into account for the nonlinear model, simulation results show that the proposed control strategy satisfied its objective, which is convergence to desired population reduction and treatment scenarios (Su ! Sud and C ! Cd). Fig. 6 depicts the desired and real populations of unaware susceptible and chronically infected classes, which imply the appropriate convergence performance using the nonlinear controller. The corresponding tracking errors are presented in Fig. 7.

8

Modeling and Control of Drug Delivery Systems 1000

1

900 800 700

Sud : Desired unaware susceptible population

0.9

C: Actual chronically infected population C d : Desired chronically infected population

0.8

Control inputs

Population of compartments

u1: Effort rate to inform the susceptible individuals

Su : Actual unaware susceptible population

600 500 400 300

0.7 0.6 0.5

200

0.4

100 0 0

5

10

15

20

Time (year)

0.3 0

5

10

15

20

Time (year)

FIG. 6 Convergence of unaware susceptible and chronically infected populations (Su and C) to their desired values (Sud and Cd).

Su − Sud

0.2

FIG. 8 Control inputs (u1 and u2) during the treatment period of HCV epidemic.

desired scenarios (20), (21) for the reduction of unaware susceptible and chronically infected compartments comply with the control input limitations. Fig. 9 illustrates the tuning of estimated parameters (θ^1 and θ^2 ) based on the designed adaptation laws (15), (16) in the presence of 20% uncertainty.

0.3

Population tracking errors

u2: Treatment rate for chronically infected individuals

C − Cd

0.1 0

5.1. System Response to Different Uncertainty Levels

−0.1 −0.2 −0.3 −0.4 −0.5 −0.6 0

0.5

1

1.5

2

Time (year) FIG. 7 Tracking errors between the desired and real values

of unaware susceptible and acutely infected compartments.

As described, two control inputs are adjusted according to the proposed nonlinear adaptive strategy in order to prevent the HCV outbreak. The first control input u1(t) denotes the effort rate to inform the susceptible individuals from the HCV and the second one u2(t) is the treatment rate for chronically infected individuals. These control inputs are considered to be normalized in Eq. (1) to be in the range of [0, 1]. The obtained values for these inputs using the proposed control strategy are shown in Fig. 8, which satisfy the physiological constraint (u1  [0, 1]). This implies that the considered

In this section, the effects of different uncertainty levels are investigated for the HCV epidemic dynamics. For this purpose, 50%, 70%, and 90% uncertainties are considered on the initial guess of parameters in θ1 and θ2 (defined in Eqs. 7, 8). Performance of the adaptation laws (15), (16) on the tuning of estimated model parameters is investigated in Fig. 10. As discussed and proven in Section 4, these adaptation laws guarantee that the estimation errors of the HCV dynamic parameters remain bounded against different uncertainty levels. Fig. 11 shows the population errors of unaware susceptible and chronically infected  classes in tracking their  desired value (S u ¼ Su  Sud and C ¼ C  Cd ). As observed in Fig. 11, the increment of parametric  uncertainties increases the magnitude of errors (Su and  C ) and their initial variations. However, after a period of time (about 0.2 year), the error magnitudes have reached zero, which means that the tracking convergence has been achieved for different values of uncertainties. In other words, the population of unaware susceptible and chronically infected compartments converged to their desired values (Su ! Sud and C ! Cd) in the existence of different levels of uncertainty.

CHAPTER 1

Hepatitis C Virus Epidemic Control Using a Nonlinear Adaptive Strategy 0.05

0.2 Estimation of Estimation of

0.16

Estimation of

0.14

(2) 1 1

Estimation of model parameters

0.18

Estimated parameters

9

(4)

(6) 1

0.12 0.1 0.08 0.06 0.04

0.04

Estimation of

1

(1) for 50% uncertainty

Estimation of

1

(1) for 70% uncertainty

Estimation of

1

(1) for 90% uncertainty

Actual value of

1

(1)

0.03

0.02

0.01

0

0.02 0 0

(A)

0.5

1

2

Time (year)

−0.01 0

(A)

1.6

0.18

1.5

Estimation of Estimation of

0.14

2

(1)

Estimation of model parameters

Estimation of

(2) 2 2

(3)

0.12 0.1 0.08 0.06 0.04

5

10

15

20

Time (year)

0.2

0.16

Estimated parameters

1.5

1.4

Estimation of

2

(1) for 50% uncertainty

Estimation of

2

(1) for 70% uncertainty

Estimation of

2

(1) for 90% uncertainty

Actual value of

2

(1)

1.3 1.2 1.1 1 0.9 0.8

0.02

0.7

0 0

0.5

(B)

1

1.5

2

Time (year)

FIG. 9 Estimation of parameters in (A) θ1 and (B) θ2 during the

treatment period based on the adaptation laws (15) and (16), respectively.

6. CONCLUSION In the present study, a new nonlinear adaptive strategy was developed to control the hepatitis C virus epidemic based on a mathematical model having uncertainties. For the first time, an adaptive feedback controller was employed to decrease the populations of unaware susceptible and chronically infected compartments based on the desired scenarios. Two control inputs were employed for this goal. The first one is the effort rate to inform the susceptible individuals from the HCV and the second one is the rate of treatment for chronically infected people. The Lyapunov stability theorem and Barbalat’s lemma were used to prove the tracking convergence to desired treatment scenarios. The

(B)

0

5

10

15

20

Time (year)

FIG. 10 Adaptation of (A) θ1(1) and (B) θ2(1) using Eqs. (15) and (16), respectively, for different uncertainty levels.

proposed control laws and adaptation laws provided the stability of the closed-loop HCV epidemic system in the presence of parametric uncertainties. Results of numerical simulations showed that by adjusting the control inputs and the estimated parameters based on this strategy, the number of the unaware susceptible and chronically infected individuals are decreased. As a result, the population of the aware susceptible was increased and the population of the acutely infected and treated classes reached out to zero at the end of the process. Moreover, the obtained results implied that the tracking convergence is achieved for a wide range of uncertainties. Designing optimal trajectories and employing unstructured uncertainties can be considered as the next steps of this research in the future.

10

Modeling and Control of Drug Delivery Systems By integrating both sides of Eq. (A.2), one can write:

1.5

Tracking error for unaware susceptible population

1

−0.5

(A)

−2.5 0

Z

(Su − Sud ) for 50% of uncertainty (Su − Sud ) for 70% of uncertainty

Vð0Þ  Vð∞Þ ¼ lim

t!∞ 0

(Su −Sud ) for 90% of uncertainty

0.5

(A.3)

Since V is negative, V (0) is larger than V (∞) and V (0)  V (∞)  0. Moreover, as mentioned previously, V is bounded based R t on the Lyapunov stability theorem. Thus, lim t!∞ 0 gðηÞdη in Eq. (A.3) exists and has a bounded value. Therefore, it is concluded using the Barbalat’s lemma that

−1 −1.5

2



lim ðλ1 S + λ2 C 2 Þ ¼ 0

t!∞

−2

0.1

0.2

0.3

0.4

(C − Cd ) for 50% of uncertainty (C − Cd ) for 70% of uncertainty (C − Cd ) for 90% of uncertainty

0

−0.2 −0.4 −0.6

0.1

0.2

(B)

0.3

0.4

0.5

Time (year)

FIG. 11 The difference between (A) unaware susceptible population and its desired value (S u ¼ Su  Sud ) and (B) chronically infected population and its desired value  (C ¼ C  Cd ) for different parametric uncertainty levels.

APPENDIX. BARBALAT’S LEMMA   



The Lyapunov function V(S u , C , θ 1 , θ 2 ) in Eq. (17) is   positive definite and its time derivative (V (S u , C )) in Eq. (19) is negative semidefinite. Thus, based on the Lyapunov stability theorem [27], V is bounded and it   is concluded that S u , C , θ 1 , and θ 2 remain bounded. Barbalat’s lemma: R t If g is a uniformly continues function and lim t!∞ 0 gðηÞdη exists and has a finite value, it is guaranteed that [27] 

lim gðtÞ ¼ 0

(A.1)

t!∞

In order to use this lemma for the controlled system of HCV outbreak, g(t) is considered to be  V : 

2



gðtÞ ¼  V ¼ λ1 S u + λ2 C 

(A.4)

REFERENCES

0.4

0.2

u

0.5

Time (year)

Estimation error for chronically infected population

gðηÞdη



0

−0.8 0

t

2

(A.2)

[1] World Health Organization (WHO), Hepatitis C. (2017) http://www.who.int/mediacentre/factsheets/fs164/en/. [2] D. Prati, Transmission of hepatitis C virus by blood transfusions and other medical procedures: a global review, J. Hepatol. 45 (2006) 607–616. [3] A. Wasley, M.J. Alter, Epidemiology of hepatitis C: geographic differences and temporal trends, Semin. Liver Dis. 20 (1) (2000) 1–16. [4] S. Zhang, X. Xu, Dynamic analysis and optimal control for a model of hepatitis C with treatment. Commun. Nonlinear Sci. Numer. Simul. 46 (2017) 14–25, https:// doi.org/10.1016/j.cnsns.2016.10.017. [5] S. Zhang, Y. Zhou, Dynamics and application of an epidemiological model for hepatitis C. Math. Comput. Model. 56 (2012) 36–42, https://doi.org/10.1016/j. mcm.2011.11.0817. [6] S.L. Chen, T.R. Morgan, The natural history of hepatitis C virus (HCV) infection, Int. J. Med. Sci. 3 (2006) 47–52. [7] A.M. Bisceglie, Natural history of hepatitis C: its impact on clinical management, Hepatol. Millenn. 31 (4) (2000) 1014–1018. [8] J. Yuan, Z. Yang, Global dynamics of an SEI model with acute and chronic stages, J. Comput. Appl. Math. 213 (2008) 465–476. [9] R.M. Klevens, D.J. Hu, R. Jiles, S.D. Holemberg, Evolving epidemiology of hepatitis C virus in the United States, Clin. Infect. Dis. 55 (suppl. 1) (2012) S3–S9. [10] S. Banerjee, R. Keval, S. Gakkhar, Modeling the dynamics of hepatitis C virus with combined antiviral drug therapy: interferon and ribavirin, Math. Biosci. 245 (2013) 235–248. [11] K.O. Okosun, Impact and optimal control of movement on a multipatch hepatitis C virus model, J. Pure Appl. Math. 5 (1) (2014) 80–95. [12] K.O. Okosun, O.D. Makinde, Optimal control analysis of hepatitis C virus with acute and chronic stages in the presence of treatment and infected immigrants, Int. J. Biomath. 7 (2) (2014) 1–23. [13] J. Khodaei-Mehr, S. Tangestanizadeh, R. Vatankhah, M. Sharifi, Optimal neuro-fuzzy control of hepatitis C virus integrated by genetic algorithm, IET Syst. Biol. 12 (2018) 154–161.

CHAPTER 1

Hepatitis C Virus Epidemic Control Using a Nonlinear Adaptive Strategy

[14] J. Khodaei-Mehr, S. Tangestanizadeh, R. Vatankhah, M. Sharifi, ANFIS-based optimal control of hepatitis C virus epidemic, IFAC-PapersOnLine 51 (15) (2018) 539–544. [15] J. Martcheva, C. Castillo-Chavez, Diseases with chronic stage in a population with varying size, Math. Biosci. 182 (2005) 1–25. [16] X. Hu, F. Sun, Threshold dynamics for an epidemic model with acute and chronic stages, Int. J. Nonlinear Sci. 12 (1) (2011) 105–111. [17] N. Ainea, E.S. Massawe, O.D. Makinde, L. Namkinga, Mathematical analysis of control strategies of HCV in a community with inflow of infected immigrants, Adv. Infect. Dis. 5 (2015) 1–13. [18] M. Shen, Y. Xiao, W. Zhou, Z. Li, Global dynamics and applications of an epidemiological model for hepatitis C virus transmission in China, Discrete Dyn. Nat. Soc. 2015 (2015) 1–13. [19] R. Shi, Y. Cui, Global analysis of a mathematical model for hepatitis C virus transmissions, Virus Res. 217 (2016) 8–17. [20] H. Moradi, M. Sharifi, G. Vossoughi, Adaptive robust control of cancer chemotherapy in the presence of parametric uncertainties: a comparison between three hypotheses, Comput. Biol. Med. 56 (2015) 145–157. [21] M. Sharifi, H. Moradi, Nonlinear composite adaptive control of cancer chemotherapy with online identification of uncertain parameters, Biomed. Signal Process. Control 49 (2019) 360–374.

11

[22] D. Boiroux, A.K. Duun-Henriksen, S. Schmidt, K. Norgaard, N.K. Poulsen, H. Madsen, J.P. Jorgensen, Adaptive control in an artificial pancreas for people with type 1 diabetes, Control Eng. Pract. 58 (2017) 332–342. [23] O. Aghajanzadeh, M. Sharifi, S. Tashakori, H. Zohoor, Nonlinear adaptive control method for treatment of uncertain hepatitis B virus infection, Biomed. Signal Process. Control 38 (2017) 174–181. [24] M. Sharifi, H. Moradi, Nonlinear robust adaptive sliding mode control of influenza epidemic in the presence of uncertainty, J. Process Control 56 (2017) 48–57. [25] R. Padmanabhan, N. Meskin, W.M. Haddad, Optimal adaptive control of drug dosing using integral reinforcement learning, Math. Biosci. 309 (2019) 131–142. [26] W.J.M. Probert, K. Shea, C.J. Fonnesbeck, M.C. Runge, T. E. Carpenter, S. D€ urr, M.G. Garner, N. Harvey, M. A. Stevenson, C.T. Webb, M. Werkman, M. J. Tildesley, M.J. Ferrari, Decision-making for foot-andmouth disease control: objectives matter, Epidemics 15 (2016) 10–19. [27] J.E. Slotine, Applied Nonlinear Control, Prentice Hall, Upper Saddle River, NJ, 1991. [28] C. Orbann, L. Sattenspiel, E. Miller, J. Dimka, Defining epidemics in computer simulation models: how do definitions influence conclusions? Epidemics 19 (2017) 24–32.

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CHAPTER 2

Integral Sliding Mode Control of Immune Response for Kidney Transplantation POURIA FARIDIa,c • RAMIN VATANKHAHa • MOJTABA SHARIFIb,c a

Department of Mechanical Engineering, Shiraz University, Shiraz, Iran, bDepartment of Electrical and Computer Engineering, University of Alberta, Edmonton, AB, Canada, cDepartment of Medicine and Dentistry, University of Alberta, Edmonton, AB, Canada

1. INTRODUCTION Acute kidney injury (AKI) is a prevalent secondary disease for patients who are in the treatment process inside hospitals. Even those who passed the recovery period of AKI are still at the risk of other complications, such as chronic kidney diseases (CKD) [1]. According to the annual report of the United States Renal Data System in 2018, people with estimated Glomerular Filtration Rates (eGFR) less than 60 mL/min/1.73 m2 are at the risk of kidney diseases. CKD has five stages characterized by eGFR magnitude: stage 1 with eGFR> 90 mL/min/ 1.73 m2 is the least dangerous stage, while people in stage 5 with eGFR< 15 mL/min/1.73 m2 are suffering from the end-stage renal disease (ESRD) that is mortal in the absence of dialysis or renal transplantation [2, 3]. eGFR that shows kidney performance is calculated by the CKD-EPI creatinine equation, based on the serum creatinine concentration in blood (a breakdown product resulting from the activity of muscles), race, sex and age [4]. In developing countries such as India and Pakistan, which together are home to one-sixth of the world’s human population, the annual incidence of ESRD is estimated at around 100 per million people. This means 100,000 and 15,000 patients are at the risk of ESRD in 1-billion and 150-million populations of India and Pakistan, respectively. When dialysis equipment is not satisfactory, kidney transplantation is the best way of treatment [5]. In developed countries such as the United States, the number of people with kidney failure (in the ESRD stage) is continually growing and has the highest increase rate in the world. Research studies have shown that 75% of children in the United States with ESRD got renal transplants, which implies the significance of this treatment [2]. In renal transplantation, there is a possibility of kidney rejection in the absence of standard health care. Pharmacological immuno-suppression is currently the most effective way to reduce the chance

of this rejection. However, using this therapy, patients’ bodies will be exposed to several viral loads and bacterial pathogens [6]. Human herpesvirus five or Human Cytomegalovirus (HCMV) is the most common and significant pathogen among renal transplant recipients. Primary HCMV infection is usually without any specific symptom, but the uncontrolled case will lead to a life-long infection in patients’ bodies. Moreover, HCMV disease can be caused by the reactivation of HCMV latent infectious viral load [7]. In recent years, a wide range of investigations and experiments have been conducted regarding the HCMV infection, its dynamics, diagnosis, risk factors and international guidelines on its management [8–12]. Emery et al. [8] opposed the popular belief that Cytomegalovirus (CMV) replicates slowly, which had been an accepted theory due to time-consuming in vitro experiments for showing up the cytopathic effects. Taking three different cases into account, they proved that CMV in vivo replication has tremendous dynamics (variation). Serological tests, standard tube cell culture technique, antigenemia assay, polymerase chain reaction (PCR), immunohistochemistry, nucleic acid sequencebased amplification (NASBA) and hybrid capture assay are recognized methods of CMV detection that each has some merits and demerits [9]. Bataille et al. [11] investigated modern immunosuppression and its risk factors on 300 people having the same trial therapy, and the case D+/R
(CMV-seropositive donor/CMVseronegative recipient) had the highest risk factor. Also, based on their observations, it was mentioned that a patient with impaired early kidney function becomes a candidate at risk. Mathematical modeling or analysis is a vital tool for investigating the dynamic behaviour of biological systems and making decisions about treatment methods

Modeling and Control of Drug Delivery Systems. https://doi.org/10.1016/B978-0-12-821185-4.00010-5 © 2021 Elsevier Inc. All rights reserved.

13

14

Modeling and Control of Drug Delivery Systems

and their duration. There has been considerable effort in employing various useful control strategies for a variety of diseases and illnesses. Some control strategies have been developed for the hepatitis B virus (HBV). For instance, Sheikhan et al. [13] compared three strategies (nonlinear feedback neural-type sigmoid, open-loop time-based fuzzy and closed-loop fuzzy) to achieve the optimum performance of the HBV control. Laarabi et al. [14] proved the existence of an optimal control method by combining two control laws to reduce the therapy cost and maximize the number of healthy cells in HBV infection. In another work [15], a fuzzy logic structure was used to solve an HBV optimal control problem. Two other control inputs were designed to have efficient drug therapy considering both hindering viral production and inhibiting new infection of HBV based on a robust adaptive Lyapunov-based control theory [16]. Hepatitis C virus (HCV) has also attracted the attention of researchers in recent years. Zhang et al. [17] developed an epidemiological model of HCV and used numerical simulations to study the influence of the model parameters based on available data obtained from China. In another study [18], the performance of an optimal controller was investigated based on an HCV model having acute-infected and chronic-infected individuals as its compartments. After that, Zhang et al. [19] utilized an optimal control measure to inhibit the prevalence of HCV while minimizing the cost and population of infected individuals. A novel optimal adaptive neuro-fuzzy controller was developed to decrease the number of HCV infected individuals using an additional genetic algorithm optimization [20]. Cancer tumor modeling and control have also been studied in this field of research. In 2001, a four-population model containing tumor and host cells, drug therapy and immune response was presented [21]. Accordingly, Babaei et al. [22] proposed a model reference adaptive control method to determine a personalized drug administration to treat cancer with parametric uncertainty. Moradi et al. [23] also developed an adaptive robust control method for three nonlinear mathematical cell-kill models of cancer in the presence of parameter uncertainties. They have extended the previous study [23] by enhancing their strategy to a modern composite adaptive control in which the model parameters were precisely identified online during the cancer chemotherapy [24]. Moreover, Khalili et al. [25] suggested an optimal open-loop control strategy for drug delivery in chemotherapy considering the human obesity effects. Moradi and Sharifi [26] also proposed a nonlinear robust adaptive sliding mode control method to reduce the number of susceptible and infected humans to zero

in an influenza outbreak regarding a five-state compartmental model of this disease. The employed model of influenza was developed by Arino et al. [27] with five state variables known as SEAIR (Susceptible-ExposedAsymptomatic-Infectious-Recovered). That model was enhanced in Ref. [28] by defining the vaccination, social distancing and antiviral rates as three possible control inputs. In this modern era, HIV prevention and treatment is also of pivotal importance. Ngina et al. [29] presented an in vivo deterministic model of HIV and presented an optimal control scheme based on that. In a new study [30], a robust sliding mode controller was formulated to reduce the population of infected CD4+ T cells with antiviral therapy according to the acquired output information. In 2000, a differential SEIR model of malaria, containing both humans’ and mosquitos’ populations and their interaction was taken into account [31]. Another epidemiological model of malaria was formulated to consider personal protection, possible treatment and vaccination strategies in two latent periods [32]. After that, Rafikov et al. [33] employed an optimal control method for a mathematical malaria model by placing genetically modified mosquitos in the environment. Furthermore, a robust nonlinear controller with adaptive gains was proposed to inhibit the prevalence of malaria with seven variables for human and mesquite compartments [34]. To study the immune response of renal transplant recipients, there has been made considerable effort to predict, model and optimally control the HCMV infection in both primary and latent cases. Flechet et al. [35] analyzed AKI-predictor as an online machine learningbased prediction tool and compared it to physicians’ ability to predict AKI in clinical uses. Based on the obtained results, it was emphasized that AKI-predictor was beneficial in terms of successfully removing false HCMV positives (error in determining a patient at high risk) and reducing clinical costs. In addition, Parreco et al. [36] compared different machine learning algorithms for predicting AKI. Wodarz et al. [37] presented a model based on mice infected with murine cytomegalovirus (MCVM), which introduced important knowledge for explaining the growth of CMV specific CD8+ T cells in human by getting older. Kepler et al. [38] formulated a fifth-order state-space model for HCMV infection that was extended by Banks et al. [39] to six-order one, considering serum creatinine of the blood as a measure of the kidney performance. Their model contained antiviral and immunosuppressive drugs as two control inputs. Then, Kwon et al. [40] simplified the model in [39], considering immunosuppressive drug as the only control input, proved the local stability of the dynamics

CHAPTER 2 Integral Sliding Mode Control of Immune Response for Kidney Transplantation and determined both uninfected and infected steadystates. They utilized a model predictive control (MPC) method as an optimal strategy to achieve a balance between over-suppression and under-suppression. In this research work, we propose a nonlinear integral sliding-mode control scheme for the first time based on the model developed in [40] to adjust the viral load and serum creatinine of the blood to satisfy standard clinical limitations. Using the proposed controller, the nonlinear behavior of this infection is taken into account in defining the control law for the immunosuppressive drug without any linearization of the model. The proposed method facilitates tracking of any desired concentration of CD8+ T cells as an immune response that specifically targets the kidney in order to maintain the serum creatinine concentration below an acceptable clinical limit during the treatment period. The performance of closed-loop dynamics in the presence of different parameter values is evaluated and the robustness of the controller is shown for various cases via comprehensive sensitivity analysis. In Section 2, the HCMV dynamic model is presented and relationships between its state variables and parameters are explained. The controller design and Lyapunov stability proof are formulated and described in Section 3. Simulations and numerical results for primary and reactivation cases, required control inputs, sensitivity analysis of the model parameters, and discussion on the antiviral therapy are mentioned in Section 4. Concluding remarks are finally explained in Section 5.

2. MATHEMATICAL MODEL OF TRANSPLANT RECIPIENTS The dynamic model described in [40] is used here to analyze the HCMV infection in transplant recipients:
 S S_ ¼ λs 1
S
βSV κs

(1)

I_ ¼ βSV
δI I
mEv I

(2)



V ¼ ρV δI I
δV V
βSV   ρ V E_v ¼ ð1
εÞ λEV + Ev Ev
δEv Ev V + κV

(3) (4)

E_K ¼ ð1
εÞλEK
δEK EK

(5)

δC κ EK  C, C ¼ λC
EK + κ EK

(6)

where S(t) is the concentration of susceptible cells (Cells/μL-blood), I(t) is the concentration of infected cells (Cells/μL-blood), V(t) is the concentration of free

15

HCMV (Copies/μL-blood), Ev(t) is the concentration of HCMV-specific CD8+ T cells (Cells/μL-blood), EK(t) is the concentration of allospecific effector CD8+ T cells that specifically target kidney (Cells/μL-blood), and C(t) is the serum creatinine concentration (mg/ dL). It is clear that Eqs. (1)–(4), presenting the response of the immune system to the viral load, are coupled with Eqs. (5), (6), presenting the response of the immune system to the newly transplanted kidney [39]. The term βSV in Eqs. (1), (3) stands for the loss rate of both susceptible cells and viral load, in which they will be transformed into the infected ones in Eq. (2). If there is no virus, susceptible cells grow normally with the rate of λs(1
S/κ s)S. The terms δII and δVV represent the natural death rate of infected cells and viruses, respectively. Moreover, infected cells’ population decreases due to the immune response (Ev) against HCMV with the term mEvI in Eq. (2). Also, the viral load increases with the death of infected cells through the term ρVδII in Eq. (3). ε is the normalized control input acting as the immunosuppressive drug dosage per day, which reduces the immune system response affecting Eqs. (4), (5) in order to maintain the creatinine concentration (C) below its threshold for ideal kidney performance. Immune response (Ev) against HCMV increases in two ways; naturally with the rate λEV and also by acting against the viral load with the term (ρEvV/V + κ V)Ev, while its natural death rate is δEvEv. Allospecific immune effector cells (EK) have natural birth and death rates modeled with the terms λEK (related to the human leukocyte antigen system) and δEKEK, respectively. The serum creatinine level of the blood (C) is a measure for kidney performance. λC and δCκ EKC/(EK + κ EK) represent its production rate and death rate (in terms of the allospecific CD8+ T cells’ activity, EK), respectively. Note that this model does not consider the HCMV-specific CD4+ T cells for simplicity and the total period of treatment is considered 450 days [40]. The numerical values of the mentioned parameters are extracted from [40] and described in Table 1.

3. CONTROL SCHEME In this section, the proposed sliding mode controller (SMC) is described to obtain a balance between normal kidney performance and proper immune system defence against the viral load. To achieve this goal, two constraints are defined on the state variables V(t) (viral load) and C(t) (serum creatinine of blood) that should be less than their defined thresholds to ensure that 1. viruses will not make dangerous situation and 2. the transplanted kidney is working properly.

16

Modeling and Control of Drug Delivery Systems

TABLE 1

TABLE 2

Parameters Values of the HCMV Dynamic Model for Kidney Transplantation [40]. Parameters

Magnitudes of Desired Parameters in Eqs. (8) and (9).

Value
3

Parameters

Value

λs

10

per day

EKf

1.5

κs

400 cells/μL-blood

Cf

1.1875

a

0.12


4

β

2 10

δI

0.2 per day

μL-blood/(copies day)

m

0.048571 μL-blood/(cells day)

ρV

9.6 copies/cells

δV

0.2 per day

λEV

0.5 cells/(μL-blood day)

ρEv

1 per day

κV

20 copies/μL-blood

δEv

0.05 per day

λEK

0.5 cells/(μL-blood day)

δEK

0.1 per day

λC

0.2 mg/(dL day)

κ EK

8 cells/μL-blood

δC

0.2 per day

Accordingly, the final magnitude of EK is obtained based on Eq. (6) as 0 ¼ λC


3.2. Lyapunov Stability Proof Now, we consider a quadratic Lyapunov function candidate V in terms of distance from the sliding surface as follows: V¼

A proper SMC firstly needs a definition for the sliding surface, which its stability could be proven by the Lyapunov theorem. In this work, an integral sliding surface is formulated in terms of tracking error as follows: 

ðt



e dt,



(8)

such that the constraint for being below the threshold value of 1.2 for C(t) [40] is satisfied. Therefore, we consider Cf ¼ 1.1875 as the final steady-state value to make sure that it will be less than the mentioned threshold.

(10)

  V ¼ s_s ¼ sðe_ + eÞ ¼ s E_ K
E_ Kd + EK
EK d

(7)

which e ¼ EK
EK d is the tracking error with respect to the desired concentration of allospecific effector CD8+ T cells (EKd). The main characteristic of this sliding surface is having a stable behaviour for the error dynamics. Therefore, the control law is developed such that the system response tends to this stable surface. Due to the existence of one control input as the immunosuppressive drug dosage, only one desired trajectory for EK is defined based on [26]:

1 2 s , 2

in which s is the sliding surface defined in (7). The Lyapunov candidate (10) is positive definite, and it is needed to be proven that its time derivative is negative definite: 

0

  EK d ¼ EK 0
EK f exp ð
at Þ + EK f

(9)

where EK0 and EKf are the initial and final (steady-state) values of allospecific CD8+ T cells, respectively. The magnitudes of these desired values are presented in Table 2, where Cf is the final value of the serum creatinine in blood and a is the reduction rate of EK in Eq. (8) during the treatment period.

3.1. Design of an Integral Sliding Mode Controller

s ¼e +

δC κ EK  1:1875 ! EK f ¼ 1:5 EK f + κ EK

(11)

For this purpose, we design the control input ε such that V 
η s tanh ðsÞ, which η is a positive constant. As a result, we first obtain εeq that makes s_ ¼ 0 based on Eqs. (5), (8), and (11): 

εeq ¼

  λEK
δEK EK + a EK 0
EK f exp ð
at Þ + EK
EK 0
EK f exp ð
at Þ
EK f λEK (12)

This control signal will lead the system to reach the desired trajectory. However, it works when the system is on the sliding surface, in the absence of any disturbances. Thus, to obtain the desired value for V to guarantee the stability, the following controller is proposed: 

ε ¼ εeq +

k tanh ðsÞ, λEK

(13)

CHAPTER 2 Integral Sliding Mode Control of Immune Response for Kidney Transplantation in which k is a positive constant. The continuous function tanh(s) is used instead of sgn(s) to avoid or reduce chattering phenomenon as an undesirable phenomenon. Employing the controller (13), V is obtained as 



V ¼ s_s ¼
k s tanh ðsÞ

(14)

It is chosen k  η to satisfy the sliding mode condition for V : 



V 
ηstanh ðsÞ

(15)

As mentioned before, η is positive. Moreover, s tanh (s) is a positive function. Consequently, V is negative definite and the system is asymptotically stable.

17

of allospecific CD8+ T cells targeted kidney (EK) is followed appropriately. In comparison with the primary case (Fig. 1), there is a little more concentration of the susceptible cells, infected cells, viral loads, HCMVspecific CD8+ T cells and serum creatinine of the blood in reactivation case (Fig. 2). However, in both cases, the desired trajectory for EK is tracked appropriately as illustrated in Fig. 3. Fig. 3 depicts the error signals in primary infection and the reactivation cases. In both cases, a small error (with the order of 10
3) is seen in the first days that tends to zero eventually.



4. SIMULATION RESULTS 4.1. Primary Infection Case In healthy individuals without a history of HCMV infection, primary infection can occur due to the immunosuppression drug. Initial values for the primary infection case are as follows [40]:

 ½S0 , I0 , V0 , EV 0 , EK 0 , C0 ¼ 400, 10
12 , 10
2 , 10
12 , 10
12 , 1 (16)

Fig. 1 shows the simulation results of the primary infection case for six variables with the immunosuppression drug using the proposed controller. It can be seen that the controller satisfies the constraints for viral load (V  3.5) and the serum creatinine (C  1.2). The viral load constraint is satisfied and only violated one time due to the over-suppression case; however, the desired value of EK is followed appropriately. A high correlation can be seen between the desired and actual trajectories, which indicates that the control performance is suitable.

4.2. Reactivation Case Latent infection is the next stage after primary infection, in which the virus remains in cells and the HCMV disease can occur [41]. The initial conditions are obtained from [39] that obtained as the equilibrium point of the untreated system. Fig. 2 demonstrates the results for the reactivation case with the following initial conditions:

½S0 , I0 , V0 , EV 0 , EK 0 , C0 ¼ 399:999, 8:56  10
7 , 7:34  10
6 , 10, 5, 1:625 (17)

Simulation results show that in reactivation case, the serum creatinine will also eventually come below its limit (after 40 days), the viral load is less than its threshold except for a period that needs an antiviral drug (197th day to 225th day), and the desired concentration

4.3. Control Signals Fig. 4 shows the normalized immunosuppressive drug dosage as the control input, for the primary and reactivation cases. It is observed that 70% of the immunosuppressive drug dosage is enough to satisfy the constraints below their threshold for both cases in Figs. 1 and 2.

4.4. Sensitivity Analysis Sensitivity analysis is a significant tool in the evaluation of systems and their controllers and making decisions. It determines the effect of magnitude change of any parameters and inputs on the system response. Its results can be used to get a broad understanding of parameters effect, validating the model and predicting future results [42]. In this section, a sensitivity analysis is developed for the HCMV infection. For this purpose, five immune system parameters λEK, δEK, λC, δC and kEK in Eqs. (5), (6) are changed and their influences are investigated in detail. These equations are the only ones that can affect the control input design. The effect of each parameter change is studied and demonstrated separately in Fig. 5 to Fig. 10. Note that only variables or outputs that will be influenced considerably in this sensitivity analysis for each case are shown here. The percentage of changes after parameter variations is calculated in comparison with the original case results. We considered 30% change for less sensitive parameters (λEK and δEK) and 10% change for parameters with high sensitivity (λC, δC and kEK). It should be taken into account that each change in λC, δC and kEK will affect the final steady-state value of EKf obtained based on Eq. (9). Regarding Fig. 5, four main variables that change significantly with respect to 30% variation in λEK, are the susceptible cells (S), infected cells (I), viral loads (V) and control signal (ε). The susceptible cells’ final population increased 7% due to a 30% reduction in λEK, while a 30% increment of this parameter results in 8% less susceptible cells for both primary and reactivation cases in

18

Modeling and Control of Drug Delivery Systems

FIG. 1 Results of the primary infection case. The top-left graph is the concentration of susceptible cells (S(t))

which is reduced to less than 350 Cells/μL-blood at the end of the treatment. The top right graph shows the concentration of infected cells (I(t)) that has some fluctuations during the treatment and eventually reaches less than 0.3 Cells/μL-blood. Middle graphs are the concentration of free HCMV (left), and EV(t) as the concentration of HCMV-specific CD8+ T cells (right). Viral loads (V) are more than their defined limit in about one month of the treatment period and show similar fluctuations to the infected cells’ response. Bottom graphs are EK(t), the concentration of allospecific CD8+ T cells that target kidney (left), and C(t), the serum creatinine concentration (right). Response value for EK(t) has tracked its desired value appropriately. The serum creatinine of the blood is also below its defined threshold for the whole treatment process. Solid lines are result values in response to the control input, dashed lines are the thresholds for viral load (3.5 Copies/μL-blood) and serum creatinine concentration (1.2 mg/dL) and dotted dashed line is the desired value of EK(t).

comparison with final values obtained in the original situation. All other variables have the opposite trend, in which the 30% increment in λEK results in 133%, 140% and 7% increase in the maximum populations of the infected cells, viruses and final value of the immunosuppressive drug, respectively. Viral load violated its threshold three times when the value of λEK is raised 30% and it never reached this threshold with the 30% reduction of λEK in both primary and reactivation cases. Also, the final value of the control signal has changed 16% compared to the final magnitude reached in the original situation, with a 30% reduction of λEK. Therefore, the value of λEK has a significant effect on both immunosuppression and antiviral therapies and it should be considered.

Fig. 6 demonstrates the susceptible cells (S), infected cells (I), virions (V) and control signals (ε) as four main values that have noticeable changes in response to a 30% variation of δEK. In both primary and reactivation cases, the final number of susceptible cells (S) approximately increases 5.5% with increasing the magnitude of δEK for 30% and vice versa (compared to the condition with original parameters). However, other variables I and V reduced 66%, 50% in their maximum values and ε decreased 14% in its final value by 30% increment in δEK, respectively. These trends are completely opposite to the ones observed in Fig. 5 for the sensitivity analysis of λEK. This change is highly noticeable in viral load (V) that its magnitude experienced a maximum around 14 Copies/μL-blood in both primary and reactivation

CHAPTER 2 Integral Sliding Mode Control of Immune Response for Kidney Transplantation

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FIG. 2 Results of the reactivation case. Solid lines are the response values, dashed lines are the thresholds for viral load (3.5 Copies/μL-blood) and serum creatinine concentration (1.2 mg/dL) and a dotted dashed line is the desired value of EK(t). Left graphs show the concentrations of susceptible cells, viruses and allospecific CD8+ T cells that target kidney, respectively (from top to bottom). Right graphs represent the concentrations of infected cells, HCMV-specific CD8+ T cells and serum creatinine for the reactivation case, respectively, from top to bottom.

cases with a 30% reduction of δEK and crossed the threshold (3.5 Copies/μL-blood) in three different periods of the treatment that could cause deleterious repercussions for the human health. Furthermore, the control signal undergoes some changes in its magnitude in which the 30% reduction of δEK in the reactivation case makes the control signal to start from the value of 1.1. This is in contrast with the assumption of normalized control input that should be between 0 and 1. Thus, the precise calculation of δEK should be considered as a priority of the HCMV modeling. Figs. 7 and 8 present the response sensitivity by 10% variation in λC. Based on Eq. (9), it is reasonable that a small change in the magnitude of λC will result in noticeable changes in the system’s response, as seen in Fig. 7 and Fig. 8. Susceptible cells had approximately small variations in this sensitivity analysis on λEK and δEK as shown in Figs. 5 and 6. While the final concentration

of these cells experienced a 40% reduction for 10% increment in λC but a negligible 8% change by 10% reduction in λC for both primary and reactivation infections compared to the original condition (Fig. 7). As observed in Fig. 7, infected cells and virions also experienced a wide range of changes such that the maximum values they reached in 10% increment analysis were about five times more than their maximum values obtained in a 30% increment and 30% reduction analysis in Figs. 5 and 6 for λEK and δEK, respectively. Moreover, a 10% decrease in λC resulted in very small variations in the concentration of infected cells and virions. In Fig. 8, small changes are seen in the tracking error (with the order of 10
3) and about 25%–30% changes in the final value of the control signal compared to the result obtained in the original condition. Other important variables (the concentrations of serum creatinine (C) and kidney allospecific CD8+ T cells (EK)) did

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Modeling and Control of Drug Delivery Systems

FIG. 3 Error signals of EK(t) in the primary infection (solid line) and reactivation (dashed line) cases. Both cases

represent a small error with the order of 10
3 which shows that the controller was employed properly to track the desired trajectories suitably.

FIG. 4 Normalized immunosuppressive drug dosage as the control input ε(t). The left graph is demonstrating

the primary infection case and the right graph is for the reactivation case, both reaching 0.7 as the steady-state drug dosage magnitude based on the final value of EKf ¼ 1.5.

not have a considerable sensitivity with respect to λC. For instance, EK followed its desired response suitably as it is evident from its error value in Fig. 8. Fig. 9 depicts the sensitivity analysis results for a 10% change in the parameter δCκ EK (a numerator in Eq. 6). The considerable changes witnessed in the previous

analysis on λC (Fig. 7) is seen to be even more intense. As an instance, the concentration of free HCMV (V) is reached 100 Copies/μL-blood as its maximum value, for both primary and reactivation cases, which is clinically unacceptable as it is 25 times higher than its admissible threshold.

CHAPTER 2 Integral Sliding Mode Control of Immune Response for Kidney Transplantation

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FIG. 5 Sensitivity analysis with a 30% variation in λEK. The left graphs present the primary case and the right graphs demonstrate the reactivation case. Top graphs are sensitivity analysis for the susceptible cells (S) in primary and reactivation cases. The second and third rows depict the infected cells and viruses, respectively. The lower graphs demonstrate the variation of the control signal (ε) in primary (left) and reactivation (right) cases. Solid lines are original values, dashed lines in the third row are viral threshold values, dot-dashed lines are values for 30% increment in λEK, and dotted lines are values for 30% decrement in λEK.

Finally, Fig. 10 illustrates the results obtained for a 10% alternation in κ EK. A high increase in the viral loads (V) and infected cells (I) is observed after a 10% increment, and the viral load in primary and reactivation cases came above its threshold. In contrast, in a 10% reduction of κ EK, the viral load never reached its threshold in neither primary nor reactivation infection.

4.5. Possible Antiviral Therapy Strategies Pre-emptive therapy and universal prophylaxis have been recognized as two main global preventive strategies to treat the HCMV viral load and prevent kidney diseases after its transplantation. According to [43], pre-emptive

therapy should be started when the CMV DNA copies are exceeded from their defined threshold, and this therapy should be continued until the CMV level drops below the detection limit. Universal prophylaxis is also started usually ten days after kidney transplantation. However, there are side effects in both of these treatment strategies, such as risks of getting leukopenia, anemia, thrombocytopenia, nephrotoxicity and cytopenias [44]. In a long-term study [45], after seven years of a clinical trial to compare the prophylaxis and the pre-emptive therapies, results showed that although both strategies were useful, the prophylaxis showed more effectiveness and better performance in preventing CMV infection or kidney disease.

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Modeling and Control of Drug Delivery Systems

FIG. 6 Sensitivity analysis for a 30% variation in δEK. The left graphs show the primary case and the right graphs represent the reactivation case. Dotted lines are for cases with a 30% reduction in δEK, dot-dashed lines stand for the 30% increase in δEK, and solid lines are the real values. Two top graphs demonstrate the susceptible cells in which their final magnitude after the 30% reduction of δEK has declined 11% in both primary and reactivation cases during the whole period of treatment, and it increased 5.5% in the presence of the 30% increment in δEK. Graphs in the second row are for infected cells in which there is a high increment in both cases for the 30% reduction of δEK that would be harmful. In the third row, the viral loads are represented. Dashed lines are the thresholds in which exceeding them for a long period exposes the human body at a high risk, which is observed in the result with a 30% reduction of δEK. The last graphs in the fourth row are the immunosuppressive drug dosages as the control signals that their final values have changed approximately 14% with the reduction and increment of δEK. In the primary case, the control inputs are ascending in original condition and with the 30%-reduced δEK, while it is decreasing for the 30% increment in δEK. This trend is completely vice versa in the reactivation case, where a 30% increment in δEK results in a rising control input and descending ones for original condition and 30%-reduced δEK.

CHAPTER 2 Integral Sliding Mode Control of Immune Response for Kidney Transplantation

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FIG. 7 S, I and V in the presence of a 10% change in λC. Susceptible cells (top graphs), infected cells (middle graphs) and viral loads (bottom graphs) all are representing a significant change in both the primary (left) and reactivation (right) cases.

5. CONCLUSION REMARKS In this chapter, a sliding mode control strategy was developed based on a nonlinear mathematical model of the immune response in renal transplant recipients, with the goal of maintaining a balance between undersuppression case (with the possibility of kidney rejection) and over-suppression case (with a viral load threatening the body). The desired trajectories were defined for allospecific CD8+ T cells that target the kidney. Investigating the sliding mode control law, appropriate dosages of the immunosuppressive drug were determined to track the desired trajectories in both primary and reactivation infection cases. The Lyapunov stability method was employed to prove the tracking convergence to the designed goals. The results showed that for both primary and reactivation infection cases, the controller maintained the serum creatinine below its admissible limit; however, the viral load violated its normal range in some periods of the treatment. As a result, considering an antiviral drug as the second

control input to hold the viral load below its limit can be studied in future work. Furthermore, the control signals reached their final constant values without any chattering in less than 50 days after initiation of treatment. Finally, a comprehensive sensitivity analysis was conducted to evaluate the effects of five main parameters of the model on the system’s response. It is obtained that the system is highly sensitive to parameters in the dynamics of the blood’s serum creatinine, in which the desired final value for the concentration of allospecific CD8+ T cells was calculated. Two other parameters in the dynamics of allospecific CD8+ T cells were also investigated as the next most effective ones in the system’s response based on the presented sensitivity analysis in this work. In future studies, some hybrid and adaptive control strategies can be developed for this nonlinear HCMV infection to take both the immunosuppressive and antiviral drugs into account, in addition to considering the possible modeling uncertainties.

FIG. 8 Error values (top) and control inputs (bottom) with 10% variation in λC. Tracking error for both primary (left) and reactivation (right) cases has the order of 10
3, having less than 1% change in the presence of a 10% variation in λC. The control signal for both cases increase by about 25% with a 10% increment of λC and decrease about 30% with a 10% decrement of λC, comparing their final values to the final value of the original situation.

CHAPTER 2 Integral Sliding Mode Control of Immune Response for Kidney Transplantation

FIG. 9 Sensitivity analysis in the presence of a 10% change in δCκEK. In either primary (left) or reactivation (right) case, variables and control inputs changes are intense with a 10% reduction of δCκ EK. Susceptible cells (top graphs) experienced a sudden drop and a negligible change in reduction and increment of δCκEK, respectively. This trend is also evident for infected cells (second row) and viral loads (third row). Control inputs (bottom graphs) have also undergone significant variations in both reduction and increment conditions.

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Modeling and Control of Drug Delivery Systems

FIG. 10 Sensitivity analysis for a 10% variation in κ EK. Susceptible cells (first row) were highly decreased in a 10% increment case and slightly decreased in a 10% reduction case from their initial status. In contrast, infected cells (second row), free HCMV (third row), and control input values (last row) increased 830%, 800% and 20% with 10% increment in κEK and decreased 80%, 80% and 20% with a 10% reduction in κEK, respectively. These comparisons are presented with reference to the original maximum values of the infected cells and free HCMV concentration, and the original final value of the control input.

CHAPTER 2 Integral Sliding Mode Control of Immune Response for Kidney Transplantation

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CHAPTER 3

Smart Drug Delivery Systems

FATEMEH SALAHPOUR-ANARJANa • PARINAZ NEZHAD-MOKHTARIa,b • ABOLFAZL AKBARZADEHa,c a

Department of Medical Nanotechnology, Faculty of Advanced Medical Sciences, Tabriz University of Medical Sciences, Tabriz, Iran, bStudent Research Committee, Tabriz University of Medical Sciences, Tabriz, Iran, cStem Cell Research Center, Tabriz University of Medical Sciences, Tabriz, Iran

1. INTRODUCTION Living systems can react to environmental stresses to fit their functionality and structure to mutations in nature through the actions of complex sensing mechanisms, regulating and actuating functions, and feedback command operations. Creating biomaterials by dynamic and adjustable features could reply to the microenvironment variations by internal or external stimuli is an important challenge. Responding to circumstance is a key feature of a smart drug delivery system (DDS) [1,2]. They are able to respond to single/multiple or internal/external stimuli such as pH, temperature, enzymes, hypoxia, oxidation, reduction, light, ultrasound, inflammation-responsive, and magnetic or electric fields. Their responses depend on the chemical or physical state of the surrounding circumstances that include swelling, destruction, or solution-to-gel transitions [3]. Since the 1960s, stimuli-responsive polymers have been used as practical biomaterials for biomedical uses including the triggered-release delivery of biological cargos. Also, liposomes have been investigated as alternative DDSs with the capability to reduce the toxic side effects [4]. Stimuli-responsive block copolymers are determined as copolymers whose blocks could undergo comparatively great and sudden, chemical or physical variations in response to minor environment stimuli [5]. Smart macromolecules such as liposomes, micelles, and dendrimers produced by stimuli-responsive polymers which network can pass from a collapsed to an expanded state at transition conditions. Also, their smart surface hydrophilicity and possessing stimuliresponsive groups as a function of a stimulus can provide a responsive interfaces [6]. In the recent years, many efforts have been made to design smart stimuli-responsive DDSs for multiple applications. They are very favorable candidates for producing “smart” platforms which could be applied in numerous biomedical uses from targeted-DDSs to tissue-specific imaging. Recognition of biological

systems can help scientists to design future “SMART MATERIALS” that perform a significant role in effective DDSs and diagnostics devices. Smart-embedded theranostics devices could control by either the patient or doctor and report treatment progress. Structural modifications of drug carriers can develop an infinite number of smart materials [7]. During the response to pathogens, inflammation is a primary natural defense mechanism. Particular features of inflammation microenvironment are overexpression inflammatory and matrix-remodeling enzymes, upregulation of special cell surface receptors, increasing permeability blood vessels, acidic pH, and high oxidative stress have been exploited in the inflammationresponsive DDSs development. Smart macromolecular DDSs could be selectively accumulated in the inflammatory region through passive targeting owing to the improved EPR (permeability and retention effect) and cell-mediated targeting of inflammation-recruited phagocytic cells such as macrophages or direct targeting to particular overexpressed cell surface receptors in the inflammatory sections [8]. This chapter focuses on some of the most common stimuli-responsive nanocarriers that were developed for smart medical technologies. Using stimuli nanocarriers for toxic drugs such as chemotherapy drugs or contrast agents to special organs or tissues of the body with a special condition, that nanocarriers can respond to its condition, could have been one of the most promising fields in the future of nanomedicine.

2. pH-RESPONSIVE DRUG DELIVERY SYSTEMS Various cellular parts have their especial pH points; for example, lysosomal and late endosome pH is approximately 4.5 [9], peroxisome pH between 6.9 and 7.1 [10], mitochondria pH is approximately 8 [11], and cancerous tissues pH is around 6.5 whereas physiological pH is approximately 7.4 [12,13]. Therefore cancerous tissues

Modeling and Control of Drug Delivery Systems. https://doi.org/10.1016/B978-0-12-821185-4.00012-9 © 2021 Elsevier Inc. All rights reserved.

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Modeling and Control of Drug Delivery Systems

have an acidic environment compared with other normal cells and also, pH of infected and inflammatory tissues differs from normal tissues. The pH differences result in a transmembrane gradient between the cellular components in a cell and among the cells. These pH changes in the physiological medium can be exploited to create sufficient pH-responsive DDSs, which would be steady at determining pH and then would start to release its cargo under planned pH. The pH-sensitive polymers with acidic groups (such as –COOH and –SO3H) or with basic groups (–NH2) in the polymer chain can respond by protonation or deprotonation to the alterations in the pH of the surrounding medium by swelling and collapsing. The swelling ratio (SR%) calculate according to the following equation: SR% ¼

Ms  Md  100, Md

where Md and Ms are the masses of samples in the dry and swollen state [14]. The pH sensitivity of the polymeric DDSs could be tailored by changing the hydrophobic carbon chain length and the carboxylated ratio to noncarboxylated alkyl acrylate monomer in alkyl acrylic acid polymers. Also, by using pH-responsive spacers, polymers can conjugate to the drugs, which degrade through the acidic pH situation inside the cancerous tissues or late endosomes to drug release. For example, in a study [15], poly(vinylpyrrolidone-co-dimethyl maleicanhydride) as a pH-responsive controlled release system has been conjugated to doxorubicin (DOX) as a chemotherapy agent and resulted in an increase the accumulation of the drug into the tumor site. Among pH-responsive polymers, poly(β-amino ester) (PBAE) (a biodegradable cationic polymer) has been applied for pH stimuli-responsive DDSs. This polymer degrades under the acidic conditions to release its cargo and resulted in an important accumulation of drugs in the cancerous sites compared with a nonresponsive polymer-based delivery system [16]. When the pH rate is reduced below 6.8, a rapid and continuous diameter enhance is detected, which is caused by the tertiary amines protonation in PBAE [17,18]. PBAE has also been used to produce nanovaccines. In a study [16], mannosemodified PBAE nanovaccines by pH-sensitive features were developed to co-deliver the melanoma cancerousassociated antigen polypeptide tyrosinase-related protein-2 (Trp-2) and the TLR4 agonist monophosphoryl lipid A (MPLA). Also, PD-L1 antagonist (an immune checkpoint inhibitor) was administrated with PBAE nanovaccines to decrease immunosuppression in the cancerous microenvironment to delay melanoma

progress. Mannosylated Trp-2 and MPLA-loaded PBAE nano-vaccines could be targeted and matured dendritic cell (DCs) and in the following to promote antigenspecific cytotoxic T lymphocyte activity against melanoma. These researchers declared, that combination therapy with PD-L1 antagonist farther enhanced antitumor efficacy by 3.7-fold and prolonged median survival time by 1.6-fold more than free Trp-2/MPLA inoculation and DC-targeting PBAE polymers as a nanotechnology platform could be a hopeful strategy to efficiently induce potent immune responses and synergistic antitumor effects.. The pH-sensitive copolymers such as N-(2-hydroxypropyl)methacrylamide (HPMA) and linear PEG copolymer nanocarriers are other platforms that have revealed success in DDSs of drugs to the cancerous targets. One of the major causes of chemotherapy defeat is multidrug resistance (MDR) of the tumors. MDR is related to overexpression of ATP-fuelled efflux pumps such as P-glycoprotein (P-gp) that pumps a major volume of chemotherapeutic drugs out of the cancerous cells. In the researches related to cancer, it is stated that more than 90% of treatment failure of metastatic cancer is caused due to MDR, and studies have noted that approximately 40% of breast cancer cells express the MDR proteins [19]. Combination chemotherapy is a useful approach to overcome MDR of the cancerous tissues. Effective combination chemotherapy needs the co-loading, co-delivery and controlled release of two various drugs with different chemophysical features. Bypassing the MDR by stimuli-responsive systems has been investigated in various studies. For example, in a study [18] a pH-sensitive graft copolymer as a dual-drug co-loaded nanoparticles (NPs) has been developed. Poly(β-amino ester)-g-β-cyclodextrin (PBAE-g-β-CD) through the Michael addition polymerization was synthesized and used to co-deliver DOX and adjudin (ADD) as a mitochondrial inhibitor against MCF-7/ ADR xenograft tumor model. DOX was conjugated to 1-adamantaneacetic acid (Aa) to produce a prodrug then encapsulated in the hole of cyclodextrin and ADD was encapsulated by PBAE. Co-loaded NPs (AaDOX + ADD@PC) could achieve precise control of drug loading, increasing cellular uptake of the drugs and cases to induce apoptosis by mitochondrial dysfunction. The scientists stated that the Aa-DOX +ADD@PC could have effective growth inhibition against MDR proteins both in vitro and in vivo through the synergistic effect of ADD and DOX, which can be afforded a promising strategy and resulted in high therapeutic outcomes for MDR tumors (Fig. 1). Liposomes are one of the most extensively used DDSs and many liposomal nanomedicines have been

CHAPTER 3

Smart Drug Delivery Systems

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FIG. 1 A pH-sensitive NP system, poly(β-aminoester)-g-(β-cyclodextrin) (PBAE-g-β-CD, PC) copolymers were synthesized to realize the co-loading and programmed release of DOX and ADD. ADD release was triggered in low pH in endo/lysosome after endocytosis and then DOX was hydrolyzed to produce a continued release in cancerous cells and finally apoptosis induced by mitochondrial dysfunction (the Aa-d-α-tocopheryl polyethylene glycol succinate (TPGS) conjugate improved the serum stability and biocompatibility of the NPs) [18].

approved for clinical applications from their discovery till now [20]. By modifying liposomes by pH-sensitive lipids such as dioleoyl phosphatidylethanolamine (DOPE), poly(organophosphazenes), and cholesteryl hemisuccinate (CHEMS) or by conjugation of pH-sensitive polymers on liposome surface have achieved an acid sensitivity and controlled DDSs that can easily internalize into the cell with endocytosis pathways and then fuse to the endosomes membrane to deliver its contents into the cytoplasm [21–23]. Since acidic sensitive systems are stable at physiological pH but get protonated over the acidic situations that cause destabilization of these structures and facilitating delivery of their payloads such as drugs, DNA, siRNA, toxins, and antigens [24]. Recently, in a study [25], pH-sensitive paramagnetic liposomes have been used as a probe in magnetic resonance imaging (MRI) for monitoring acidic pH in tumors. This pH-sensitive liposomal MRI contrast agent has been the potential to use as a marker of low pH in the cancerous interstitium. Dipalmitoylphosphatidylethanolamine (DPPE)/dipalmitoylglycerosuccinate (DPSG) liposomal GdDTPA-BMA displayed a marked increase in T1-relaxivity (r1) when the pH decreases below the physiological level in the blood, due to aggregate of the liposomes and following the leakage of GdDTPA-BMA. By grafting of polyethylene glycol (PEG) to DPPE, blood circulation time increases owing to sterical stabilization after intravenous injection. The scientists declared that the relaxometric pH-response of the DPPE/DPSG/DPPE-PEG system reduced as a function of DPPE-PEG mol percent, therefore, a compromise would be needed between long blood residence time and an appropriate liposomes pH-sensitivity. They introduced a strategy to compensate the decreased

pH-sensitivity, thus, they encapsulated Gadofosveset (a low-molecular weight Gd-chelate with high affinity for albumin) inside DPPE/DPSG liposomes that caused a considerably higher relaxometric response due to the release of gadofosveset at low pH and subsequent binding to albumin. After the internalization of nanoparticles (NPs) into the cell through different endocytosis pathways, early endosomes are first where they delivered then mature to the late endosome and finally to the lysosome. Therefore the endosomal escape of the nanocarriers is one of the main biological limitations related to DDSs. If nanocarriers could not escape from this rigid chemical environment, it will result in the degradation of NPs and its content. Membrane fusion and membrane destabilization are mechanisms that have been explored the endosomal escape. However, nonspecific membrane fusion can be the limitation of these methods. To overcome this limitation in a study [26], pH-responsive liposomes were made up of 3β-[N-(N’,N’-dimethylaminoethane)carbamoyl]cholesterol hydrochloride (DC-liposome) plus embedding a tiny percentage of the cationic nitrogen of the ammonium moiety in their structures. This small percentage was enough until DC-liposome to display pH-responsive features while their biocompatibility was maintained. The fluorescence-based experiment proved this fact that DC-liposomes could show pH-dependent cationic features by DC-moiety protonation at acidic pH. Also, the endosomal colocalization study exhibited localization of DC-liposome in the early endosome compared to the late endosome is higher, therefore, suggesting endosomal escape is more possible and elevated cationic and fusogenic properties of DC-liposomes at acidic pH can mediate membrane fusion with the anionic endosomal membrane by electrostatic

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interactions, resulted in the endosomal escape. The researchers declared that DC-liposomes have the potential to break the endosomal barriers to improve the therapeutic efficacy of encapsulated cargoes. Also, it was investigated that DOX-loaded DC-liposome presented higher cytotoxicity than the free DOX because of DC-liposomes’ endosomal escape capability. Another class of nanocarriers is micelles which have been developed pH-responsive DDSs [27]. Copolymers can form micelles and by protonation of the groups such as carboxylic or amines on the surface of the watersoluble copolymer, micelle can response to the pH variation and able to start degradation and release own payloads [28]. For example, poly(L-lactide)-b poly (2-ethyl-2-oxazoline)-b-poly(L-lactide) (PLLA-PEOPLLA) triblock copolymers and diblock copolymers such as (PEO-PLLA) have been used in the pH-responsive micelle’s systems [24]. Dendrimers are another class of nanomaterials that have been investigated to develop pH-responsive DDSs [29,30]. Dendrimers composed of pH-responsive monomers such as 2,2-bis(hydroxymethyl)propanoic acid or by modifying of terminal ends with acidsensitive groups including hydrophobic acetal groups that are conjugated to desired cargo to produce a pH responsive delivery system that able to cleave off the dendrimer structure in acidic situation resulting in the cargo release [31,32]. In a study [13], The 3.5th generation of dendritic chitosan-coated silica magnetic (DCSM) NPs with a lot of amino and carboxylic acid groups on the surface were synthesized and two different anticancer drugs, DOX and methotrexate (MTX) were loaded to DCSM NPs by electrostatic interactions and simultaneously intracellular delivery of these drugs was investigated. The amino functional groups in the chitosan polymeric chain are responsible for its solubility owing to the protonation in acidic media, moreover, NPs with chitosan coating is biodegradable and biocompatible. At low pH rates, the amine functional groups (–NH3+) and carboxylic acid (–COOH) on the nanocarrier surface were protonated, and at high pH rates, Fe-OH, carboxylic acid, and Si-OH groups are deprotonated (Fe-O, COO, and Si-O) resulting negatively charged zeta potential and results in DCSM NPs act as a pH-responsive nanocarrier. The scientists reported that DOX@MTX@nanocarrier has shown high drug release at cancerous conditions (pH 5.4 and 4.0) while having a low drug release in the simulated bloodstream at physiological circumstances (pH 7.4) and co-delivery of MTX with DOX to MCF7 cell lines can generate synergistic anticancer effects and enhance the treatment efficacy and decrease toxic side effects of drugs.

3. REDOX-RESPONSIVE DRUG DELIVERY SYSTEMS It is important to develop types of block copolymers that are able to respond to dual redox (both oxidants and reductants) responses under mild conditions. The redoxresponsive polymer through a disulfide (SdS) or diselenide (SedSe) bonds that are embedded in liposomes, micelles, and dendrimers structures can be exploited as glutathione (GSH)-triggered drug release at a high GSH level environment. Relying on this fact that there is a redox potential change between the oxidizing intracellular space and the reducing extracellular area which can be potentially used to deliver the redox-sensitive DDSs into the cells that are prone to redox reactions due to components including SdS or SedSe linkages [5,33]. Redox-sensitive DDSs after entering the cancer cell by endocytosis or other entrance ways, disulfide linkage in their structures start to disrupt in the late endosomes environment rich of GSH and resulted in facilitating their payload releasing. GSH is a small tripeptide derived from three amino acids: glutamate, cysteine, and glycine that synthesized naturally in the cell cytosol and within the cell, above 98% exists in the thiolreduced form (GSH). GSH is present inside specific intracellular sections such as mitochondria and endoplasmic reticulum and acts as a coenzyme, cofactor, and/or substrate for some of the enzymes, and participate in some of the redox reactions [34]. In normal cells, low GSH concentrations make stable disulfide bonds of proteins but cancerous cells demand high levels of GSH (more than three times higher) due to rapidly cells proliferating [33]. A certain level (1–10 mM) of the presence of GSH has been regarded as an important distinction between normal and cancerous tissues. The GSH pathway plays a key role in the reduction of the disulfide linkage in the reducing intracellular environment by maintaining an elevated level of GSH [35]. The disulfide crosslinking also guarantees the stability and integrity of the redox-sensitive DDS and limits the possibilities of early release of the payload until reaching to a rich environment of GSH. A stimuli-responsive system such as redox-sensitive DDS could be used as a reliable delivery biosystem for nucleic acid-based therapeutics such as ASN, pDNA, siRNA, and miRNA or other biomolecules such as proteins and peptides for the treatment of various genetic diseases. Since these biomolecules are highly prone to degradation, therefore, the successful delivery of them is a great challenge and requires many efforts. One of the most used approaches to exploit the redox stimuli systems is applying positively charged groups such as polyaspartamide in the polymer backbone that entrap

CHAPTER 3 nucleic acids electrostatically, and on the other hand, disulfide linkage has embedded in the polymer chain resulting in releasing at the right time/place. For example, gelatin thiopolyplexes have been used as potential redox responsive nanosystems for nucleic acids such as pDNA delivery. Nucleic acids have been directly conjugated to the positively charged thiolated polyethyleneimine to form polyplexes that have shown successful delivery of DNA into the cells with high transfection yield. In yet another study, GSH-sensitive-polymer coated chitosan particles were applied for designing biosystems stabilized with a disulfide bond to provide gene delivery. Including FDA approved redox-responsive drug can mention is Mylotarg® (gemtuzumab ozogamicin) that is a redox-responsive anti-CD33 antibody conjugate that has approved by the FDA to treat adults with newly diagnosed CD33-positive in acute myeloid leukemia (AML) or patients over the age of 2 with CD33-positive AML whose disease returned or did not respond to previous treatment [36]. Liposomes-based redox-responsive DDSs have also been investigated to enhance the delivery facility to the special target tissues. These liposomes are formed by phospholipids are linked by a cleavable disulfide bond and stay stable until reaching the reducing environment inside the target cells where the disulfide bond cleavage by GSH results in destabilization and deliver cargo. Thiocholesterol lipid-based liposomes have also been used successfully gene delivery into the cells with a reducing environment.

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TOS (tocopherol succinate) is a complex of vitamin E family and is used as a potential antineoplastic. In a study [33], TOS was grafted to hyaluronic acid (HA) through disulfide bonds and resulted in HA-SS-TOS conjugates then assemble into amphiphilic nanomicelles thereafter paclitaxel was loaded into micelles and create HA-SS-TOS-PTX particles which can separate when exposed to GSH-rich environments in vitro and in vivo. Furthermore, HA-SS-TOS-PTX micelles due to HA could be proficiently taken up with the epithelial cancer cells that overexpress CD 44 because CD44 is a receptor of HA and HA can regulate cell proliferation and migration through CD44 (Fig. 2). In another study [37], a redox-responsive drug delivery micelle system has been reported by coupling deoxycholic acid (DA), as a hydrophobic moiety, with heparosan, as a linear polysaccharide can induce fast and specific internalization into the cancer cells and exhibit a PEG-like “stealth” feature, through a disulfide bond to form HSD micelles that is loaded with DOX for the laryngopharyngeal carcinoma treatment. The amphiphilic glycan conjugates (HSDs) self-assembled into nanoscale micelles in an aqueous medium thereafter DOX was loaded into the HSDs micelles. DOX@HSD micelles internalize by clathrin-mediated endocytosis pathway into the FaDu cancer cells and display GSH-triggered drug release response and this group reported an approximately 100% release rate in a 10-mM GSH environment can induce inhibition of FaDu cancer cells a minimum of 10-fold selectivity

FIG. 2 Redox-responsive HA-SS-TOS-PTX micelles exposed to GSH-rich environments in epithelial cancer

cells release PTX after SdS linkage cleavage by GSH [33].

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Modeling and Control of Drug Delivery Systems

FIG. 3 Structure of PEG-PUSeSe-PEG copolymers and redox-responsive disassembly of PEG-PUSeSe-PEG micelles [5].

relative to that of COS-7 normal cells. DOX@HSD can be a promising redox-responsive DDS besides good biocompatibility in the clinical treatment of laryngopharyngeal carcinoma given that the GSH level is two to three orders higher in the cytosol of cancer cells compared with that in the extracellular environment. The high GSH level will trigger a fast micellar disassembly, which leads to the release of DOX into the cytosol of cancer cells. Selenium compounds have been extensively applied in the pharmacochemistry as an antioxidant in the glutathione peroxidase (GPx) activity, among which diselenide is a proper dual redox response because it has good activity in the presence of oxidants and reductants. Commonly, SedSe bonds oxidize to seleninic acid in the presence of oxidants and reduce to selenol in a reducing condition. In a study [5], PEG-PUSeSe-PEG amphiphilic triblock copolymer was synthesized with one hydrophobic diselenide-containing block and two hydrophilic PEG blocks and its self-assembly behavior in water and in the oxidants or reductants presence was investigated. In this work, the diselenide groups were first imported into a diol composition, which obtains the desirable solubility. Then the diselenide containing polyurethane (PUSeSe) blocks were synthesized via polymerization of toluene diisocyanate (TDI) in light excess with diselenide-containing diols and finally terminated by PEG monomethyl ether. PEGPUSeSe-PEG self-assemble in an aqueous environment and create micelles. It was predictable that the SedSe bonds would cleavage in the presence of oxidants

(H2O2) or reductants and the consequent disassembly of the PEG-PUSeSe-PEG micelles occur. After adding H2O2, the micellar structure was turned into irregular aggregates and then decomposed into small aggregates of various nanometers in size during 3 hours of oxidation. This group has shown that the micelles of PEGPUSeSe-PEG are also responsive to the reduced environment because the Se-Se bonds are completely sensitive and prone to cleavage in the presence of reductants such as reduced glutathione (GSH). Hence when GSH was added to PEG-PUSeSe-PEG micelle, they became broken and, then, tiny aggregates formed during the reduction process. Therefore it is confirmed that the diselenide groups in the polymer backbone can form micelles with dual redox responsiveness (Fig. 3).

4. THERMORESPONSIVE DRUG DELIVERY SYSTEMS The cancer cells are more sensitive to heat-induced damage than the normal cells because of their rapidly dividing nature [38]. This fact causes to use hyperthermia as a supplement treatment besides chemotherapy and radiation for the destruction of cancer cells. Incorporation of temperature-sensitive components in the nanocarriers such as liposomes, micelles, and dendrimers along with other NPs such as gold NPs or superparamagnetic iron oxide particles (SPIONs) that can produce heat in external stimuli presence including alternating magnetic field (AMF) or near-infrared (NIR) light. These structures have developed hyperthermia as an adjunct to radiotherapy and chemotherapy for the treatment of

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FIG. 4 Poly N-isopropylacrylamide (PNIPAAm) temperature transition behavior (LCSTNIPAAm ¼32°C) and (light yellow dots are water molecules).

various cancer. Use of these NPs to localize carriers with the antitumor drugs in the cancerous sites to ensure that only tumor cells are exposed to elevated temperature without side effects for normal cells. This coupling strategy has shown promise in the delivery of drugs to the target tumors and leads to improved accumulation and efficacy of the drug [39–42]. Temperature-sensitive polymers also named as shape-memory polymers have been applied in the development of minimally invasive surgery biomedical devices [43,44]. In aqueous solution, temperaturesensitive polymers display low critical solution temperature (LCST) depending on their transition behavior from monophasic to biphasic when the temperature is changed wherein below LCST they are water-soluble but become insoluble above it. This feature makes them as thermoresponsive DDSs. Most common LCST polymers have been investigated as temperature-sensitive polymers are poly(N-isopropylacrylamide) (NIPAAm), poly(vinyl amide) families, poly(oligoethylene glycol (meth)acrylate) families. One of the most used thermoresponsive polymers is poly(NIPAAm) and its copolymers that have been used as thermosensitive amphiphilic polymers. In the LCST temperatures, hydrophilic/hydrophobic interactions are equilibrium

whereas, at temperatures below LCST, hydrogen bonding between the water molecules and the acrylamide (AAm) groups of poly(NIPAAm) exists and hydrophilic interactions become predominant, thereforeNIPAAm chains are solubilized in water, but, above the LCST, hydrogen-bonded network is rapidly interrupted and water molecules release entropically favored of the polymer chains, and hydrophobic interactions become predominant and the transparent liquid turns cloudy and aggregate to form clusters and finally turns to gel. Cloud point is dependent on polymer concentration and the lowest cloud point is referred to as the LCST [3, 4] (Fig. 4). However, the LCST temperature of poly(NIPAAm) is lesser than the regular human body temperature (32°C < 37°C), therefore its structure must be further tuned to enhance its LCST to a range that is compatible with human health uses. To adjust the poly(NIPAAm) LCST above the temperature of body, water-soluble units including methacrylate (MAA) or acrylic acid (AAc) groups have been incorporated into the poly(acrylamide) (poly(AAm)) backbone, these groups by increasing the overall hydrophilicity of the polymer chain cause an increase in the LCST of poly(AAm) and while turn it into a good stimuliresponsive DDS [45,46].

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FIG. 5 Schematic demonstration of the temperature-dependent drug-release kinetics from the chemically

cross-linked thermosensitive hydrogels and comparing two different release profiles in two different states. (A) At TLCST, the release profile is a two-step manner: the fast release of drug molecules occurs by the initial depreciation of the hydrogel, succeeded by the slow release of drugs from the condensed structure [4].

Poly(AAm) can be cross-linked to form macroscopic gel-like networks and maintain its structural integrity against the dissociation of single polymer chains during the reversible solvation/desolvation process across the LCST. Cross-linked polymer gels due to possessing structural integrity are similar to solid matters but in fact, they own the features of both fluids and solids altogether. These cross-linked gels can hold up to 99 wt% of solvent molecules inside a low-density network. Therefore, small guest molecules can rapidly move in and out of the network by the diffusion process during reversible volume changes. Hydrogels have been used for the localized delivery of an extended range of proteins, drugs, and oligonucleotides to date. According to the nature of the cross-links between polymer chains in solution, polymer gels are categorized as having either physical (noncovalent) cross-links or chemical (covalent) cross-links. In physically cross-linked gels, hydrogen bonding, hydrophobic and ionic interactions hold the gel network together and could be reversibly manipulated under specific conditions. The degree of cross-linking in physically cross-linked gels depends on concentrations of polymers and temperature. Chemically cross-linked gels are stronger because the polymer chains are connected by covalent bonds and the release kinetics of cargos from cross-linked poly(AAm) hydrogels depend on the cross-linking degree that could be accurately controlled by temperature during the synthesis. The degrees of cross-linking in the polymer gels tune the porosity of the hydrogels. Higher degrees of crosslinking in the polymer gels cause small pore size and resulted in slow-release rates of the embedded cargo.

The release kinetic of a chemically cross-linked polymer gel network is regulated by its LCST and the crosslinking degree. These macroscopic gels were frequently used as materials for localized drug-delivery or tissue engineering due to their ability to sustain drug release over an extended duration is extremely acceptable. In recent years, the nanoscale forms of these gels have been developed as a platform for the systemic delivery of biologically active agents, these agents increase drug efficacy and reduce side effects because of enhancing the bioavailability and good pharmacokinetic profiles [47,48]. Fig. 5 shows the release mechanism of a chemically cross-linked polymer gel network [4]. Dual stimuli-responsive systems have also been investigated in many studies. For example, in another study [49], a dual pH/temperature-sensitive superparamagnetic nanogel (diameter of less than 30 nm) has delivered two anticancer drugs, MTX and DOX. Magnetic nanogels showed an apparent pH/thermotriggered controlled drug release in a stable behavior and that was distinguishable between tumor tissues and normal tissues. Co-administration of MTX with DOX also shown higher cytotoxicity to MCF7 and MDA-MB-231 cell lines compared with single drugloaded or free dual drug forms. The scientists asserted drug-loaded pH/thermo-sensitive nanogels have the potential to be applied for cancer treatment because they have a low early drug release during blood circulation but when they reach tumorous tissue, drug release becomes rapid. The membranes of temperature-sensitive liposomes are formed from different temperature-

CHAPTER 3 response phospholipids. By mixing various lipids with various gel-to-liquid transition temperatures could design liposomes with the favorite gel-to-liquid transition. For example, liposomes composed of dipalmitoylphosphatidylcholine (DPPC) as primary lipid and distearoylphosphatidylcholine (DSPC) as a co-lipid undergo phase-transition from gel-to-liquid crystalline and lamellar-to-hexagonal transition resulted in the release-loaded elements during such transitions [50]. Using of temperature-sensitive liposomes in combination with localized hyperthermia can increase the temperature-dependent effects and particularly releases the trapped drug in the heated tumor tissue. In a study [51], a 1.2-dipalmitoyl-snglycero-3-phosphoglyceroglycerol (DPPGOG)-based liposomal formulation was synthesized that enabled long circulation time with durable and effective drug release under moderate hyperthermia (41–42°C) and DPPGOG facilitates temperature-triggered drug release from these liposomes. The researchers declared that grafting PEG to lipids can possibly be used for clinical applications and the mean area under the curve for tissue drug concentration was raised more than sixfold by these liposomes compared with nonliposomal drug delivery. Integrating stimuli-responsive polymers and liposomes together to obtain triggerable, targetable and theranostics capabilities nanoscale platforms can fulfill all the disadvantages of these structures when used lonely, such as lack of drug-release triggers and the instability of naked liposomes. Polymers can rapidly experience a phase transition to a gel which triggers conformational variations in the chains of polymer and a coil-to-globule transition in response to external stimuli including variations in temperature. Polymers having LCST could be applied to prepare thermosensitive liposomes. Below the LCST, polymers stabilize the liposomes in their hydrated form, but above the LCST, polymers destabilize the liposomal due to temperature transition behavior and liposomal integrity structural starts to disintegrate and resulting in the release of its cargo. Using NIPAAm copolymers as LCST polymers in the structure of the thermoresponsive liposome has been reported in many studies [52–54]. The liposomes membrane could be simply modified with the triggered response polymers and small molecules to facilitate releasing of cargo by stimuli. Through a postsynthesis modification strategy, a network of stimuli-responsive polymers can be integrated onto the liposomes surface to produce a multifunctional nanoscale DDSs that allows for multidrug loading, triggered drug, release targeted delivery, and theranostic abilities.

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5. HYPOXIA-RESPONSIVE DRUG DELIVERY SYSTEMS The penetration of drugs to tumors is an important issue therefore, designing carriers that are able to deliver drugs efficiently, has been a challenging matter forever. The cancerous matrix is a heterogeneous environment composed of irregular blood vasculature and raised interstitial fluid pressure prevents penetration of drugs to the solid cancerous matrix and resulted in weak therapeutic effects. One of the promising nanocarriers, to acquire satisfactory efficacy, is hypoxia-responsive DDSs that exploit features of the cancerous environment and could assist as an important therapeutic target [55]. Based on our knowledge of cancerous tissues, tumor metabolizing is based on glycolytic pathways that provide oxygen and food. Cancerous tissue always faces a shortage of oxygen (less than 5 vs 40–60 mmHg in healthy tissues [56]) and food because of its high metabolization leads to an acidic condition that could be exploited by hypoxia-responsive DDSs to deliver drugs efficiently. The zeta-potential of the nanocarrier surface could be increasingly altered by replying to the hypoxia gradient therefore hypoxia-responsive nanocarrier could enhance the positive surface charge by responding to hypoxia gradients and resulted in deep penetration in the cancerous matrix. Studies have shown that positively charged nanocarriers have a better penetration capability due to overcoming the hydraulic pressure gradient and exhibit efficient binding plus internalization to angiogenic endothelial cells (Fig. 6) [57]. Given that, the reticuloendothelial system removes positively charged nanocarriers from circulation but stimuli-responsive nanocarrier could produce positive charge on the carrier surface in the cancerous site by the response to cancerous condition such as lack of oxygen (hypoxia) and acidic condition (refer to part 2 in this chapter). Nitroimidazoles [58], nitrobenzyl alcohols [59], and azo linkers [60] are used in nanocarrier structures as a hypoxia-responsive group which can reduce under hypoxic conditions and result in variations in surface charge and hydrophobicity or hydrophilicity of the nanocarriers (Fig. 7). In a study [59] anticancer theranostics FDU-DB-NO2 drug delivery system designed for solid tumors treatment that can specifically be activated by hypoxia. This prodrug has investigated to deliver an anticancer drug floxuridine (FDU) and a fluorescence dye precursor 40 -(diethylamino)-1,10 - biphenyl-2-carboxylate (DB) for selective two-photon imaging and a hypoxic trigger 4-nitrobenzyl group that can track real-time drug release in hypoxia condition of solid tumors. 4-Nitrobenzyl group of FDU-DB-NO2 reduces by NTR/NADH under

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Modeling and Control of Drug Delivery Systems

FIG. 6 Schematic of the applicability of a hypoxia-responsive nanocarriers with stable circulation capability

and improved penetration efficacy in the hypoxic condition in the cancerous site due to increasing the positive surface charge via responding to hypoxia gradients. [57].

FIG. 7 Groups and linkers are used in nanocarrier structures as a hypoxia-responsive.

hypoxic conditions and a fluorescence dye 7-(diethylamino) coumarin generates, therefore, the developed FDU-DB-NO2 is locked in normal cells, whereas it is unlocked in cancerous cells by hypoxia and results in fluorescent dye generation along with FDU release. The amount of fluorescent dye production and FDU release were regulated by hypoxic status and enhanced with the diminution of the O2 concentration. In this study has been confirmed that FDU-DB-NO2 shows high cytotoxicity against hypoxic MCF-7 and

MCG-803 cell lines whereas no cytotoxicity against normoxic BRL-3A cells and exhibited efficient inhibition on tumor growth of MCF-7-cell-inoculated xenograft nude mice (Fig. 8). The hypoxic status, concentration of O2, and amount of FDU release in cancerous tissue and spheroids were imaged with fluorescence. In another study [60], dually hypoxia- and singlet oxygen-responsive polymeric micelle nanosystem has been reported to deposit photosensitizer in the disease site and resulted in improving photodynamic therapy (PDT)’s anticancerous efficacy. In this study, hydrophilic methoxypolyethylene glycol (mPEG) block attach to hydrophobic polypeptide block (aspartic acid) hypoxia-responsive azobenzene linker and producing a self-assembling amphiphilic copolymer conjugated with imidazole as the side chains. mPEG-Azo-PAsp-IM micelles (189  19 nm) achieved by the self-assembly process was loaded with photosensitizer chlorin e6 (Ce6). Azobenzene linker as the hypoxia-sensitivity moiety would able the dePEGylation and improved micelles cellular uptake, while the imidazole as the singlet oxygen-responsive moiety can lead to rapid Ce6 release and micelle disassembly. Moreover, the singlet oxygen via reducing the oxygen level due to short half-life resulted in enhanced internalization and

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FIG. 8 Structure of FDU-DB-NO2 and its function in hypoxic cancerous cells [59].

increase the intracellular Ce6 concentration. The cellular uptake study established that the dually responsive micelles could deliver significantly more Ce6 to the lung carcinoma (LLC) tumor-bearing mouse model, which resulted in an extremely enhanced photosensitizer delivery and impression of PDT antitumor. Radiochemotherapy is the common clinical treatment for malignant glioma but in this method, there are 2 significant challenges; first, there is increased resistance to radiation owing to intratumoral hypoxia and blood-brain barrier (BBB) limits the delivery of the chemotherapeutic agent to the brain. In a study [58] an MLP (hypoxic radiosensitizer-prodrug liposome) has developed to deliver DOX to the glioma cancerous and to overcome the challenges mentioned earlier and realizing a synergistic chemo-/radiotherapy for treating malignant glioma. DOX is an important anticancer drug that is used to treat various types of cancers through the inhibition of progression of the enzyme topoisomerase II and DNA damage induction. Since DOX has a very weak effect when utilized in glioma therapy, because it cannot cross the BBB, and hence cannot achieve adequate concentrations in glioma cells to produce satisfactory toxicity. However, in brain tumors, some of the BBB architecture is damaged and allowed nanocarriers to cross the BBB. In this work, hypoxic radiosensitizer nitroimidazoles were conjugated with lipid molecules by a hydrolyzable ester bond to make MDH. MDH was mixed together with DSPE-PEG2000 and cholesterol to form MLP liposomes, which have strong radiosensitivity and to raise DOX release under cancerous hypoxic condition, owing to the chemical variation of

nitroimidazoles under hypoxic conditions (Fig. 9). According to this fact, hydrophobic nitroimidazole can converts to hydrophilic aminoimidazole through the selective bioreductions series when presented by hypoxic situations and causes to cleave liposomes structures and releasing loaded DOX (Fig. 7). This group figures out that the combination of MLP/DOX and radiotherapy significantly inhibited glioma growth and is a hopeful applicant as a DOX delivery system to enhance the antitumor treatment effects on glioma. According to studies, small variations in the structures and functions of block copolymer aggregates could result in the release of loaded cargoes, which will have great potential in the fields of controllable drug and gene delivery as a smart delivery system that can response its surrounded condition.

6. OTHER EXTERNAL STIMULI-RESPONSIVE SYSTEMS Light-responsive carriers through the external light illumination represent a manner to trigger release of drug at the desired target [61]. The photosensitive systems could reach the on-off drug release event due to the close or open the nanostructures when stimulated through either a 1-time or repeatable irradiation of light [62]. Though, considering the light wavelength limitation for practical therapy, presently the light penetration depth confines the noninvasive uses for deep tissues [47]. In this regard, for example, a light-controllable red blood cellbased DDS with decorating Ce6, a photosensitizer, on the cell membrane for loading and remotely controlled release of DOX was successfully engineered (Fig. 10) [63].

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Modeling and Control of Drug Delivery Systems

FIG. 9 Formation of MLP/DOX liposomes drug-delivery system as the hypoxia-triggered, radiotherapy

sensitive and its mechanism as hypoxic cell radiosensitizer to DOX release into the cell nuclei to obtain synergistic chemo-/radiotherapy glioma treatment [58].

FIG. 10 The development of DOX@RBC-Ce6, in which DOX was loaded inside RBCs, whereas Ce6 was

inserted into membrane of RBC through hydrophobic interaction. With light irradiation (660 nm), the membrane structure of RBC would be disrupted in this approach and the loaded drug would be rapidly released [63].

CHAPTER 3 Magnetic responsive systems can offer a noninvasive platform for spatially and temporally control of the carriers to the targets, and drugs are released by the exposure of the external magnetic field [64]. The magnetic NPs (MNPs) large surface-to-volume ratio offers abundant active sites for conjugation of biomolecules, so permitting precise design to achieve their intended smart purposes by using a localized external magnetic field, including therapeutic delivery, long circulation time in the bloodstream, and target specificity. In the study, a dual magnetic/pH responsive nanoapproaches based on PEGylated Fe3O4 NPs for DOX delivery have developed [65]. Newly, ultrasound (US) has been widely applied in clinics for therapy and diagnosis owing to its high safety and intrinsic tissue penetration [66]. The US-responsive nanocarriers development for ultrasonography increases US methods to be an effective and unique process to capture drug carriers and at the desired sites trigger drug release by tuning the frequency of US, duty cycles and exposure time [67]. In this regard, the stable ultrasound-responsive curcumin-loaded chitosan/perfluorohexane nanodroplets for on-demand drug delivery and contrast-US imaging was developed. The synthesized nanocarrier efficiently released entrapped molecules of curcumin under the action of ultrasound in a controlled manner and presented strong US contrast even at low concentrations [68]. In the other study, an ultrasound-responsive O-carboxymethyl chitosan nanodroplet for controlled DOX delivery was reported [69].

7. CONCLUSIONS In the past decades, nanosystems have developed as drug carriers. Nanocarriers have shown more capability of increasing retention time in the bloodstream and the chance of accumulation in target malignant sites. On the other hand, they are more noticeable because they generate synergistic antitumor effects besides the drugs. The environment and circumstance of the cancerous sites and its surrounding, such as acidic atmosphere, elevated GSH levels, degradative enzymes in lysosomes and endosomes and hypoxia condition can be exploited for designing smart DDSs that are more efficient than nonsmart carriers to deliver treatment/tracking agents to cancer cells and expected by using these smart systems the drug’s adverse side effects related to high dosage is decreased. Among the mentioned stimuli-responsive biomaterials in this chapter, polymer-based approaches are the most hopeful because they can be produced in massive scales and by the broad chemical functionalities range, improved postsynthetically in a simplistic way

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and prepared into biomaterials in many various forms such as surface films, patterns, solids, or solutions. Polymers could be managed to vary their molecular formations in response to external stimuli including temperature, resulting in variations in transparency, density, conductivity, solvent-uptake capability, or swelling degree. Scientists are hopeful to develop effective treatments with fewer side effects for chronic and incurable diseases. These strategies may afford promising platforms for selective imaging, real-time tracking drug release, and personalized solid cancerous treatment.

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CHAPTER 4

Polymeric Transdermal Drug Delivery Systems

MOHAMMAD SHAHROUSVANDa • NADEREH GOLSHAN EBRAHIMIa • HADI OLIAIE b • MAHSA HEYDARIa • MOHAMMAD MIRa • MOHSEN SHAHROUSVANDc a

Polymer Engineering Department, Chemistry Engineering Faculty, Tarbiat Modares University, Tehran, Iran, bDepartment of Polymer Engineering and Color Technology, Amirkabir University of Technology (Tehran Polytechnic), Tehran, Iran, cCaspian Faculty of Engineering, College of Engineering, University of Tehran, Rezvanshahr, Iran

1. INTRODUCTION Drug delivery systems (DDSs) consist of a drug component as an active agent and a carrier component. Polymers are widely used as drug carriers in the pharmaceutical industry. These substances are used not only to control the drug release but also to protect drugs from ambient moisture during storage or to prevent drug damage while passing through the digestive tract. Some of the main features of all polymers used in drug delivery are biocompatibility, good mechanical properties, and desirable pharmacokinetics [1]. The main purpose of the drug presence in DDSs is the acceleration of the therapeutic process. The choice of drug depends on many factors such as its role in the improvement of the disease and even the type of carrier. Why is it that drugs with special properties/applications are prescribed in controlled release of DDSs? Of course, it should be noted that drugs with a long or very short half-life, high therapeutic dose necessity, and a low solubility cannot be delivered with such types of DDSs. So, additives (as of release accelerators or retardants) may help these systems to achieve an optimal drug release. To provide the desired release kinetic in designing a controlled DDS, there are many factors involved, which can be generally divided into two operational phases: First one is the factors affecting at drug loading stage, and the other is about those determinants that directly interfere at the drug release period. All of these factors must be qualitatively and quantitatively well known to be appropriately applied into the carriers. Also, the type of targeted tissue is another factor that determines what type of DDS and which release mechanism is acceptable [2]. One of the main applications of tissue engineering is the preparation of skin alternatives for the healing of various wound types that fall totally into the skin

scaffolds category. An efficient skin scaffold should exhibit good mechanical properties, provide a suitable chemical nano/microstructures, and also pledge a proper surface to facilitate cell attachment, cell proliferation, and cellular differentiation. Besides, some chronic skin wounds require a controlled release of specific drugs at the wound bed, so the ability of loading drug within the skin scaffolds is crucial and ultimately leads to the formation of transdermal DDSs (TDDSs) [3–5]. In this chapter, the basic principles of the physiological structure of the skin, the process of spontaneous healing, and the types of wounds are discussed first. The following provides brief information and requirements for wound dressings as drug carriers in TDDSs. Also, the drugs which are commonly used in these systems are classified and studied in terms of their nature. Then, according to thermodynamic principles, affecting factors on the drug release are investigated in both loading and releasing phases. These thermodynamical factors are shown to be linked to the release kinetics. In the final step of this chapter, parallel to the noted basics, different TDDSs get introduced as well as their properties and applications are discussed and compared.

2. SKIN STRUCTURE AND WOUNDS Skin is the largest organ in the human body that naturally has a protective role against external hazards such as bacterial invasion or minor mechanical damage. The outermost layer of the skin, epidermis, consists of dead cells and keratinocytes that are separated from the dermis layer and transported to the surface. The epidermis has high impermeability that controls the rate of body water loss. The role of this layer is to provide physical integrity and flexibility of the skin as well as supporting vessels,

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lymph nodes, and nerve bundles. Some scientific sources suggest another layer to the structure of the skin called hypoderm and is often composed of fat. It is located underneath the dermis and causes the skin to be connected to the underlying tissues (e.g., muscle) (Fig. 1) [6]. A wound is called a defect in the integrity of the skin structure caused by various factors such as mechanical stress, surgical procedure, and burns. Losing integrity due to wounding can impair the function of the skin

and even in the worst state, cause death. The process of normal wound healing that performs spontaneously, is divided into four overlapping stages: hemostasis, inflammation, proliferation, and remodeling (Fig. 2) [7, 8]. The body’s immediate response to injury is hemostasis, which occurs at the site of injury to stop bleeding. At this stage, blood vessels contract, and also platelets and inflammatory cells (due to histamine secretion) accumulate in the wound area by binding to collagen

FIG. 1 Schematic of skin layers.

FIG. 2 Schematic of different stages of a normal wound-healing process.

CHAPTER 4 in the extracellular matrix (ECM). The duration of the inflammation phase is approximately 3 days. Neutrophils are the dominant cells at this stage that activate within 24 h after injury. The main function of neutrophils is the elimination of pathogens, foreign substances, and dead cells through the phagocytosis process. At the end of this step, macrophages phagocytose the neutrophils. The decrease in the number of inflammatory cells in the wound bed indicates the onset of the proliferation stage. At this stage, the proliferation of fibroblast and vascular endothelial cells and simultaneously, collagen deposition are performed to replace the temporary fibrin matrix at the site of the wound. These processes allow angiogenesis, tissue granulation, reepithelialization, and wound contraction to take place. The proliferation phase begins about 2–3 days after getting injured and continues until complete wound closure. The last stage of wound healing is called remodeling. This stage begins 2 weeks after the injury occurrence and can last for more than 1 year. During this stage, all of the activities related to inflammation and proliferation stages are ended, as well as the extra endothelial cells, macrophages and myofibroblasts are excreted through apoptosis. In general, cell migration into the ECM, remodeling and then the destruction of temporary ECM by matrix metalloproteinases (MMPs) are critical points in wound healing [9]. Wounds can be categorized from different perspectives such as origin, depth, length of the recovery period, and so on. A summary of the wound classification is shown in Fig. 3. One of the most important classifications is based on healing duration, divided into acute and chronic wounds [10]. Wounds caused by traumatic accidents or surgeries that lead to healing process are called acute wounds. If any factor prevents or lengthens the healing process, it may be the cause of potential risks, such as infection in the wound area. Such a wound can lower the body’s immunity and will be prone to become chronic. The wounds that do not heal within 90 days are called chronic. Common ways to treat these wounds are very difficult and expensive because of the possibility for infection prevalence. Such wounds often require surgery and sometimes lead to amputation or even death. Pathophysiologically, in chronic wounds the inflammation stage of the healing process becomes very prolonged and cannot transmit to the proliferation stage. Also, the other factors that cause a wound to be chronic include poor blood supply, persistent microorganisms and infection, disturbance in macrophage activity, an inordinate increase of MMPs level, lack of growth factors, and the presence of senescent cells at the site of injury [11]. With these all, given the incidence and progression of chronic wounds, the need for a proper, effective, and

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reliable treatment that can help the completion of the healing process is going to be critical.

3. WOUND DRESSINGS Various traditional methods such as allografting have been used to improve the healing process in chronic wounds, which most of them have been plagued by limitations. Thus the focus became directed toward new alternative approaches. This made the efforts to be centralized on developing the advanced wound dressings, which are consistent with wound pathophysiology, have an efficient influence on the wound-healing process, and would reduce the likelihood of side effects, secondary infections, bacterial attacks and at the same time have adequate physical and mechanical properties to protect the wound bed. The capability of loading drugs and other essential additives (growth factors, antibacterial or reinforcement agents, etc.) into these dressings would be a vital feature that has led to the formation of TDDS. In such systems, the tendency is to reduce the drug side effects and to manipulate their release profiles in an appropriate order. On the other hand, chronic wounds are covered with a layer of abnormal tissue, which separates the surface of injured tissue from the outside environment. Therefore the drug must first pass through the dead tissue to reach the live cells. This causes a significant amount of drug to stay inactivated and the efficacy to be decreased. High level of exudate production in chronic wounds is also another factor that can reduce the rate of penetration in such a local drug delivery. Therefore it is ideal to develop special carriers that can deliver an adequate dose of a drug to the target tissue without causing secondary injury in the wound bed [12, 13]. Advanced dressings for this purpose are mostly based on natural/synthetic polymers with various forms such as hydrogels, hydrocolloids, electrospun nanofibers, micro/particles, microneedles, foams, and films. Each of these dressings is applied to a particular type of wound that some of them will be discussed further. However, since the main focus of the chapter is on TDDSs, it is necessary to first study the drug types and also thermodynamical/kinetical aspects of drug–polymer interactions in drug release systems [14–16].

4. DIFFERENT DRUG TYPES Until now, various drugs have been used for DDSs. They can be divided into hydrophilic and hydrophobic groups. In some sources, metal-based nanoparticles have been also classified as drugs due to their antibacterial activity, antioxidant performance, and positive effects on healing process

FIG. 3 The wound type classification chart.

CHAPTER 4 [17]. Generally the ideal drugs for transdermal drug delivery should have several characteristics: I. partitioning coefficient (log P or υ) between 1 and 4 (excessive partition characteristic is not suitable for skin drug delivery); II. low melting temperature (below 200); III. minimum of shelf-life about 2 years; IV. short half-life (up to 10 h); V. no local irritation, sensitization and side effect; VI. a daily dose of less than 20 mg/day; VII. molecular weight less than 500 Da; VIII. permeability coefficient similar to skin (greater than 0.5
103 cm/h); and IX. the ability to release in both hydrophilic and lipophilic (a case of hydrophobic) phases [18, 19].

4.1. Hydrophilic Drugs In many DDSs, achieving a sustained release profile is considered a new challenge. In general the use of those small-molecule drugs (hydrophilic/hydrophobic) that are compatible with the polymeric matrix creates a preferential drug partitioning, which increases the encapsulation efficiency in the drug loading phase. However, it

Polymeric Transdermal Drug Delivery Systems

49

should be noted that such a parameter cannot guarantee a sustained drug release. The burst release for hydrophilic drugs is possible in two ways. In the first one, it is observed that hydrophilic small-molecule drugs have low solubility in nonpolar solvents/solutions, so the drug partitioning takes place on the surface of the polymer matrix during the loading phase, and thus it may be the cause of burst release in physiological fluid. In the second case, studies about water-soluble polymeric systems show that hydrophilic drugs such as ciprofloxacin, ampicillin, metronidazole, and cefazolin are readily miscible in the matrix and reciprocally are easily released into the body’s aqueous environment. Such a process also makes the system susceptible to burst release. However, studies based on the use of hydrophilic drugs in water-insoluble polymeric systems, with the assumption of no significant surface partitioning, indicate a sustained release. Therefore it is important to understand the types of drug-polymer interactions or, more generally, the physicochemical properties of drug release systems. Table 1 shows some of the hydrophilic drugs that have been used in nanofiber carriers so far, along with some of their characteristics [20].

TABLE 1

Some features of hydrophilic small-molecule drugs used in wound caring and their release information from electrospun mats. RELEASE DETAILS

Drug name

Log P

Polymeric matrix

Loading value (% wt./wt.)

1h

2h

Others

Ciprofloxacin

0.57

PVP PLCL/PDEGMA PVA/Alginate

0.4 10 –

– 12% 30%

– 20% 40%

60% (1 min) 80% (220 h) 85% (6 h)

Ampicillin

0.88

AL-BSA

PMMA/Nylon6 PCL PLLA

5 10 20 1–20 16.7 10

23% 17% 7% – 75% –

37% 25% 10% – 80% –

99% (96 h) 81% (96 h) 40% (96 h) 30% (6 h) and 50% (12 days) 98% (24 h) 98% (48 h)

Captopril

1.02

PLGA PLCL

10 10

– –

– –

100% (48 h) 78% (48 h)

Metronidazole

0.15

PCL Chitosan/PEO

4.8–14.4 1 5 15

20% 52% 70% 70%

40% 75% 80% 100%

90% (24 h) – – –

Cefazolin

0.4

Chitosan/PEO Gelatin

1 10

– 10%

20% 30%

65% (24 h) 95% (17 h)

AL-BSA, Amyloid-like bovine serum albumin; PCL, poly(ε-caprolactone); PDEGMA, poly(di(ethylene glycol) methyl ether methacrylate); PEO, poly(ethylene oxide); PLCL, poly(lactic-co-ε-caprolactone); PLLA, poly(L-lactic acid); PLGA, poly(lactic-co-glycolic acid); PMMA, poly(methyl methacrylate); PVA, polyvinyl alcohol; PVP, polyvinylpyrrolidone. Reproduced from M. Gizaw, J. Thompson, A. Faglie, S.Y. Lee, P. Neuenschwander, S.F. Chou, Electrospun fibers as a dressing material for drug and biological agent delivery in wound healing applications, Bioengineering 5(1) (2018) 1–28, with permission.

50

Modeling and Control of Drug Delivery Systems

4.2. Hydrophobic Drugs When compared with hydrophilic drugs, hydrophobic ones can generally produce sustained release profiles over a long period of time. The reason is the poor solubility of the drug in physiological body fluid that makes a preferred partitioning in a water-insoluble polymeric matrix rather than diffusion into the fluidic release media. Phenytoin, nifedipine, ketoprofen, vancomycin, and curcumin are several hydrophobic drugs which slow release profiles have been reported for them. Table 2 mentions some of the hydrophobic drugs that have been loaded in nanofiber carriers up to now. It is important to note that depending on the required time interval for the release process, a large variety of components and composition ratios are commonly used. The antiinflammatory drugs and coagulation factors are most proper for the early stage of healing and require a rapid release. Water-soluble polymers (dissolution mechanism) and systems with minimal polymer-drug interactions are often used for this class. However, proliferation and remodeling phases are counted as late stages of healing and require sustained drug delivery. In this case, blending, drug encapsulation in a polymeric matrix, and other methods that

provide a high value of polymer-drug interactions are frequently utilized [20].

4.3. Metal-Based Nanoparticles Nanotechnology and the use of nanoparticles in modern biomedical fields are rapidly expanding and coupling together. Therefore the application of some metal-based nanoparticles has been noticed recently due to their unique properties such as low cytotoxicity, bacteriostatic, and bactericidal activities. It is mentioned in the literature that some of these nanoparticles can accelerate the wound-healing process [21]. • Silver nano/particles and its derivatives Silver (Ag) and its derivatives (such as silver nitrate (AgNO3)) are known as an antibacterial agent and are commonly used to treat burns and infections and to promote wound healing. Ag nanoparticles (AgNPs) exhibit greater antibacterial and antifungal properties due to increase in surface-to-volume ratio and are therefore used in advanced wound-dressing design. It has been recently observed that with the presence of AgNPs in dressing, a significant decrease in inflammatory cytokines and oxidative stress occurs and the process of tissue regenerating is accelerated. Zhou

TABLE 2

Some features of hydrophobic small-molecule drugs used in wound caring and their release information from electrospun mats. RELEASE DETAILS

Drug name

Log P

Polymeric matrix

Loading value (% wt./wt.)

1h

2h

others

Curcumin

3.62

PHBV PCL/GT

3 4.7 3

55% 65% –

65% 67% –

70% (5 h) 78% (5 h) 65% (20 days)

Ketoprofen

3.29

PCL/Gelatin PVA PNVCL-co-MMA Cellulose acetate

5 5 20 15

– 50% 5% 10%

– – – –

40% (20 h) and 80% (45 h) 62% (48 h) 35% (24 h) 60% (48 h)

Nifedipine

2.49

Eudragit PU PNIPAAm/PU PVA

10 4.2 12 2

40% 15% 8% 27%

50% – 10% 29%

70% (8 h) 75% (72 h) 23% (30 h) 88% (48 h)

Phenytoin

2.26

PCL PVA/PCL

2 2

5% 11%

8% 15%

16% (48 h) 47% (48 h)

Vancomycin

1.11

Alginate

10

10%

–

60% (48 h)

Methylene blue

3.61

PHB/PEG

–

32%

–

90% (7 days)

GT, gum tragacanth; PEG, polyethylene glycol; PHB, poly(R-3-hydroxybutyrate); PHBV, poly(3-hydroxybutyric acid-co-3-hydroxyvaleric acid); PNIPAAm, poly(N-isopropylacrylamide); PNVCL-co-MAA, poly(N-vinylcaprolactam-co-methacrylic acid); PU, polyurethane. Reproduced from M. Gizaw, J. Thompson, A. Faglie, S.Y. Lee, P. Neuenschwander, S.F. Chou, Electrospun fibers as a dressing material for drug and biological agent delivery in wound healing applications, Bioengineering 5(1) (2018) 1–28, with permission.

CHAPTER 4 et al. synthesized silver chloride and AgNPs with reduced graphene oxide (Ag/AgCl/rGO nanoparticles), which allowed the release of silver ions under appropriate environmental conditions. Application of these Ag/AgCl/rGO nanoparticles in burn wound of mice, accelerated wound healing and wound closure, and also increased re-epithelialization and deposition of collagen fibers [22]. AgNPs mechanism of action: Initially, AgNPs attach to the cell membrane of bacteria and penetrate them. There, they interact with DNA and proteins containing sulfur and phosphorus and then release silver ions. Subsequently the bacterial respiratory system is targeted and its cell division will be disrupted that ultimately lead to bacterial cell death [23]. • Gold nanoparticles Gold nanoparticles (AuNPs) are biocompatible and widely used in drug delivery and wound-healing systems. Unlike silver, AuNPs do not exhibit antimicrobial activity as a single substance. Therefore they must be combined with other biomolecules to be optimally used in medical applications. Collagen, gelatin, and chitosan can be easily combined with AuNPs that results in enhancement of the wound-healing process. AuNPs mechanism of action: AuNPs antibacterial activity occurs in two ways: (1) when AuNPs penetrate inside the bacterial cells, they reduce the ATP levels there, resulting in damage to metabolism and ultimately make the bacterial cell death. (2) AuNPs with ROS-independent mechanisms cause cell death in bacteria [21]. • Zinc oxide (ZnO) nanoparticles Zinc oxide (ZnO) is an inorganic antibacterial agent that can stay stable when it is in contact with organic matters. Zinc, with long life in body cells, has been extremely successful in healing wounds especially burns. The topical application of zinc is inflammation reduction, improving re-epithelialization, and decreasing bacterial growth in chronic wounds. Besides, zinc acts as a regulator for keratinocyte migration and phagocytosis, and plays an important role in remodeling ECM. The ZnO mechanism of action is based on the formation of ROS species and its effect on wound healing depends on the light conditions, size, and concentration of the nanoparticles. Similar to AgNPs, these nanoparticles have a high antibacterial activity due to their small size and high surface to volume ratio [24]. However, Raguvaran et al. showed that high concentrations of ZnO nanoparticles are toxic to body cells. A similar result was previously observed by Shalumon and coworkers [25]. Thus it is generally recommended that the content of the ZnO nanoparticles used in the

Polymeric Transdermal Drug Delivery Systems

51

dressings should be sufficient to provide antibacterial properties and at the same time should be in a biocompatibility window that guarantees no cytotoxicity occurrence in the wound site. Table 3 shows a brief explanation for common metal-based nanoparticles in different types of wound dressings [21].

5. THERMODYNAMIC/KINETIC OF DRUG DELIVERY SYSTEMS Based on the release types, DDSs are divided into three categories: physical, chemical, and biological systems, each having subcategories as follows: I. Physical DDSs: • Diffusion-controlled systems: (related to the systems that diffusion controls the release rate) ✓ Matrix systems: dissolved, dispersed, porous, hydrogel, and bioerodible systems ✓ Reservoir systems • Ion exchange systems • Osmotically controlled systems • Hydrodynamically balanced systems II. Chemical DDSs: • Immobilization of drug mechanisms • Prodrugs III. Biological systems • Gene therapy In “diffusion-control systems,” the diffusion phenomenon can be studied in two aspects: (1) the first one is related to the design/manufacture phase of the DDS (before release). At this stage, drug penetration occurs across the polymer chains. Therefore the interactions between the polymer and the drug can be a determining parameter. In more detail the type of interaction, penetration, and distribution/dispersion of the drug in the polymer chains have a direct effect on the next stage (release stage). (2) The release stage involves the placement of DDS in the vicinity of physiological fluid (in vitro or in vivo condition). Through this process the drug molecules penetrate the fluid. Thus a three-component interaction between polymer-drug-physiological fluid plays an essential role in this section [26].

5.1. Parameters Affecting the Drug Release Factors that may influence drug release profiles rely on the properties of scaffold, the features of the drug, the interaction between these two components, and ultimately the environmental conditions. More precisely, the effective factors include carrier microstructure (polarity, crystallinity, glass transition temperature, free volumes, crosslinking density (CLD), etc.), carrier geometry (size, shape, and porosity), carrier material composition (polymeric

52

Modeling and Control of Drug Delivery Systems

TABLE 3

An overview for common metal-based nanoparticles in wound-caring applications. Nanoparticle name

Biomedical activities

Remarks

AuNPs

• •

Antibacterial activity Increasing the absorption of nanoparticles in the skin Promoting the cell proliferation Preventing the cancer progression and metastasis

Their mechanisms of action are as follows: (1) penetrating inside the bacterial cells and reducing the ATP levels; (2) performing ROS-independent mechanisms. Their activities mostly depend on roughness and particle size

• • AgNPs

• • • • •

Antibacterial activity Antiinflammatory effects Wound-healing acceleration Cosmetic attributes improvement Preventing the skin carcinogenesis

Their mechanism of action is penetration inside the bacterial cells that results in DNA damages and disrupting their cell duplication. Their activities mostly depend on particle size and particle shape

ZnO NPs

• • • • • • •

Antibacterial activity Antielastase function Antikeratinase activity Antioxidant application Antiinflammatory effects No cytotoxicity in low concentrations Theranostic properties in breast cancer

Their dominant mechanism of action is ROS formation Their activities mostly depend on size and concentration

TiO2 NPs

• • • •

Antibacterial activity Pronouncing high stability Suitable photocatalytic activity Effective antifungal activity against fluconazole-resistant strains

Their dominant mechanism of action is ROS formation. Their activities mostly depend on crystal structure, particle size, and particle shape

MgO and CuO NPs

• • •

Antibacterial activity Pronouncing high stability With low cost and easy availability

Their mechanism of action is the bacterial cell membrane damage and causing the leakage of intracellular contents and eventually death of bacteria

CeO2 NPs

• • • • •

Antibacterial activity Antioxidant application Wound-healing acceleration Having angiogenic properties Decreasing the reactive oxygen levels

–

CuO, Copper (II) oxide; CeO2, cerium (IV) oxide; MgO, magnesium oxide; NO, nitric oxide; TiO2, titanium dioxide.

components and additives), carrier degradability, the distance of encapsulated drug to carrier surface, drugpolymer diffusivity, drug solubility in polymeric matrix/ solution, drug partitioning coefficient, drug loading value, and environmental conditions (such as electrical charge, pH of polymeric solution, or physiological fluid, temperature, and pressure) [27]. Some of these items will be argued with more details: • Free volumes The diffusion into polymer chains occurs through the amorphous regions. Therefore diffusivity depends on the mobility of the polymer chains or in better speech, on the available free volumes of the system. Among the effective factors on free volumes content,

the most important one is the difference between the system temperature and its glass transition temperature (Tg). If the temperature of system is lower or close to Tg, the available free volumes will be reduced that makes the diffusion be also decreased (Fig. 4) [28]. Besides, the experimental studies about porous scaffolds show that drug penetration can be carried out either by drug continuous motions or by its discontinuous jumps between the polymer chains. The criterion for distinguishing these two penetration types is the ratio of the penetrant (drug molecule) size to the average pore size in the polymeric scaffold molecule size (γ ¼ drug average pore size ). For ratios smaller than one (γ < 1), penetration occurs through jumping of the

CHAPTER 4

Polymeric Transdermal Drug Delivery Systems

Glassy

53

Volume

Rubbery

Hole free volume

Free volume

Interstitial free volume

Occupied volume

Tg Temperature FIG. 4 The specific volume versus temperature typical graph for polymers [28].

drug molecules from pore to pore and few polymer segments get involved in drug penetration. The ratios greater than one (γ > 1) means that the drug motion into smaller pores is not possible and inevitably happens through polymer segments. The quantitative study of the effect of free volume on drug penetration process is related to thermodynamics of mixing and concepts of drug partitioning in physiological fluid [29]. • Polymer crystallinity and crosslinking density In semicrystalline polymer chains, as the crystalline domains increase, the dense regions get extended and, as a result, the drug movements face a high level of tangles, which reduce permeation. Increasing the CLDs also restricts the mobility of polymer segments, thereby reducing diffusivity. The importance of these two factors gets highlighted in controlling the drug release [26, 30]. • The polarity of polymer chains Polymer modifications (changing the polarity) by grafting or copolymerization with hydrophilic monomers increase the rate of drug release through the polymer chains. Inversely, combining hydrophobic functional groups to the primary polymer chains may provide a controlled release condition [31]. • Presence of fillers and plasticizers Adding plasticizers to the polymeric matrix increases permeability. The low concentration of these substances increases the free volumes because of decreasing the interchain interactions. This phenomenon enhances the mobility of polymer chains. However, adding fillers to the matrix decreases diffusivity due to an increase in tortuosity of the drug

penetration pathway. It is also important to note that fillers with smaller size, can cause the mentioned effect at lower concentrations. • Drug partition coefficient between the fluid and polymer In studying the drug diffusivity, drug distribution between the polymer and the physiological fluid is of particular importance. The drug distribution is determined by balancing the drug activity within existent phases (Eq. 1): αdp ¼ αdf

(1)

where αdp and αdf are the drug activity in the polymer phase and in its release medium, respectively. This equation can be expressed by substituting the activity coefficient (ν) and concentration (c) as Eq. (2) or (3): ðυcÞdp ¼ ðυcÞdf

(2)

cdp υdf ¼ ≡K cdf υdp

(3)

K is the partition coefficient, the same as ratio of drug solubility in the polymer to its solubility in the physiological fluid. The partition coefficient of a drug affects the drug release rate, and simultaneously depends on the drug solubility value in release medium [32]. • Drug solubility in polymeric matrix/solution In a polymeric matrix/solution, the rate of drug transfer within polymer chains is proportional to the slope of “drug concentration graph versus time.” This slope depends on the drug solubility in free volumes of amorphous domains. It is why that improving the drug solubility in the “matrix DDSs” (water-insoluble ones) reduces the risk of burst release.

54

Modeling and Control of Drug Delivery Systems

In dissolved/dispersed systems (subsets of matrix systems), drug molecules must leave their primary crystalline structure to be dispersed in the polymeric matrix. The stage of drug dispersion is called dissociation, which depends on the drug lattice energy and temperature. If the free energy of mixing gets greater than or equal to the drug lattice energy, dissociation will happen. The next step, namely, the dissolution of drug molecule into the polymer, is more complex and depends on the chemical nature of the polymer and the temperature of the system. Generally, those dispersed compounds can be dissolved in a polymeric matrix that have similar polarities and are in undersaturation condition. Eq. (4) states the relationship between solubility and melting temperature of the drug:
 4Hm ln Cp ¼ K  RT m

(4)

where Cp , 4 Hm, and Tm are mole fraction solubility, enthalpy of melting, and melting temperature of drug crystals, respectively. R is also the universal gas constant. This equation illustrates that the melting temperature of the drug, its melting enthalpy, and the inherent polymer/drug properties are factors that influence drug solubility in polymer [32]. Further quantitative details will be discussed in the next subsection. • Drug loading Drug loading is the placement of a specific drug value in a carrier. The “matrix systems” containing low amounts of a drug (undersaturation condition) are in fact, “dissolution systems.” Their release rate is initially proportional to the square root of time (t1/2), and after approximately 60% of release, the rate drops exponentially. For a high initial amounts of loading, the dispersed system is obtained. This makes a saturation that its release rate is still proportional to the square root of time. The first and second rules of Fickian diffusion are used to express the release rate in these systems. Finally, if the initial amounts of the drug exceed the saturation state, solid particles will appear in the matrix. After the initial drug release in such systems, some pores and continuous ducts will be replaced instead of particles. Later the pores will be filled by physiological fluid and then the release will take place through the formed channels [33]. • Environmental factors As mentioned earlier, environmental factors (electrical charge, pH of the polymeric solution or physiological fluid, temperature and pressure, etc.) also influence the way and speed of drug release. Some

of these factors will be discussed with details in Chapter 6 for drug-containing hydrogels. Although the effect of noted factors has been extensively investigated in experimental researches, most of the results are case-based and qualitative. Thus through using thermodynamic concepts, the researchers attempt to correlate the factors quantitatively. Final approaches may be promising to control the type of drug release and to regulate the miscibility of components in DDSs. Following is an overview of drug solubility in a polymeric matrix (one of the most important parameters affecting drug release kinetics) based on quantitative thermodynamic principles.

5.2. Principles of Drug-Polymer Solubility Based on Thermodynamics of Mixing In general, predicting the compatibility of two substances at a particular temperature/pressure depends on two issues: (1) the sign and magnitude of the free energy of mixing, and (2) the dependence of free energy of mixing on the components mixing ratio. Several theories have been developed to explain the change in free energy when two species are mixed. The Flory-Huggins (FH) theory predicts that for polymer-small molecules (drug) systems, the entropy of mixing is always favorable and should be relatively constant. Given the favorable entropy of mixing, the criterion for the miscibility will be switched on enthalpic interactions. Therefore it is important to calculate the FH interaction parameter (χ), which is an enthalpic term. The negative interaction parameter represents a system that exhibits strong adhesive interactions leading to miscibility. However, the positive interaction parameter indicates a system that has stronger cohesive interactions. Therefore to understand the miscibility in multicomponent mixtures, the focus should be on the relative balance between adhesive and cohesive interactions. Until now, different methods have been used to estimate the FH interaction parameter in drug-polymer systems [34]. In this section, some methods are investigated and compared which most researchers have performed their related studies based on them.

5.2.1. The primary solubility parameter approach The differences of solubility parameters are mentioned as a way to predict the miscibility of mixing in systems. Thus through the solubility parameters, χ parameter is estimated as follows (Eq. 5): χ¼

2 Vsite
δdrug  δpolymer RT

(5)

CHAPTER 4 where Vsite is a hypothetical lattice volume and is equal to the molecular volume of the drug, and (δdrug  δpolymer) is the difference of solubility parameters for drug and polymer. Most of predictions based on mentioned equation conflict with experimental results. This is because the above equation only considers van der Waals interactions, while the interactions between many drugpolymer systems are of other enthalpic species such as hydrogen bonding [35]. This has led the researchers to look for more accurate approaches.

5.2.2. The melting point depression method According to thermodynamic of mixing, through dissolution of a crystalline drug in an amorphous polymer, the melting temperature of the drug will be depressed. Parallel to the FH model, the magnitude of melting point depression can be related to the solubility of the crystalline drug in the polymer (Eq. 6):   pure      ΔHM 1 1 1  1  ϕdrug + 1   ¼ ln ϕ pure drug mix λ R TM TM  2 + χ 1  ϕdrug polymer MW =ρpolymer that; λ ¼ drug MW =ρdrug

(6)

where ΔHpure and Tpure are the fusion enthalpy and the M M melting temperature of the pure drug, respectively. R is the universal constant of gases, λ denotes the molar volume ratio of polymer, and drug (ρdrug and ρpolymer are the densities of drug and polymer, respectively), χ is the FH interaction parameter, and T expresses the onset of melting absolute temperature at the desired volume

Polymeric Transdermal Drug Delivery Systems

55

fraction of the drug (ϕdrug). To obtain the solubility of the drug in the polymer at room temperature, the melting temperatures should be determined in the various volume fractions of the drug, and then be fitted with the earlier mentioned equation. It should be eventually extrapolated to 25°C. It is noteworthy that the melting temperature of drug-polymer mixtures can be obtained by the differential scanning calorimetry (DSC) test at a specified heating rate. The exact ϕdrug is determined by high-performance liquid chromatography (HPLC) analysis in the laboratory. However, in the drug-polymer (two-component) systems the relation of ϕpolymer ¼ (1  ϕdrug) can be replaced in Eq. (6) to rearrange it as Eq. (7). More detailed descriptions of the experimental protocol are mentioned in Marsac and coworkers research [36]. 

   R 1 ln ϕdrug + 1  ϕpolymer + χϕ2polymer ¼ ΔHfus λ TM 0 1 (7) wdrug =ρdrug @ A     that; ϕdrug ¼
 wdrug =ρdrug + 1  wdrug =ρpolymer 1

 mix

1



pure TM

where wdrug is the mass fraction of drug, and also ρdrug and ρpolymer are the densities of drug and polymer, respectively. By plotting this equation against ϕ2polymer, a graph can be obtained that is linear at low volume fractions of polymer and its slope represents the χ parameter. According to Fig. 5, Baghel plotted such an equation for calculating χ in the dipyridamole (DPM)-poly (vinylpyrrolidone) (PVP), and in the cinnarizine (CNZ)-poly(vinylpyrrolidone) (PVP) systems [37]. Any lack of linearity across the concentration range most

FIG. 5 The plot of calculation the interaction parameter for the (A) DPM-PVP and (B) CNZ-PVP systems based on Eq. (7). ((From S. Baghel, H. Cathcart, N.J. O’Reilly, Theoretical and experimental investigation of drugpolymer interaction and miscibility and its impact on drug supersaturation in aqueous medium, Eur. J. Pharm. Biopharm. 107 (2016) 16–31, with permission.))

56

Modeling and Control of Drug Delivery Systems

likely reflects the dependence of the interaction parameter to the combination ratio [38]. The occurrence of melting point depression depends on the physical interaction between the drug and the polymer. Therefore to use this method, there must be sufficient physical interaction between the components. The melting point depression method is favorable for those systems that the Tm of polymer is low. In this state, when the drug starts melting, the polymer is spending its terminal zone, and thus they can interact and form an equilibrium in the liquid phase. Therefore it is clear that as the melting point of the drug gets depressed closer to the Tg of polymer, a weaker physical interaction between the drug and polymer will happen so that no depression may take place ultimately. In simple words, if the polymer is not completely liquid-like at the Tm of the drug, their mixing would be unfavorable. The same results are applicable to the polymer blends. Finally the results show that the melting point depression method is in good agreement with experimental data, but it is important to note some point of limitations: 1. The melting point depression method estimates the interactions between the drug and the analogue and/or monomer in the liquid state, meaning it does not take into account the fundamental physicochemical differences between the monomers and polymers. It is evident that in comparison with the polymer chain (with covalent bonds of the monomers), the liquid unbonded monomers have relatively unlimited molecular motions that provide the interaction with drug molecules without steric effect. Therefore this method overestimates drug solubility in the polymers. 2. One of the underlying assumptions of this method is that “the interactions between the drug and the analogue (and/or monomer) in the liquid state are similar to those in the solid state.” In other words, the solvent should not affect the molecular structure of the drug/ polymer, and the interaction between them (e.g., through protonation or deprotonation mechanisms). 3. As this method requires a liquid analogue and/or polymer monomer, it cannot be applied to all of the polymers. 4. This method is based on DSC thermal measurements that are time-consuming because of the slow dissolution or crystallinity kinetics of the drug. The curve fitting of Eq. (7) is also performed at higher temperatures and then will be extrapolated to 25° C. For this reason, the final predictions are associated with a degree of uncertainty, which is correlated with some factors such as accuracy of measurements, the

magnitude of temperature extrapolation, and the validity of the underlying assumptions of the proposed model (e.g., the assumptions in FH theory) [36–39].

5.2.3. The polymer in solution method Knopp estimated the polymer-drug solubility at room temperature by introducing a new polymer in solution method. In this study, experimental information was obtained by the simple shake-flask method and HPLC quantification, and afterward analyzing was performed through linear regression, mentioned in Eq. (8): Xdrug ¼ a:Cpolymer + b

(8)

where Xdrug, Cpolymer, and b are the drug solubility in polymer solution, the polymer concentration in solvent, and the drug solubility in pure solvent, respectively. Both of “a” and “b” are the fitting parameters. With determining the graph slope (a) in linear regression, the solubility of drug-polymer in solid-state ^ solid (X drugpoymer ) will be obtained. Besides, because the solubility of the drug is uncertain even in a pure solvent, it is suggested to predict the estimation interval by Eq. (9). 

^ solid X drug



0

^ ¼X drug  t0:025, N2
ðσ a + 2σ b Þ
solid

pffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi 1 + 1=N

(9)

The terms σ a and σ b are the standard deviations of fitting parameters (a and b) that are assumed to be independent parameters. Also “t0.025, N2 is the 2.5% quantile in the t-distribution and N is the number of measurements.” According to Eq. (8), as the concentration of polymer increases, the solubility of the drug enhances, which may be a reflection of the interaction between the drug and the polymer. In other speech, it can be concluded that if the solvent is inert, the drug solubility in polymer will be a linear function of polymer concentration for each drug-polymer composition ratio. Knopp showed the linear regression of celecoxib (CCX) solubility versus PVP concentration, for ethanol (▲) and methanol (■) solvents at 25°C. The solid lines represent the best fitting and the dashed lines obtain from nonlinear prediction intervals (Fig. 6). Some advantages of this method can be listed as below: 1. It does not require advanced equipment or complex nonlinear data processing but rather is based on a simple shake-flask method and HPLC quantification that can be examined in most laboratories. 2. This method can rapidly winnow those polymers which are suitable for DDSs containing solid dispersions as drugs.

CHAPTER 4

Polymeric Transdermal Drug Delivery Systems

57

FIG. 6 The plot of increase in CCX solubility as a function of PVP concentration. ((From M.M. Knopp, et al.,

A promising new method to estimate drug-polymer solubility at room temperature, J. Pharm. Sci. 105(9) (2016) 2621–2624, with permission.))

3. The results predicted by this method are in good agreement with the experimental values and have more accuracy compared with the melting point depression method. It may be due to calculations at room temperature (without extrapolation to lower temperatures). Of course, this method is only applicable in the presence of inert solvents and those polymers that have perfect dissolution in the solvent (perfect dissolution >100 mg/mL) [39].

5.2.4. Calculation the activity coefficient of the drug based on FH lattice theory The solubility of a crystalline substance in a typical low molecular weight solvent is given in Eq. (10): 
 ΔGfus ΔHfus T ¼ 1 ln υdrug xdrug ¼ Tm RT RT ð ðT config 1 T 1 ΔC config P dT  ΔCP dT + RT Tm R Tm T

(10)

where υdrug is the drug activity coefficient and xdrug shows the molar fraction of the drug. ΔGfus is the change of free energy between the supercooled liquid of drug and its crystal (solid-state). ΔCconfig denotes the heat P capacity difference between the liquid and the crystal (at the Tg). T is also the temperature that is supposed for estimation.

The υdrug can be a criterion of solubility, or more accurately, reflects the nonidealities of the mixing. With assuming the polymer to act as a solvent, the above relationship can be considered to describe the solubility of crystalline drugs in the polymeric matrix. Based on this and upon the concept of the FH lattice theory, Eq. (11) calculates the υdrug of the drug for a drug-polymer system:  
 ϕdrug 1 ln υdrug ¼ ln + 1  ϕpolymer + χϕ2polymer λ xdrug

(11)

According to the earlier mentioned equation, the magnitude of υdrug depends on enthalpic-based interactions in the system, which is reverberated in χ parameter. On the other hand, the difference between the molecular sizes of the system components is a criterion of entropy that is shown by logarithmic terms in the noted equation. These terms also affect the υdrug value. If the molecular size of components is alike (e.g., small molecule drugs vs low molecular weight solvents), the entropy of mixing can be assumed ideal. With this assumption, those two logarithmic terms in Eq. (11) will be ignored. In such a drug-solvent system, υdrug only depends on the partial molal heat of mixing for the supercooled liquid solute with solvent, which χ reflects it. Based on Eq. (10), it is obvious that if the enthalpy of mixing or the χ is positive, the solubility will decrease. Such a system shows that the cohesive interactions of the drug are much higher than adhesive ones. In other

58

Modeling and Control of Drug Delivery Systems

words, those compounds (drugs) with high melting temperature or large heat of fusion (proportional to lattice energy) do not tend to form a significant interaction with the polymer. Thus these drugs will exhibit very little solubility in the matrix. Mutually a negative enthalpy of mixing or a negative χ parameter indicates that the adhesive interactions of the drug have higher values, which increases the solubility of the compound (drug). From Eqs. (8), (9), it is generally apparent that the solubility of a drug in polymeric matrices is a function of its melting temperature, the heat of fusion, and also the enthalpy of mixing (reverberated in the χ parameter). That is why the drugs with higher heat of fusion or melting temperature are less soluble in the polymer. This speech is in agreement with aforesaid explanations in the melting point depression method (Section 5.2.2). Marcas et al. observed similar results in comparing the solubility of nifedipine and felodipine in the PVP matrix. The felodipine has a lower heat of fusion (proportional to lattice energy), so its solubility is expected to be higher. In more detail the solubility parameter of felodipine with the ΔHfus ¼ 30.83 KJ/mol and Tm ¼ 414.75 K is 25.0 J1/2/cm3/2; while this parameter for nifedipine with the ΔHfus ¼ 39.89 KJ/mol and Tm ¼ 445.25 K is equal to 22.9 J1/2/cm3/2 [38]. Also, as Table 4 shows, Baghel et al. confirmed the aforementioned statement in a distinct study [37]. Remarkably, if the drug solubility content is poor, the probability of its supersaturating in the matrix increases. As a result, it tends to recrystallization that reduces the miscibility of the drug-polymer system. This may enhance the burst release in the “release phase” as mentioned earlier [38–40].

5.3. The Relationship Between the Thermodynamic and Kinetic Principles of Drug Release However, determining the thermodynamic concepts about drug-polymer miscibility and the activity

coefficient of the drug is essential for tracking the system in drug loading/releasing phase, but it is not sufficient for comprehensive analysis of DDSs. It is because these two phases depend on many parameters mentioned previously. So for well predicting the system, all the affecting parameters should be determined according to thermodynamic fundamentals. Finally, for better controlling and characterizing DDSs, the results should be correlated to kinetic principles of drug release in physiological fluid. For example, free volumes have a critical effect on the drug loading/releasing phase. Therefore, utilizing some thermodynamical approaches (including modified FH theories such as Ruzette-Mayes method [41]) may cover the effect of free volumes in drug-polymer systems. Also, the geometry/type of carrier, the type of targeted tissue, and the type of required release mechanism are some issues that have been further scrutinized as kinetics of release. Since the qualitative and quantitative bases of release mechanisms have been investigated in many previous studies, they are not discussed in this chapter anymore [41]. It is important to note that beside the exclusive mechanisms of release, adequate carriers according to their types, are essential for optimal drug delivery to different tissues. Because the focus of this chapter is on TDDSs, common types of carriers will be reviewed for determining their features and applications.

6. POLYMERIC TRANSDERMAL DRUG DELIVERY SYSTEMS Choosing the right dressing for the wounded skin is crucial and depends on the risk of infection occurrence, the type and size of the wound, the amount of exudate, and so on. In general, appropriate wound dressings should absorb excess exudates and at the same time, regulate the amount of moisture in the injured area. It should also prevent pathogens and secondary infections, and additionally be permeable to oxygen. Besides, the

TABLE 4

Typical useful information applied to calculate free energy of mixing and solubility. Materials name

MW (g/mol)

Density (g/cm3)

Molecular volume (cm3/mol)

ΔHfus (KJ/mol)

Tm (K)

Solubility parameter (J1/2/cm3/2)

DPM

504.63

1.40

360. 23

29.06

441.26

29.07

CNZ

368.51

1.13

326.12

38.33

394.40

21.05

PVP K30

40,000

1.23

32,520.33

–

–

25.02

CNZ, cinnarizine; DPM, dipyridamole. Reproduced from S. Baghel, H. Cathcart, N.J. O’Reilly, Theoretical and experimental investigation of drug-polymer interaction and miscibility and its impact on drug supersaturation in aqueous medium, Eur. J. Pharm. Biopharm. 107 (2016) 16–31, with permission.

CHAPTER 4 design should be on a way that the wound dressing can create an adequate contact to skin, that is, covering the wound surface completely and standing firm on the injured tissue. Another notable requirement is the simple removal of the dressing from the surface of the skin without any side damage. Medical gauzes are the most applicable products for wound dressing. Gauze is made up of woven or nonwoven fabrics based on natural or synthetic fibers such as cotton and polyester fibers. They can absorb exudate from wounds and keep the environment moist. However, they cannot be a good barrier against microorganisms. Therefore designing and fabricating engineered dressings with all the aforementioned capabilities is of interest to researchers today. In addition the ability of loading drugs in these dressings is one of the unique features that has made them attractive as a TDDS. Some advantages of a TDDS include the following: I. preventing the first pass of metabolism; II. preventing the gastrointestinal maladaptation; III. a predictable and long duration activity; IV. minimal side effects; V. possibility of using short half-life drugs; VI. improving the drug effectiveness and minimizing its fluctuation; VII. stability of plasma concentration against potent drugs; and VIII. ease of treatment and self-management capability; Some disadvantages of TDDS are also as follows: I. as mentioned earlier, many hydrophilic drugs tend to burst release and cannot penetrate the skin too slowly that lowers the drug efficacy; II. in some cases, skin patches may be the reason for problems like irritating, itching and swelling; III. the barrier function of the skin or the patients’ age may affect the function of the drug; IV. TDDSs have poor compatibility with ionic drugs in the release condition; V. a large dose of drugs (above 10 mg/day) are difficult to load; and VI. drugs with a too low or too high value of the partition coefficient cannot reach a systematic circulation [7, 42]. The following is a brief explanation of the biocompatible carriers that have been used for drug loading as TDDSs up to now.

6.1. Film-Based Dressings Film-based dressings are durable, comfortable, transparent, easy to improve, cost-effective, semipermeable against water vapor and oxygen and impermeable to liquids. These properties prevent exudate leakage and

Polymeric Transdermal Drug Delivery Systems

59

protect the skin against bacterial invasion. Thin-film dressings are often polyurethane-based, with favorable elasticity and transparency. The flexibility of such polymeric dressings allows for easy movement of body parts that are under dressing. However, this type of dressing is not suitable for mild exudative wounds and superficial pressure wounds because they require short-term dressing replacement to prevent bacteria growth. Therefore, film-based dressings are less suggested for chronic wounds [43, 44].

6.2. Hydrogels Hydrogels are three-dimensional (3D), hydrophilic, and polymeric networks that can absorb large amounts of water or exudate. Their tendency to water is due to the presence of hydrophilic functional groups such as alcohols, carboxylic acids, sulfonic acids in the polymer chain structures. These systems are mainly composed of single-component or multicomponent polymers, with a minimum water absorption capacity of 20%. Hydrogel-based dressings can strongly keep the wound bed moist, rehydrate eschar, and aid autolytic debridement. The versatility of hydrogel dressings depends on their ability to create covalent or noncovalent crosslinks that controls their swelling ratio and dimensional stability. Common hydrogel-based wound dressings are amorphous or film-like. They can be easily placed in the wound area and simply removed if necessary. The most important advantage of hydrogels is to be injectable and suitable for in situ formation, which enables them to be positioned correctly and reduces the pain during implantation surgery. Protein-based hydrogels have been widely cited in the literature for wound healing. This is because the collagen is one of the main constituents of natural ECM, so that is why the gelatin or collagen-based scaffolds can result in positive effects on wound healing. Another type of hydrogel dressing is sodium/calcium alginate extracted from seaweed. Alginate is a type of natural polysaccharide that has excellent water absorption. As a result, when it is applied to wounds, large amounts of exudates can be absorbed and at the same time, bacterial activity may be significantly limited. Therefore alginate dressings are very useful for wounds with high level of exudate secretion. Also, the calcium ions existing in these dressings enhance stability and interconnection of the physical structure and promote healing. Fibrin-based hydrogels improve wound healing with the development of vascularization and cell adhesion. However, the difficulty of controlling their mechanical properties, their slow mechanism of crosslinking, the risk of an immune response or the transmission of

60

Modeling and Control of Drug Delivery Systems

infectious diseases have limited the use of such scaffolds in the treatment of wounds. Despite the tremendous benefits of hydrogels, the major concern about this type of dressing is their impermeability to gas (oxygen) which limits their use against infection. Therefore hydrogel dressings are often used in conjunction with antibacterial compounds. Another significant limitation for hydrogels is their rapid loss of water (dehydration), which necessitates the use of hygroscopes in their structure [7, 45, 46].

6.3. Hydrocolloid-Based Dressings Hydrocolloid-based dressings are moisture-retentive dressings that contain gel-forming agents (such as sodium carboxymethylcellulose and gelatin) and are usually combined with foam and film backings to form absorbent, self-adhesive, and waterproof wafer. These dressings exhibit semipermeability to water and oxygen and, therefore, absorb exudates from the wound bed. However, these systems are not recommended for infectious wounds with the possibility of hypoxic injury and excessive moistening of wound beds [7].

6.4. Electrospun Nanofibers As mentioned, the ECM in the dermis layer is composed of collagen nanofibers, which results in structural integrity and firmness of the skin tissue. Therefore the manufacture of ECM-mimicking scaffolds can be promising for the healing of various wounds. Among the available methods for nanofiber fabrication, electrospinning is considered as an efficient, inexpensive, and easy technique. The electrospun scaffolds have a high surface to volume ratio, which increases the contact points of cells with the scaffold surface, thereby enhancing the interaction and interconnection between them. The presence of porosity between the produced nanofibers facilitates oxygen permeability and enables fluid absorption. In addition, the pores in nonwovens are small enough (typically 1–10 μm) to prevent bacteria from penetrating. On the other hand, the experimental results have shown that the presence of electrospun nanofibers in the wound area reduces necrosis, enhances vascularization and then accelerates the healing process (Fig. 7A–D). It is also possible to load bioactive molecules such as growth factors, nanoparticles, and drugs into or on the surface of the nanofibers, which is inspirational for drug delivery purposes. Remarkably, the nanofibrous scaffolds have shown that they can maintain the activity and release of drugs for a relatively long and controlled period. Nonetheless the commercial use of electrospun mats as TDDS is still in its infancy.

The electrospinning process is applicable to the most of natural and synthetic polymers. Many studies have investigated the presence and effect of electrospun collagen in several tissue engineering fields (skin, cornea, and the blood vessel system). Other natural polymers, including fibroin, silk, chitin, chitosan, fibrinogen, or their mixtures have been also prepared as wound dressings by electrospinning, which have shown favorable cell attachment, cell growth, and cell infiltration under in vitro condition. Among synthetic polymers, electrospun nanofibers based on polyurethanes, polycaprolactone, poly-L-lactide, poly(lactic-co-glycolic acid), and so on have been suggested for application in wound dressings. In some recent studies, drug addition to nanofibers has been investigated toward manufacturing controlled release systems for wound healing. Many antibacterial, antifungal or antibiotic agents have been used for this aim, which some of them were noted before (Tables 1 and 2). Before utilizing nonwovens under in vitro or in vivo conditions, it is usually necessary to modify them through crosslinking. This process aims to stabilize the nanofibrous structure of nonwovens (especially for watersoluble nanofibers). One of the most common methods is chemical crosslinking by glutaraldehyde vapor. In this step, despite the crosslinking between the nanofibers, the nanodimensions and the porous structure remain in the system. However, one of the major problems in electrospun mats is their pore size distribution (smaller than 10 μm), which significantly reduces cell penetration rate and cell growth in the thickness direction of the scaffold. Nonwovens are also among the 2D scaffolds, while 3D ones are more desirable for many biomedical applications. Therefore many efforts are being made now to develop new methods for producing 3D electrospun scaffolds with larger pore sizes [7, 46, 47].

6.5. Micro/Nanoparticles Biodegradable nanoparticles based on polysaccharides, polymers, and some minerals/metals are capable of loading and releasing bioactive molecules (such as growth factors, drugs, and proteins). They have been recently considered for the treatment of various wound types. The nanoparticle encapsulation (such as ZnO, TiO2, and AgNPs) within scaffolds has been recently investigated to create antibacterial properties and to prevent local infections and unorganized collagen deposition in wound area. One of the most facilitate ways to fabricate the drugloaded nanoparticles is simple/double emulsifying. In this method, two immiscible fluids get mixed, which makes some spherical droplets of first fluid be dispersed

CHAPTER 4

Polymeric Transdermal Drug Delivery Systems

61

FIG. 7 Nanofiber mats for healing the chronic wounds. (A) The SEM images of gelatin methacryloyl (GelMA)

nanofibers before and after 24 h incubation in PBS. (B, C) The effect of incorporation of GelMA nanofibers and gelatin constructs on the flap necrosis and vascularization. (D) Images of implanted electrospun nanofibrous mats and the adjacent skin flap. (d–f ) The color laser Doppler detection of skin flaps perfusion (7 days postsurgery), confirming the better therapeutic outcome of GelMA nanofiber. ((From S. Saghazadeh, et al., Drug delivery systems and materials for wound healing applications, Adv. Drug Deliv. Rev. 127 (2018) 138–166, with permission.))

inside the second one. This combination can be repeated with a third fluid to form double emulsions. The particle size in the emulsifying method can be controlled by varying the viscosity of the solutions and the mixing speed. Molding is another applicable method to produce polymeric nano/microparticles. However, the particle size in this method is generally larger than that in emulsifying method. The advantage of molding method is the monodispersity of particle size distribution, which provides a predictable kinetic of drug release. Though, one of its major limitations is controlling the particle shape. For example, it is difficult to fabricate spherical particles in this method. Another way for fabrication the drug-loaded nano/ microparticles is microfluidic platforming. The advantage of this method is the ability to make multicompartmental droplets. Droplet microfluidic systems are such as emulsifying systems, except that particles are formed by flow-induced shear stress. The main challenge in this

method is the low efficiency and limitation of particle size and geometry (Fig. 8) [7, 48].

6.6. Biofoams Biofoams have high porosity, interconnected pores, excellent fluid absorption properties, oxygen permeability, and suitable thermal isolation property. They are therefore desirable candidates to treat various types of wounds. Biofoams are capable of absorbing more exudates than thin-film dressings. Collagen-based biofoams are more widely used because of their good mechanical and physical properties, enhancing fluid uptake and re-epithelialization in wounds. However, the main disadvantages of conventional swine/bovine collagen are its rapid degradation and the risk of transgenic disease transmission to humans. Gelatin-containing biofoams, beside the elimination of collagen-related detriments, have good degradation profiles and favorable angiogenic properties. However, their porosity and water absorbance are lower

62

Modeling and Control of Drug Delivery Systems

FIG. 8 Approaches for fabrication of drug delivery tools. (A) The process of fabricating poly(lactic-co-glycolic

acid) (PLGA)-based drug carriers through double emulsion process. (B, C) Micrographs of the fabricated particles containing chlorhexidine (CHX) and platelet-derived growth factor (PDGF-BB). (D) The effect of dual drug delivery on wound-healing rate. (E) Droplet-based microfluidic platform for fabrication of porous microgels. (F) Fabrication of multicompartment drug carriers using micro fluidic systems. ((From S. Saghazadeh, et al., Drug delivery systems and materials for wound healing applications, Adv. Drug Deliv. Rev. 127 (2018) 138–166, with permission.))

than other natural hyaluronic acid, chitosan, or alginatebased hydrogels. Biofoams based on synthetic polymers such as polyesters have better elasticity and mechanical properties. In this category, polyurethane foams are frequently used which usually consist of two or three layers. The layer in contact with the wound surface is hydrophilic and the backing layer is hydrophobic which prevents leakage. In general, the porous structure in foam-based dressings is affordable and impressive for wound healing, but their too high water absorbance due to the large pores

reduces the effectiveness of drug release. Thus the hydrogels are more favored for producing the TDDSs with high-absorbance features. Another challenge with foam-based dressings is that they are usually nonadhesive and require backing or bandage. Eventually, these dressings are recommended for the early stage of healing in wounds with bleeding [18, 49].

6.7. Microneedles Microneedles are a group of TDDSs that increase the diffusivity of the drug to the stratum corneum by creating

CHAPTER 4 micron-size pores in the skin. These carriers do not stimulate the nerves and are therefore painless. Microneedles are also known as excellent drug protectors and gradual release systems. A multilayer structure of microneedles can be designed to deliver the drugs needed for each stage of wound healing at appropriate times. Microneedles can be classified into four groups: (1) Solid microneedles that allow drug penetration by creating pores in the epidermis. (2) Drug-coated microneedles that perform drug delivery from their outer surface while creating pores. (3) Drug container microneedles that are soluble in the physiological environment of the skin. After their penetration within the tissue and sojourning there, their degradation begins and simultaneously the encapsulated drug inside them gets released. (4) Hollow microneedles that penetrate the tissue and remain there. They facilitate drug transfer from the upper reservoir to the target site [7] (Fig. 9).

Polymeric Transdermal Drug Delivery Systems

63

7. CONCLUDING REMARKS In tissue engineering, the field of wound healing is growing rapidly and a large number of research groups are investigating different aspects of wound pathogenesis and providing optimal ways to accelerate the healing process. Therefore the development of skin scaffolds or wound dressings is of great importance. In particular the type of interaction between the drug and polymeric system has a direct effect on the way of drug loading and also its release mechanism under physiological conditions. So depending on time required for the release process, different compounds and different mixing ratios can be used. Further the processability and biocompatibility of many polymeric materials have made them useful for drug delivery in addition to dressing applications. In other words, they are the golden keys for producing TDDSs. For example, the release of antiinflammatory drugs

FIG. 9 Microneedle arrays as transdermal drug delivery tools in skin care. (A) Schematic showing different

types of microneedles used for transdermal drug delivery. (B) Bilayer microneedles with dissolvable tips carrying insulin before and after implantation. (C, D) Microneedles with swellable tips used for better adhesion of skin flaps to the surrounding tissues. (E, F) Swellable microneedles used as self-locking drug delivery tools. ((From S. Saghazadeh, et al., Drug delivery systems and materials for wound healing applications, Adv. Drug Deliv. Rev. 127 (2018) 138–166, with permission.))

64

Modeling and Control of Drug Delivery Systems

and coagulation factors are some agents for the early stages of healing and require a rapid release. For this category, water-soluble polymers (dissolution mechanism), and systems with minimal polymer-drug interaction, are suggested. However, proliferation and remodeling processes are among the late stages of healing and entail sustained drug delivery. In this case, blends, drug encapsulation processes in the polymeric matrix, and systems with a high level of polymer-drug interaction are often used. However, delving deeper into the interface between nanotechnology, polymer engineering, wound pathophysiology, tissue engineering, and DDSs will provide engineered systems with optimal properties and cost.

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CHAPTER 5

Stimuli-Responsive Polymers as Smart Drug Delivery Systems MEHDI JAHANBAKHSHI • MOHSEN SHAHROUSVAND

Caspian Faculty of Engineering, College of Engineering, University of Tehran, Rezvanshahr, Iran

1. INTRODUCTION As technology advances, biomaterials and polymer scientists are looking forward to provide intelligent drug carriers with environmental stimuli. Living systems in the body of organisms respond to different environmental changes. Polymer science experts try to mimic this behavior by natural and/or synthetic macromolecules, which is named as stimuli-responsive polymers, considering the physical, chemical, and biological changes [1]. Smart polymers have innovative applications in the biomedical field, such as delivery systems for therapeutic agents, tissue engineering scaffolds, cell culture supporters, biological separators, and sensors. In 1988 the University of Michigan researchers made an intelligent polymer using electrorheological fluids in which viscosity of this smart polymer changed abruptly with the slightest flow of electricity [2]. Currently, it is widely used for smart polymers in the field of biopharmaceuticals. So, this chapter discusses the manufacturing methods, applications, properties, and future of various types of these materials in drug delivery technology.

2. POLYMERS AS RESPONSIVE DRUG DELIVERY SYSTEMS Polymers are a group of giant molecules that are made up of a structural unit called a monomer. In a general classification, polymers can be divided into addition polymerization and condensation polymerization. Molecules having at least one unsaturated carbon-carbon bond in their structure can be known as a monomer in addition polymerization. If there is exactly one unsaturated band in the monomer structure, the final product of polymerization will be a linear polymer. If some of the monomers in the reactor have more than one carbon double bond in their structure, a cross-linked network will be generated. Addition polymerization has three stages: initiation, propagation, and termination that require initiator molecules for production of active

radicals. However, condensation polymerization, also called stepwise polymerization, requires monomers with at least two reactive functional groups [3]. If the monomer functionality is greater than two, the reaction product will be a crosslinked network. Therefore, a wide range of monomers can be produced by any of the polymerization methods to produce a variety of products capable of responding to stimulus. According to Cabane et al., Stimuli are divided into three categories: physical, chemical, and biological [4]. – Physical stimuli such as light, temperature, ultrasound, magnetism, mechanical and electrical stimuli, and pressure that often change the dynamic energy levels of chains. – Chemical stimuli such as solvent, ionic strength, and electrochemistry or pH that increase or decrease intermolecular interactions, for example, between polymer with solvent or between polymer chains. – Biological stimuli such as enzymes or receptors that affect the actual function of molecules, such as enzymatic reactions or the detection of molecule receptors. There are also polymers that respond to more than one stimulus at a time. External stimuli-sensitive polymers exhibit an instantaneous and reversible response due to changes in the stimulus factor, for example, its hydrophilic structure becomes hydrophobic or vice versa. In general, smart polymers, depending on their nature, due to the external stimulus can exhibit at least one of the following responses, as shown in Fig. 1 [5]: I. If the polymer is linear, it can be coiled and recoiled. II. If the polymer is crosslinked, it can swell or shrink. III. If the polymer is grafted to the surface, it can change the surface energy. Sensitive polymers can be homopolymer, random copolymer, diblock copolymer, multiblock copolymer, dendrimer, interpenetrating polymer networks (IPNs), and others. Fig. 2 shows some structures made by these polymers that have stimuli-responsive features.

Modeling and Control of Drug Delivery Systems. https://doi.org/10.1016/B978-0-12-821185-4.00015-4 © 2021 Elsevier Inc. All rights reserved.

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Modeling and Control of Drug Delivery Systems

The structures shown in Fig. 2 can give rise to coils, core-shell, polymersomes, and brush-like surfaces. Each of these structures has a different drug release.

In sum, the basis of the performance of an intelligent polymer consists of receiving, processing, and responding to stimuli based on the thermodynamics of the transitions.

2.1. Thermoresponsive Polymeric Drug Delivery Systems

FIG. 1 Classification of the responsive polymers by their

physical forms.

There are polymers that respond to temperature stimuli under different mechanisms. Temperature is one of the most commonly used pathological stimuli in drug delivery systems (DDS), which are divided into two categories: internal and external temperature stimuli [6]. Internal temperature changes are caused by tumors, inflammations, infections, and so on, but external stimuli are supplied by a heat source. Since the temperature range in the external method is higher and its efficiency and ease of use is higher, this method has been of much interest to researchers. The temperature stimulus actually affects the intermolecular forces and responds by forming and destroying them. The phase transition in polymer solutions was first studied by Flory and Huggins, which investigated the solvent-polymer and polymer-polymer interactions [7, 8]. This phase transition can be divided into four categories based on nature of interactions [9]:

FIG. 2 Various structures made by stimuli responsive polymers.

CHAPTER 5 Stimuli-Responsive Polymers as Smart Drug Delivery – – – –

van der Waals interactions Hydrophobic interactions Hydrogen bonds Ionic interactions The difference in the strength of the interactions and the response activation energy depends on the nature of the polymers. Providing activation energy of response is the origin of the difference in the phase behavior of polymers with respect to temperature changes. Accordingly, thermos-mediated DDS have been described based on the upper critical solution temperature (UCST) and the lower critical solution temperature (LCST) [10]. Most polymers exhibit LCST behavior. Although most polymers exhibit LCST behavior, in both models the transition from single-phase to two-phase is studied in a specific composition. As shown in Fig. 3, the polymer chain has the highest swelling in its singlephase region but shrinks if it enters its two-phase region. Therefore the drug is loaded when the polymer chains are in the single-phase region and when the chains are shrunken the drug exits the polymer. Therefore LCST behavior seems more appropriate and easier for applications of intelligent drug release. As shown in Fig. 3, the polymer solution can be single-phase or two-phase in a constant concentration of polymer at different temperatures. When the ambient temperature is lower than the LCST of polymer, the

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drugs are encapsulated in relaxed polymeric network in vitro. Once the thermal-sensitive drug carriers reach a tumor or inflammatory site, which have usually above normal body temperature, the polymer chains begin to shrink and release encapsulated drugs for therapeutic purposes. Advantages of UCST polymers can be described for the preparation of self-assembled structures that disassemble in environments with temperatures beyond UCST. Therefore the critical temperature of the phase transition for LCST/UCST-sensitive drug carrier must be within the body temperature range. The design of critical temperature-regulated structures depends on the knowledge of polymerization engineering, which can reach the desired temperature by copolymerizing hydrophilic and hydrophobic monomers with arbitrary sequences and lengths. The sol-gel phenomenon for temperature-responsive DDS is regulated by a semihydrophilic-to-lipophilic ratio on the polymer chain and is an energy-driven phenomenon that depends on the free energy of mixing or the enthalpy or entropy of the system [11]. So the temperature-responsive DDS can be divided into two types: negatively thermosensitive and positively thermosensitive, based on coil-to-helix conversions. Table 1 shows some of polymers with LCST or UCST behavior in the critical temperature region.

FIG. 3 The UCST and LCST behaviors of thermal responsive polymers.

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2.1.1. Poly(methyl vinyl ether)

TABLE 1

LCST-type and USCT-type polymers and their critical temperature.

The most interesting thing about this polymer is its transition temperature of 37°C that exactly corresponds to the body temperature. Poly(methyl vinyl ether) (PMVE) is a vinylic polymer with LCST behavior which can be copolymerized with other monomers [23].

Phase transition temperature in aqueous solution (°C)

References

PNIPAM

30–34

[10, 12]

Poly(N,Ndiethylacrylamide)

33

[13]

Poly(N-vinylcaprolactam)

30–50

[14, 15]

2.1.3. Polypeptides

PEO-b-PPO

20–85

[16]

25

[17–19]

mPEG-b-P(AAm-co-AN)

43.1

[20]

Poly(allylurea)

40–65

[21]

2,6Diaminopyridine-based polymers

40–65

[22]

Polypeptide is an amphiphilic polymer with folded structures that can exhibit temperature-sensitive behavior. Among these polymers are elastin-like structures which have LCST-type response to temperature [24]. He et al. prepared OEGylated polypeptides bearing Y-shaped pendants and studied the structure and thermoresponsive properties relationship [25]. They were demonstrated polypeptide bearing Y-shaped pendants showed significantly increased hydrophilicity than the linear-shaped counterpart. This complex shows a reversible LCST-type phase transition in the range of 37.5–42.5°C.

Polymer

2.1.2. Poly (N-ethyl oxazoline)

LCST type

UCST type PAAm/PAAc IPN

Although this polymer has a temperature-responsive behavior, its transition temperature is relatively high to be used as a DDS. Therefore various methods such as fabrication of copolymers and IPNs with other monomers are used to lower the transfer temperature.

2.1.4. Poly(N-vinylcaprolactam) Poly(N-vinylcaprolactam) is a nonionic, nontoxic, water-soluble, thermally sensitive, and biocompatible polymer which synthesized by free radical polymerization. Similar to PMVE and PNIPAM, this polymer has a transition temperature around body temperature (33°C), and that is why it has attracted the attention of many scholars.

2.1.5. Poly(N-isopropylacrylamide) FIG. 4 The Structure of some of the most used thermal-

responsive polymer in DDS.

Since the hydrophilic/hydrophobic balance of polymer chains is dependent on solvent interactions and the additives can alter the solvent’s ability, the effect of the additives on temperature-responsive behavior cannot be ignored. All additives such as salts, co-solvents, and surfactants can alter the quality of the solvent and can therefore alter the polymer-solvent interactions and critical temperature transitions [23]. Some of the most commonly used temperature-responsive polymers are discussed in this section and their structures are shown in Fig. 4.

Poly(N-isopropylacrylamide) (PNIPA, PNIPAAm, NIPA, PNIPAA, or PNIPAm) is a fast on-off switching temperature-responsive polymer that can be synthesized from N-isopropylacrylamide by free-radical polymerization. This polymer can be cross-linked by N,N0 -methylene-bis-acrylamide or N,N0 -cystamine-bisacrylamide (CBAm) [26]. This polymer is one of the most commonly used temperature-responsive polymers that, if its temperature goes above 32°C, this swollen hydrogel becomes shrunk and exhibits a 90% reduction in volume. It can then remove the loaded drug from its structure. However, there have been reports that the use of poly(N-isopropylacrylamide) is inappropriate due to its ammonium structure and nonbiodegradability and ability to react with blood platelets [27].

CHAPTER 5 Stimuli-Responsive Polymers as Smart Drug Delivery

2.1.6. Poly(acrylic acid-co-acrylamide) If the acrylic acid and acrylamide monomers were polymerized by radical polymerization simultaneously in the reactor, the product would be an IPN copolymer that would exhibit UCST thermal behavior. This means that the poly(acrylic acid-co-acrylamide) is singlephased at temperatures above 25°C and, if lowered, enters the two-phased region and shrinks. This feature can be used to load and release the drugs [28].

2.2. pH-Responsive Polymeric Drug Delivery Systems If for temperature-sensitive polymers, we divide the stimulus into two categories: internal and external, but here the major stimulants are internal. For example, the gastrointestinal tract has all three pHs: acidic, neutral, and alkaline. The stomach is acidic and the intestine is alkaline. It is also found that tumor-bearing areas and inflammatory wounds have lower pH levels than the healthy tissues, which provides an opportunity for the development of carriers of pH-sensitive smart drugs based on pH-responsive polymeric materials [29]. The variation in pH of different body tissues is shown in Table 2. pH-sensitive polymers are polyelectrolytes that have acid or base groups in their structure that either capture or release protons. The absorption or desorption of hydrogen is affected by changes in the pH of the medium, ionic strength, and type of counterions which depends on dissociation constant (Ka) of polyacids or polybases. Polyacids swell at a pH greater than the pKa of the polymer while polybases swell at a pH less than the pKa of the polymer. As the pH increases, the

TABLE 2

pH in various tissues and cellular compartments [29, 30]. Tissue/cellular compartment

pH

Blood

7.35–7.45

Stomach

1.0–3.0

Duodenum

4.8–8.2

Colon

7.0–7.5

Early endosome

6.0–6.5

Late endosome

5.0–6.0

Lysosome

4.5–5.0

Golgi

6.4

Tumor, extracellular

7.2–6.5

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polyacid becomes ionized and repulsion between the anions in the polymer structure causes hydrodynamic volume of the polymer to increase. Likewise, for a polybase at a low pH, it is ionized and its cationic structure repels and swells it. Figs. 5 and 6 illustrate the structure of some polyacids and polybases, respectively. But ionization is not the only method of pH-sensitive drug release. Changes in pH can affect the solubility and phase behavior of the polymers and thereby load and release the encapsulated drugs. Another approach is to degrade the polymer backbone at pHs of acidic or basic, causing the drug to release. It should be noted that in this method the pH changes should not be such as to cause drug damage. The use of hydrophobic monomers when synthesizing is another approach to control the swelling rate of pH-responsive polymers. When the polymer is in neutral conditions, the polymer chains are not ionized and there is no repulsion between the chains, therefore the intensity of the phase separation of the hydrophobic segments is amplified and the difference in osmotic pressure causes the drug to release from the polymer mass.

2.3. Biological-Responsive Polymeric Drug Delivery Systems Biologically responsive polymers are one of the smart DDS that are sensitive to internal biological stimuli which are inherently present in the natural system. These systems usually are conjugated natural or synthetic polymers with biological components of the body such as antigen-responsive polymers, glucose-sensitive polymers, and protein-responsive polymers. The advantage of these systems, in addition of being able to control biomarkers, is that they are biocompatible due to the biological components in their structure and the immune system is less motivated. Enzymes are biological catalysts that exclusively accelerate some biological reactions. For example, lipases hydrolyze lipid bonds. So, enzymatic smart DDS have sensitive component in their structure which in the face of its effective enzymes, it undergoes changes in its structure and releases its own drug. Therefore, hydrolases are the most widely used enzyme in intelligent drug delivery. For example, proteases, glycosidases, lipases, kinases, and phosphatases can trigger drug delivery when the carrier is stabilized by peptide, polysaccharide-based, lipid, and phosphate links [31]. Glucose-responsive polymers are other biological smart polymers that release the bioactive compounds for insulin therapy. These smart carriers can mimic the

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Modeling and Control of Drug Delivery Systems

FIG. 5 The structure of some polyacids usable in pH-responsive polymer systems. (From G. Kocak,

€tu €n, pH-Responsive polymers, Polym. Chem. 8(1) (2017) 144–176, with permission.) C. Tuncer, V. Bu

original endogenous insulin secretions which display variability in the inconsistency of glucose. Therefore these responsive polymers can be used in both glucose-sensing and insulin-delivery applications. Of

course, enzymatic hydrolysis reactions are required to carry out the mechanism of glucose or insulin release. One of the systems that have been increasingly effective in controlling insulin release is the sensitivity of

CHAPTER 5 Stimuli-Responsive Polymers as Smart Drug Delivery

FIG. 6 The structure of some polybases usable in pH-responsive polymer systems. (From G. Kocak,

€tu €n, pH-Responsive polymers, Polym. Chem. 8(1) (2017) 144–176, with permission.) C. Tuncer, V. Bu

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Modeling and Control of Drug Delivery Systems

poly(acrylic acid) carriers to gluconic acid-induced from glucose oxidation. Other approaches for insulin-modulated DDS are lectin-based carbohydrates and polymers with phenylboronic groups that release insulin when the glucose concentration increases. Mechanisms of insulin release in these carriers are based on enzymatic reaction and decrease of cross-link density in their network.

2.4. Ultrasound-Responsive Polymeric Drug Delivery Systems Based on physiological conditions, sometimes it is not possible to use thermal, magnetic, and other stimuli for the release of drugs. Ultrasound-responsive polymeric DDS (URDDS) are one of the novel smart carriers with enhanced permeation of drug through the implosion of these microbubbles [32, 33]. One of the most widely used areas of ultrasound waves is ultrasound imaging. Since the permeability and drug release of these systems occur in the presence of ultrasound waves at frequencies of 20 kHz or greater, they can be very effective in biocompatibility and their cellular uptake. This method is an external stimulant drug release strategy that depends on the intensity, frequency, and duration of use of ultrasound waves. This method is most commonly used in cancer chemotherapy. The URDDS carriers can be micro/nanobubbles, micro/nanodroplets, micelles, and emulsions (Fig. 7). The polymers used in these methods must be degraded by the waves so that their strength is less rigid than the tissue membrane. When drug carriers arrive at the target site that can be monitored by various methods, they are disrupted by the waves. URDDS are exciting topics in drug release systems that have recently received attention in the theranostics, tumor angiogenesis, transdermal drug delivery, and delivering targeted therapy [33].

FIG. 7 Drug carriers responsive to ultrasonic waves.

2.5. Electro-Responsive Polymeric Drug Delivery Systems The electric field-sensitive polymers that have ionizable groups in their structure change their features in response to a slight change in the electric current. As mentioned earlier, these polymers are also pH sensitive. These polymers based on their ability to convert electrical energy as a stimulus to the mechanical response used in various fields such as DDS, artificial muscle actuations, transductors, and sound dampers. When the electric field is applied to the electrolyte polymer, it ionizes, causing pH changes in the environment and releasing the drug according to the described mechanism. There are four stages to this process [27]: – diffusion – electrophoresis of charged drug – forced convection of drug out of the gel along – liberation of drug on erosion of electro-erodible polymers It is therefore predictable that parameters such as field strength, position of electro-responsive polymeric drug carriers in body, thickness and geometry of carriers, nature of polymers, and the applied voltage affect the response quality of these polymers. The polymers used in this approach are those polymers introduced in the pH-responsive section that can be natural, synthetic, or a combination of both such as chitosan, agar, alginate, hyalouronic acid, allylamine, vinyl alcohol, acrylonitrile, methacrylic acid, hydroxyethyl methacrylate, and acrylic acid [34–38].

2.6. Other Responsive Polymeric Drug Delivery Systems Recently, other polymers have also been introduced that have been less studied in their stimuli such as shear stress-responsive polymers, redox-responsive polymers, shape memory polymers (SMPs), and molecularly

CHAPTER 5 Stimuli-Responsive Polymers as Smart Drug Delivery imprinted polymers (MIPs). The use of shear stresses can affect the resistance behavior of polymers. For example, if a polymer is shear thinning, the drug is stripped of its structure due to stress. Shear thinning and thickening are the rheological properties of polymers. This sensitivity can also be provided by the aggregation of nanoparticles. Magnetic nanoparticles can coalesce or expand under the influence of an external magnetic field, and these changes cause drug release. Redox-responsive polymer drug carriers act based on oxidation-reduction responsive process such as polymers containing of disulfide and diselenide linkages [39]. These polymers were synthesized by living/controlled polymerization such as atom transfer radical polymerization and reversible addition-fragmentation chain transfer polymerization [3, 40, 41]. SMPs are capable of remembering the initial shape and returning it even after relatively high deformations [42]. These polymers initially have a specific structure in ambient conditions. They are then deformed in abnormal conditions and returned to their original shape when exposed to environmental conditions. This feature can be smartly used to load and release drugs [43]. The driving force of SMPs is their transition temperatures. MIPs are artificial and intelligent systems formed in the presence of a molecule as a template [44]. Pattern molecules are able to organize and link to the target molecule. The stability, ease of preparation, and low cost of the constituents make it possible to replace MIPs with antibodies or enzymes in catalytic, separable chemical sensors, and smart delivery systems [45]. These structures are based on a pattern that is capable of absorbing or releasing specific molecules that make their use unique. Another new approach to smart drug delivery is to combine two or more methods simultaneously, for example, pH and thermal-responsive carriers. It is logical that in such cases efficiency and quality will increase [46, 47].

3. FUTURE TREND AND CONCLUSION This chapter attempts to have an analytical look at the fabrication and function of polymers responsive to the stimuli used in drug release and the most commonly used polymers were introduced in each case. Temperature-sensitive polymers undergo phase separation as the temperature changes and the drug can be loaded or released from/into the polymer bulk. LCST and UCST polymeric systems enter the two-phase and single-phase regions with increasing temperature,

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respectively. pH-sensitive polymers were investigated and it was found that polymers capable of ionization or phase separation in environments with different pHs could be used in this method intelligently. Polyacids swell at pH greater than the pKa of the polymer while polybases swell at pH less than the pKa of the polymer. URDDS are one of the novel smart carriers with enhanced permeation of drug through the implosion of these microbubbles. These drug carriers are subjected to specific energy waves when they reach the desired location, and the drug exits their structure. Biologically responsive polymers such as antigenresponsive polymers, glucose-sensitive polymers, and protein-responsive polymers are sensitive to internal biological stimuli which are inherently present in the natural system. Electrically responsive polymers have similar behavior to pH-sensitive polymers due to ionization in the medium. There are various other novel methods such as shear stress-responsive polymers, redox-responsive polymers, SMPs, and MIPs, each with its own advantages and disadvantages for use in drug release systems. Any of the methods introduced in this chapter, or a combination of them, can help the expression of drug release systems for high-tech applications.

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CHAPTER 5 Stimuli-Responsive Polymers as Smart Drug Delivery

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[42] A. Abbasi, G.M.M. Sadeghi, I. Ghasemi, M. Shahrousvand, Shape memory performance of green in situ polymerized nanocomposites based on polyurethane/graphene nanoplatelets: synthesis, properties, and cell behavior, Polym. Compos. 39 (11) (2018) 4020–4033. [43] S. Erkec¸oglu, A.D. Sezer, S. Bucak, Smart delivery systems with shape memory and self-folding polymers, in: Smart Drug Delivery System, InTech, Rijeka, Croatia, 2016. [44] M.S. Vosoughi, M. Fallah-Darrehchi, P. Zahedi, Expanded graphene oxide-supported molecularly imprinted polymer nanoparticles based on polystyrene for dibenzothiophene removal, J. Sulfur Chem. 40 (5) (2019) 539–553. [45] O.I. Parisi, M. Ruffo, F. Puoci, Molecularly imprinted polymers for selective recognition in regenerative medicine, in: Nanostructured Biomaterials for Regenerative Medicine, Elsevier, 2020, , pp. 141–163. [46] L. Tan, J. Liu, W. Zhou, J. Wei, Z. Peng, A novel thermal and pH responsive drug delivery system based on ZnO@ PNIPAM hybrid nanoparticles, Mater. Sci. Eng. C 45 (2014) 524–529. [47] S. Zhao, R. Tian, B. Shao, Y. Feng, S. Yuan, L. Dong, L. Zhang, Z. Wang, H. You, One-pot synthesis of Ln3+-doped porous BiF 3@ PAA nanospheres for temperature sensing and pH-responsive drug delivery guided by CT imaging, Nanoscale 12 (2) (2020) 695–702.

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CHAPTER 6

Efficacy of Polymer-Based Wound Dressings in Chronic Wounds BLESSING A. ADERIBIGBE

Department of Chemistry, University of Fort Hare, Eastern Cape, South Africa

1. INTRODUCTION Wound care is a global health burden. Chronic wounds are common in the elderly population [1–3]. In the United States, it is estimated that 3% of the population with chronic wounds are older than 65 years and most of them suffer from open wounds. Over 2% of the total population of the United States has also been estimated to be affected by chronic wounds. In 2016, a report from Wales estimated a 6% cases of chronic wounds with a 5.5% cost to the National Health Service. Globally the United States and Europe are the biggest wounddressing markets. The annual cost for wound care in 2014 was reported to be an average of $2.8 billion [3]. It has been estimated that this amount will increase to $3.5 billion in 2021 [3]. The market research report in 2018 indicates that the global wound-closure products market will increase by more than $15 billion in 2022 (Internet source: Advanced Wound Care Market Outlook and Forecasts). It is estimated that by 2024, the wound care market targeting surgical wounds and chronic wounds such as ulcers will exceed $22 billion due to factors such as the advancement of technology [3] (Internet source: Advanced Wound Care Market Outlook and Forecasts). Factors that contributes to a wound becoming chronic are infections, poor circulation, age, alcohol, poor nutrition, obesity, diseases such as diabetes, cancer, HIV/AIDS, and some medications (antiinflammatory drugs, chemotherapy) [4]. The use of wound dressings that is not suitable for the wound type also contributes to a wound becoming chronic [5]. An ideal wound dressings should have features such as good capability to remove excess wound exudates, biodegradable, permeable to gaseous exchange, easy to use, affordable, protect the wound from infections, provide mechanical protection, affordable, biocompatible, reduce surface necrosis of the wound, does not cause trauma to the wound on removal, keeps the wound environment moist and dry and does not cause allergic

reactions [6]. Wounds can be classified as acute and chronic. Acute wounds represent an injury to the skin that heals by the normal phases of wound repair whereas chronic wounds need longer time to heal [7]. The healing process of wound is complex and complicated. Different strategies are used for the management of wounds depending on the nature of wounds. Wound occurs on the skin as a result of damage which can be physical or thermal. The skin is referred to the largest organ in the human body with an area of 2 m2 [8, 9]. The skin is made up of three important layers known as the dermis, epidermis, and hypodermis. The skin is very important in protecting the internal organs from ultraviolet (UV) radiation and invasion of microorganisms. It is also useful in regulating body temperature [10] and for the detection process of the sensory and immune system of the body [11]. Factor to be considered when selecting a wound dressing for the treatment of a wound are the location, type, depth, amount of exudates, wound adhesion, and infections. Wound dressings are prepared from a combination of biopolymers, synthetic polymers, and antibiotics. Biopolymers used for the preparation of wound dressings are either animal or plant origin. The biopolymers obtained from animals exhibit high porosity, biocompatibility, and biodegradability with good water uptake. Biopolymers such as collagen exhibit good hemostatic property and are useful for controlling bleeding wounds. Furthermore, the hemostatic mechanism of collagen-based wound dressings showed that platelets and hemocytes can adhere to the surface of the wound dressings followed by proliferation [12]. However, animal-based biopolymers are limited by their poor mechanical properties and their capability to transmit diseases to humans. Plant-based biopolymers used in the development of wound dressings are nontoxic in nature, biocompatible with good antimicrobial activity [10]. The present of hydrophilic functionalities in the plant-based biopolymers enhance the water

Modeling and Control of Drug Delivery Systems. https://doi.org/10.1016/B978-0-12-821185-4.00018-X © 2021 Elsevier Inc. All rights reserved.

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uptake capacity of wound dressings prepared from plant-based wound dressings. Synthetic polymers have been used widely in the preparation of wound dressings due to their excellent mechanical properties and flexibility. However, their use is limited by their poor biocompatibility and biodegradability. To overcome the aforementioned limitations, they are combined with biopolymers resulting in wound dressings with good features [10]. The loading of antibiotics into wound dressings is a good design approach suitable for the management of chronic wounds [6]. Antibiotics are useful in wound healing. The use of suitable concentrations of antibiotics for the treatment of infections is useful. However, high amounts of antibiotics can result in systematic toxicity [6]. To overcome the systematic toxicity of antibiotics in wound treatment, antibiotics are embedded in wound dressing which exhibit sustained and controlled drug release mechanisms. Some polymer-based wound dressings loaded with antibiotics for the treatment of microbial-infected wounds are hydrogels, foams, beads, dermal patches, films, nanoparticles, hydrocolloids, nanofibers, membranes, and others [7]. This chapter highlights the recent development of polymer-based wound dressings incorporated with antibiotics for the treatment of chronic wounds.

2. TYPES OF WOUNDS AND HEALING PHASE Wounds are classified as chronic and acute wounds. Examples of chronic wounds are pressure ulcers, diabetic ulcers, and vascular ulcers. The common features of chronic wounds are infections, excessive inflammation, and poor capability of the dermal cells to respond to repairs [13]. Chronic wounds are characterized by features such as a high level of proteases resulting in the destruction of extracellular matrix (ECM) thereby preventing the wounds from moving into a proliferative phase and persistent infections [13, 14]. The altered activities of proteases which are common in chronic wounds hinder the wound-healing process. A high level of proteases has been reported in wound exudates of chronic wounds such as pressure

FIG. 1 Wound-healing phases.

ulcers. The high level of proteases results in the degradation of fibronectin, an important protein useful in the remodeling of ECM. The high level of proteases is also responsible for the degradation of selected growth factors [15]. The high levels of cytokines in chronic wounds impede the wound-healing process and causes changes in normal skin fibroblasts [15]. Suppression of the activation of macrophages which is useful for the release of cytokines and growth factors to induce cells such as keratinocytes, endothelial, and fibroblasts is reduced in chronic wounds resulting in a weak inflammatory response [16]. Chronic wounds are characterized by delayed re-epithelization resulting from the failure of migration of keratinocytes and the altered composition of ECM such as fibronectin. The fibroblasts are unable to respond to growth factors and reduce the synthesized amount of collagen [14, 17]. The overexpression of peroxynitrite, a free radical formed from the combination of nitric oxide with hydroxyl-free radicals also delays the healing of chronic wounds [18]. It affects inflammation, vasculature, and collagen deposition [18, 19]. In chronic wounds, slough and necrotic tissues accumulate in the wounds resulting from the altered cellular environment, poor blood supply, debris from dying cells, and others [20]. The slough is composed of pus, dead cells, fibrin, leucocytes, microorganisms, and protein materials [20, 21]. The accumulation of slough and necrotic tissues in chronic wounds enhances the colonization of bacteria which impedes the wound healing by extending the inflammatory response, hindering wound contraction and re-epithelization [20–22]. Acute wound healing involves an organized mechanism which is divided into four phases namely: hemostasis, inflammation, proliferation, and remodeling phase (Fig. 1) [23]. The hemostasis and inflammation phase occurs simultaneously. Hemostasis phase occurs immediately after injury whereby the blood clotting takes place to prevent excessive bleeding. The platelets are activated at the hemostasis phase by type 1 collagen resulting in the release of growth factors and glycoproteins which causes the platelet to aggregate. During the platelet aggregation, clotting factors are released with the deposition of a fibrin clot at the wound site.

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Efficacy of Polymer-Based Wound Dressings in Chronic Wounds

The fibrin clot acts as a wound matrix in which the platelet produces growth factors that are important for wound healing [24]. In the inflammation phase, an influx of leucocytes at the wound bed is significant. Cells such as the monocytes, macrophages, neutrophils are useful in the inflammation phase. The wound is cleansed of infections and debris. Proinflammatory cytokines and growth factors activate and recruit fibroblast and epithelial cells. Neutrophils express reactive oxygen species (ROS), proinflammatory cytokines and proteases at the surface of the wound. Neutrophils recruit cells that clean-up the wound surface [24, 25]. In the proliferation phase, the inflammatory and migratory cells induce cellular proliferation resulting in the formation of granulation tissue and epithelization. A good blood supply suitable for the delivery of nutrient, gas, and metabolite exchange is important at this phase resulting in the proliferation of dermal and epidermal cells within the wound bed [23,26]. Angiogenesis also occurs at this phase resulting in the formation of new blood vessels. The wound is totally covered with epithelium with the formation of granulating tissues [25]. In the final wound-healing phase, there is a restoration of the integrity of the epidermal. The proliferation of the fibroblasts within the wound with the production of ECM-forming granulation tissue and new blood vessels are visible in this phase. The contraction of the wound resulting from fibroblast motility with a re-organization of the matrix takes place. The formation of scar tissue is also visible at this phase characterized by a tensile strength which is comparable with the unwounded skin. An increased number of apoptotic cells in the final phase of wound healing results from TGF-A (a tumor necrosis factor) and FGF-2 (a stimulator of cell proliferation). The formation of scars resulting from wound healing such as hypertrophic scar and keloid formation has been reported to be caused by the inability of myofibroblasts to undergo timely apoptosis [24, 25].

FIG. 2 Classification of wound dressings.

81

3. WOUND DRESSINGS Wound dressings are classified as traditional, bioactive, biologics, and interactive wound dressings (Fig. 2) [25] (Internet source: Wound Dressing). The traditional wound dressings are designed primarily to cover the wound surface such as gauze and gauze cotton composites. However, their use is limited by their poor permeation of vapor and ability to cause bleeding, leakage of exudates that promotes bacterial infections and damage to the newly formed epithelium [25,27] (Internet source: Wound Dressing). Interactive wound dressings such as foams, films, spray, gels, and composites are prepared from biopolymers and synthetic polymers. They exhibit distinct features such as good moist environment for improved wound healing, promote debridement, and enhance re-epithelialization [28]. Biologics-based wound dressings are classified as allografts, stem cell therapies, skin equivalents, tissue derivatives, and xenografts [25]. They are used for the treatment of chronic wounds. However, their use is limited by the risk of transmission of diseases and infections, they are expensive, the body may reject them due to immune reactions [25,29]. Biologic woundhealing therapies also involve the use of bioactive agents which can be obtained from plants. These bioactive agents exhibit biological properties such as antiinflammatory, antioxidant, and antimicrobial activities [30]. Monoterpene is a family of naturally occurring terpene-based compounds common in essential oils. Their level of toxicity is low making them useful in wound dressing. Biologic dressings induce autolytic debridement and granular wound bed [30]. Skin substitutes are used to replace the function of the skin and they are often used in chronic wounds. They accelerate the rate of healing with reduced side effects. They are classified as either acellular or cellular. Acellular skin substitutes are composed of a scaffold of fibronectin, hyaluronic acid, or collagen. The cellular skin substitute is composed of cells such as keratinocytes and fibroblasts [30].

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3.1. Hydrogels

Bioactive wound dressings are incorporated with antimicrobials and growth factors for accelerated wound dressings. They are used for the treatment of chronic wounds such as pressure ulcers and diabetic foot ulcers. They can overcome microbial infections and promote healing. Examples of bioactive dressings are hydrocolloids, collagens, alginates, hydrofibers, and chitosan. [25,29]. Examples of antibiotics incorporated into wound dressings are shown in Fig. 3A and B. Polymer-based wound dressings have been designed and incorporated with antimicrobials agents (Fig. 3C).

NH2

H2N

Hydrogels are cross-linked polymeric networks prepared by the reaction of monomers (Fig. 4). They exhibit a good capability to swell and sustain a significant amount of water within their network but they do not dissolve in water. They are characterized by good flexibility which is similar to natural tissue. Their ability to absorb water is attributed to the presence of hydrophilic functional groups in the polymeric backbone and their inability to dissolve in water is due to the cross-linking between network chains [31]. Factors that affect the

OH

H N

S

O HO O HO OH O

HN

O Amoxicillin

NH2

Cl

OH

O

NH2

R

N

O

HO

O

N

Benzalkonium chloride

Carvacrol

NH2 Gentamicin

O

OH O

O

F

OH

N H

O O

N Tetracycline

Ciprofloxacin

HO

OH

HO

N

N

NH2

N

NH2 O

H3CO

O

OH

HN

HO

OH O OH

NH2

H2N

OH

NH2 O

O

O O

HO P O OH

HO

HN

O X HO

Phosphomycin

Y OH

Neomycin

O HO

O OH OH

O O

(A)

Usnic acid

FIG. 3 (A) Structure of some antibiotics loaded to polymer-based wound dressings.

(Continued)

CHAPTER 6

Efficacy of Polymer-Based Wound Dressings in Chronic Wounds O

O

OH OH NH2

O N

S H2N

N NH

OH

Cl O

O

N

NH O

O

N O

S

Nitrofurazone

Cl O

Chloramphenicol

N

O

HN

N O

N O

O

O

Ceftazidime

H N

H N

N H

NH

NH

Cl

NH

H N

N H

Cl

NH

HO

OH O

OH O

N H

O

O

N OH

O O

Chlorhexidine

O OH

H2N Erythromycin

N H2N

HO HO

OH

HO

OH O

N

O

NH

O OHC

O H2N

NH2

OH

Streptomycin

HO OH

OH

O HN

H N

O N H

O Cl HO

O

NH2 O H N N H

O O

H N

OH

N

O

O NH

N

O

O

O

O

NH

OH

S N

O

O O

H N

OH

O

Cl

O

HO

Piperacillin OH

O

H2N OH O Vancomycin

O S

N N N

N O O

(B)

OH

Tazobactum

FIG. 3, CONT’D

(B) Some structures of some antibiotics loaded to polymer-based wound dressings. (Continued)

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(C) FIG. 3, CONT’D

(C) Polymer-based wound dressings incorporated with antibiotics.

FIG. 4 Hydrogel loaded with drugs.

properties of hydrogels are the types of polymers used and the degree of cross-linking. They can be prepared with tailored features such as biological response to stimuli, biodegradability, and mechanical strength [31]. The distinct properties of hydrogels are their good porosity which is useful for sustained drug delivery [32]. Researchers have reported hydrogels incorporated with different types of antibiotics for wound dressings resulting in accelerated healing (Table 1). Singh et al. designed hydrogel from the combination of carbopol and gum acacia. It was loaded with gentamicin, an antibiotic drug for enhanced wound-healing

potential. In vivo studies on Swiss albino mice of strain Balb C revealed faster wound healing characterized by a significant development of fibroblasts and blood capillaries. The hydrogels also exhibited nonthrombogenic, good antioxidant, and mucoadhesive activity. The hydrogels permeability to oxygen and moisture and impermeability to micro-organisms was significant. Their high absorption capability of (8.772
0.184 g/g) in simulated wound fluid further revealed their potential as wound dressings for acute wounds. The release profile of the loaded drug from the hydrogels was by using Fickian diffusion [33]. In another report by Singh

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Efficacy of Polymer-Based Wound Dressings in Chronic Wounds

85

TABLE 1

Hydrogels Loaded With Antibiotics for Wound Dressing. Polymers

Drugs loaded

Carbopol and acacia gum

Therapeutic outcomes

References

Gentamicin

In vivo studies revealed accelerated wound healing characterized by a significant development of fibroblasts and blood capillaries. The hydrogels were nonthrombogenic, good antioxidant with mucoadhesive activity and were permeable to oxygen and moisture but impermeable to microorganisms

[33]

Poly(2-hydroxyethyl methacrylate), carbopol, and acacia gum

Moxifloxacin

The hydrogel uptake of wound fluid was high with a low hemolytic capability of 0.95% suggesting that the high hydrophilic nature of the hydrogel network decreased its interaction with the red blood cell thereby reducing the degree of disruption of the red blood cell. In vivo studies revealed the formation of fibroblasts and blood capillaries with the absence of inflammation

[34]

Hyaluronic acid, adipic acid dihydrazide

Vancomycin

The hydrogel displayed excellent biocompatibility with high inhibition capability against methicillin-resistant Staphylococcus aureus

[35]

Chitosan and PEG

Ciprofloxacin

The hydrogels exhibited sustained release of ciprofloxacin for 24 h. The hydrogel antibacterial activity on Escherichia coli was more than 80% and sustained for 12 h

[36]

Poly(vinyl alcohol) and poly(acrylamide)

Amoxicillin

The release of the loaded antibiotics from the hydrogel was influenced by the pH and temperature of the release media. The drug-loaded hydrogel antibacterial activity was significant on the gram-positive bacteria

[37]

Poly(N-vinyl-2-pyrrolidone) and poly(ethylene glycol)

Neomycin

The drug release profile was sustained for 48 h. The drugloaded hydrogel displayed antibacterial effects on S. aureus. The drug release from the hydrogel diffused in the agar followed by binding to the DNA structure of the bacteria resulting in cell lysis

[38]

Gellan gum

Vancomycin

The in vitro growth inhibition effect on S. aureus and methicillin-resistant S. aureus in the presence of these hydrogels confirmed that vancomycin antibacterial activity was preserved. The hydrogels were nontoxic

[39]

Poly-1-vinyl-2-pyrollidone and carbopol

Moxifloxacin

The drug release profile of the loaded drug was slow and sustained which was influenced by the high surface area and porous structure of the hydrogels

[40]

Poly(ethylene glycol)

Ciprofloxacin

Irradiation of the hydrogel with a UV light of 365 nm triggered a drug release. The antimicrobial activity of the drug-loaded hydrogel was significant on S. aureus

[41]

Poly(vinyl alcohol) and poly(acrylamide)

Gentamicin

The hydrogels were permeable to oxygen and water vapor but inhibited bacterial growth

[42]

Pullulan

Gentamycin sulfate

The water uptake of the hydrogels was high (4000%) with a high hemostatic capability. The hydrogel ability to provide a moist environment over wound bed was significant. The hydrogels were biocompatible, cytotoxic and significantly inhibited bacterial proliferation. The drug release was fast in the first 2 h followed by a gradual drug release from the hydrogel over a period of 40 h

[43]

Continued

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Modeling and Control of Drug Delivery Systems

TABLE 1

Hydrogels Loaded With Antibiotics for Wound Dressing—cont’d Polymers

Drugs loaded

Keratin

Therapeutic outcomes

References

Ciprofloxacin

In vivo studies revealed a reduced wound contraction and re-epithelialization with the presence of collagen-rich granulation tissue and myofibroblasts

[44]

Carboxymethylcellulose

Tetracycline

The nanocomposite hydrogels swelling capacity was significant with good antibacterial activity on S. aureus

[45]

Poly(vinyl alcohol), sodium alginate and poly(vinyl pyrrolidone)

Neomycin

The wound curing effect of the hydrogel was attributed to its capability to provide a moist environment with good swelling capacity which is useful for the migration of cell growth factors

[46]

and Dhiman, 2-hydroxyethyl methacrylate-based hydrogel loaded with moxifloxacin for wound healing was prepared. The hydrogel uptake of wound fluid was 7.22 g/g suggesting its capability to prevent the leakage of excess exudates when used on high exuding wounds. It exhibited a hemolytic capability of 0.95% which was low suggesting that the high hydrophilic nature of the hydrogel network decreased its interaction with the red blood cell thereby reducing the degree of disruption of the red blood cell. However, the hydrogel’s thrombogenic nature was high and is attributed to the carboxylic groups on the surface of the hydrogel network which induce factor XII thereby promoting blood clot formation. The water vapor transmission rate was 348.57 g/m2/day and the release of the loaded drug was through a non-Fickian diffusion mechanism. In vivo studies further revealed the formation of fibroblasts and blood capillaries with the absence of inflammation. The hydrogel also exhibited free radical scavenging ability and the surface roughness influenced the mucoadhesive nature of wound dressings. The hydrogel was reported to be a potential wound dressing that can promote skin regeneration [34]. Liao et al. prepared hyaluronic acid-adipic acid dihydrazide hydrogel loaded with vancomycin. The hydrogel displayed excellent biocompatibility with high inhibition of methicillin-resistant Staphylococcus aureus [35]. Sharma et al. developed hydrogels by cross-linking chitosan with bifunctional PEG glyoxylic aldehyde followed by loading of ciprofloxacin. The hydrogels exhibited sustained release of ciprofloxacin for 24 h. The hydrogel antibacterial activity against Escherichia coli was more than 80% and sustained for 12 h [36]. Bajpai et al. loaded amoxicillin into poly(vinyl alcohol)-g-poly(acrylamide)-based hydrogels prepared by

graft copolymerization method. The amount of amoxicillin released increased with increasing content of poly(vinyl alcohol) in the hydrogel network but decreased with increase in the content of polyacrylamide in the hydrogel network. The release of amoxicillin from the hydrogel was also influenced by the pH and temperature of the release media. The drug-loaded grafted hydrogel also displayed antibacterial activity against gram-positive bacteria. The polar nature of polyacrylamide resulted in the amoxicillin molecules being held firmly by the network chains and the reduced amount of the drug release [37]. Zafalon et al. prepared hydrogels from poly(N-vinyl-2-pyrrolidone) and poly(ethylene glycol). The hydrogels were loaded with neomycin followed by gamma irradiation to induce cross-linking and sterilization. The drug release profile was sustained for 48 h. The drug-loaded hydrogel displayed antibacterial effects on S. aureus in vitro [38]. The antibacterial results indicate that the release of neomycin from the hydrogel diffused in the agar followed by binding to the DNA structure of the bacteria resulting in cell lysis. The drug release of the antibiotic was rapid after application and reached equilibrium over a period of 8 h. The water uptake into the hydrogel was rapid with a maximum swelling of 1100% obtained in 24 h. The aforementioned findings indicate that the drugloaded hydrogel is suitable for the treatment of high exuding and infected wounds [38]. Shukla and Shukla reported hydrogels prepared from 1% wt./vol. gellan and 1 mM CaCl2. Vancomcyin, was loaded in the hydrogels directly and indirectly by carbon nanoparticles. The release of vancomycin from the hydrogel with carbon nanoparticles was for 9 days and for 6 days in hydrogels without the nanoparticles. The drug release from the hydrogels was influenced by

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Efficacy of Polymer-Based Wound Dressings in Chronic Wounds

an intermolecular interaction between the drug and the polymer network. The hydrogel in vitro growth inhibition effect on S. aureus and methicillin-resistant S. aureus in the presence of these hydrogels confirmed that vancomycin antibacterial activity was preserved. In vitro viability of fibroblasts and mesenchymal stem cells, common wound-healing cells for 1 and 9 days revealed the viability of the cell indicating the nontoxic nature of the hydrogels [39]. Singh and Dhiman prepared hydrogels with wound fluid absorption capacity of 11.37 g/g. The hydrogel was prepared using poly-1-vinyl-2-pyrollidone and carbopol. It was loaded with moxifloxacin. The drug release profile of drug-loaded hydrogel was slow and sustained which was influenced by the high surface area and porous structure of the hydrogel. The hydrogel exhibited a high absorption of the simulated wound fluid and the hydrogel was nonhemolytic, nonthrombogenic, with good oxygen and water vapor permeability. The swelling capability of the hydrogel in simulated wound fluid was 1137.35% when compared with the swelling in pH 2.2 buffer which was 287.49% and 1100.04% in phosphate buffered solution. The presence of ciprofloxacin in the hydrogel also contributed to the pH sensitivity of the hydrogels [40]. Shi et al. prepared crosslinked hydrogel which was incorporated with ciprofloxacin by a linker cleavable by light. Irradiation of the hydrogel with a UV light of 365 nm triggered a drug release. The antimicrobial activity of the drug-loaded hydrogel was significant on S. aureus [41]. Singh et al. prepared sterculia crosslinked poly(vinyl alcohol) and poly(vinyl alcohol-acrylamide) hydrogel wound dressings loaded with gentamicin. The hydrogels absorbed 4.80 and 6.32 g/g simulated wound fluid and the release of antibiotic drugs was by non-Fickian and Case II diffusion mechanisms, respectively. The hydrogels were permeable to oxygen and water vapor but inhibited bacterial growth. The release of antibiotic from the hydrogel was slow [42]. Li et al. prepared pullulan-based hydrogels by chemical cross-linking followed by the loading of gentamycin sulfate. The tensile strength of the hydrogels was in the range of 0.663–1.097 MPa which was dependent on the degree of cross-linking. The water uptake of the hydrogels was high (4000%) with a high hemostatic capability. The water vapor transmission rate was in the range of 2213–3498 g/m2/day. The water retention ability of the hydrogel was in the range of 34.74%–45.81% over a period of 6 days suggesting that the hydrogels can provide a moist environment over the wound surface thereby preventing the dehydration of the wound surface and scab formation. The hydrogels were biocompatible, cytotoxic, and significantly

87

inhibited bacterial proliferation. The drug release was fast in the first 2 h followed by a gradual drug release from the hydrogel over a period of 40 h [43]. Roy et al. loaded ciprofloxacin, an antibiotic into keratin-based hydrogels. In vivo studies were performed on 10-mm full-thickness wounds inoculated with 106 colony-forming units of Pseudomonas aeruginosa. The hydrogels reduced the amount of the bacteria in the wound bed by over 99.9% when compared with the untreated wounds at days 3, 7, and 11. The hydrogels also promoted a reduced wound contraction and re-epithelialization at day 7. At day 11, the wounds treated with the hydrogels displayed myofibroblasts and collagen-rich granulation tissue. Between day 7–11, an increase in macrophages in the wound bed was visible. Loading ciprofloxacin into the hydrogels prevented wound infection and did not interfere with the healing process. The rate of drug release from the hydrogels was in the range of 50%–63% in 5 days and it was not dependent on the initial concentration of ciprofloxacin [44]. Namazi et al. reported nanocomposite hydrogel prepared from mesoporous silica MCM-41 as a nanodrug carrier into carboxymethylcellulose hydrogel. The hydrogel was loaded with tetracycline and methylene blue. The nanocomposite hydrogel-swelling capacity was significant with antibacterial activity on S. aureus [45]. Choi et al. prepared hydrogels loaded with neomycin. They were prepared from varied amount of poly(vinyl alcohol), sodium alginate, and poly(vinyl pyrrolidone) by freeze thawing method. The hydrogel composed of poly(vinyl alcohol), sodium alginate, poly(vinyl pyrrolidone), and neomycin in the ratio 10/0.8/0.8/1 released 85% of the loaded drug within 2 h. The wound curing effect of the hydrogel was attributed to its capability to provide a moist environment with a good swelling capacity which is useful for the migration of cell growth factors. However, the loaded drug did not have any significant influence on the wound curing ability of the hydrogels [47].

3.2. Hydrocolloids Hydrocolloid dressings are composed of materials such as pectin, carboxylmethylcellulose, gelatin, and cellulose incorporated in an adhesive film or foam (Fig. 5). They are characterized by an interaction between the materials and the wound exudates resulting in their swelling. However, they are not permeable to gas, vapor, water, and bacteria [47–49] . Their transparent appearance makes it possible to easily monitor the wound healing and wound exudate. They leave residue in the wound bed and are not suitable for high exuding wounds, bleeding wounds, and infected wounds [47].

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Modeling and Control of Drug Delivery Systems

Enhanced angiogenesis and the formation of granulation tissue

Hydrocolloid loaded with antimicrobial agents FIG. 5 Illustration of hydrocolloid loaded with antimicrobial agents.

TABLE 2

Hydrocolloid and Foams Loaded With Antibiotics. Wound dressings

Polymers

Drug

Therapeutic outcomes

References

Hydrocolloid

Alginate

Benzalkonium chloride

The hydrocolloid displayed excellent antimicrobial activity on selected strains of bacteria used in the study such as Staphylococcus aureus, Escherichia coli, and Pseudomonas aeruginosa. A significant epithelialization in the wound periphery was significant in vivo. The upper dermal matrix revealed a wellstructured granulation tissue with the absence of degeneration

[50]

Hydrocolloid

Poly (isobutylene)

Centella asiatica

The wound healing using the hydrocolloid was 95% when compared with the control on day 29. The wound dressings accelerated epithelialization. Accelerated healing was significant when compared with the control

[51]

Hydrocolloid

Silk fibroin

Neoderm

In vivo studies revealed the appearance of healthy tissue at 14 days without signs of edema at 21 days. A significant increase in the density of collagen fibers suggesting excellent structural integrity of the tissue during wound healing. A regenerated a dermis layer was visible at day 7 indicating accelerated wound healing

[52]

Kendall Foam

Polyurethane

Poly (hexamethylene biguanide)

It is suitable for the management of overgranulation at gastrostomy sites. It prevents infection

[53–57]

The advantages of hydrocolloids are enhanced angiogenesis and the formation of granulation tissue. However, their debriding capability can result in an increase in the size of the wound, they also cause skin maceration around the wound bed and produces unpleasant odor which can be confused for infection [48] (May et al., 2012). There are some research reports on hydrocolloids designed for wound dressings (Table 2). Thu et al. developed a bilayer hydrocolloid film prepared from alginate and loaded with ibuprofen in the upper layer [58]. The dressing bilayer displayed unique

mechanical and rheological properties. The bilayers also exhibited slower hydration rate and low drug flux in vitro when compared with the single layer. The bilayers displayed a significant healing rate in vivo, characterized by the formation of epidermis with accelerated granulation tissue formation [58]. Jin et al. prepared hydrocolloids loaded with benzalkonium chloride by a hot melting method [51]. The wound dressing gave exhibited good swelling capability, skin adhesion, mechanical strength, and flexibility when compared with the commercial wound dressing. It also displayed excellent antimicrobial activity against S. aureus, E. coli,

CHAPTER 6

Efficacy of Polymer-Based Wound Dressings in Chronic Wounds

and P. aeruginosa. The wound dressing in vivo induced a coating with significant epithelialization in the wound periphery. The upper dermal matrix displayed a wellstructured granulation tissue with the absence of degeneration in the dermal matrix. The hydrocolloids restored the normal skin tissues. Hiranpattanakul et al. prepared chitin/chitosan hydrocolloid for wound dressing. The hydrocolloids water adsorption, excellent antibacterial activity, and biocompatibility make them potential wound dressings. Increasing the chitosan microbeads: chitin weight ratio reduced the water absorption. They displayed 100% antibacterial activity after 30 min exposure to E. coli and S. aureus. They were biocompatible with high water absorption and good degradability [59]. Jin et al. developed sodium alginate-based Centella asiatica-loaded hydrocolloid wound dressing [50]. The use of alginate to prepare the hydrocolloids enhanced their capability to retain fluid and also provide a moist environment for the wound. The hot melt pressuresensitive adhesives and elastomers used were poly(isobutylene) and SIS which induced the chemical inertness of the dressings. Petroleum resin hydrocarbon improved the adhesive nature of the dressing together with the strength and flexibility of the dressings. Rat induced with diabetes and wounds were used for the in vivo studies. The excision wound models were evaluated for 1 month and the wound healing using the hydrocolloid was 95% when compared with the control on day 29. The wound dressings accelerated epithelialization. In infected wound models inoculated with S. aureus inflammation occurred. A severe form of inflammation and hemorrhage was significant in the control and a decreased wound size resulting from accelerated epithelialization was visible in the wounds in which the hydrocolloid dressing was used. In abrasion wound model, CA-loaded hydrocolloid also revealed accelerated healing when compared with the control. The moist environment around a wound makes the wound soft and induces the migration of keratinocytes and fibroblasts, cell growth factors, and cytokines thereby accelerating epithelialization and wound healing [50,51]. The reoccurrence of keloids can be prevented by combining hydrocolloids with magnets. The magnet compresses the wound thereby increasing discharge while the hydrocolloid dressings provide a moist environment for the wound. In patients with earlobe keloid who had surgical excision followed by hydrocolloid dressing combined with magnet resulted in good postoperative outcomes. No significant acute wound effects with a good long-term recurrence rate when compared with patients treated with conventional dressing

89

methods [60]. Wound dressing composed of a layer of hydrocolloid and a layer of activated carbon was developed. The activated carbon exhibited bacteriostatic or antimicrobial activity by inhibiting the growth of bacteria and also absorbed wound fluids and unpleasant smell. The hydrocolloid layer induced wound healing by providing a moist warm environment suitable for enzymatic processes and efficient wound healing. The layer also induced debridement of dead tissue and exhibited good hemostatic activity. The combination of an activated carbon layer with a hydrocolloid layer resulted in a synergistic effect and accelerated wound healing [52]. Lee et al. developed hydrocolloid dressings loaded with silk fibroin nanoparticles for the treatment of burn wounds [61]. In vivo studies revealed the appearance of healthy tissue at 14 days without signs of edema at 21 days. A significant increase in the density of collagen fibers suggested excellent structural integrity of the tissue during wound healing. A regenerated dermis layer was visible at day 7 indicating accelerated wound healing when compared with gauze and Neoderm. The water uptake of the prepared hydrocolloid was similar to Neoderm. Increase in the content of the nanoparticles in dressing retained the shape of the dressing after the addition of water and the tensile strength also increased. The hydrocolloids were biocompatible.

3.3. Foams Foam-based wound dressings are porous and are able to absorb fluids [62]. Their thickness varies and they can be either adhesive or nonadhesive. The contact layer of the foam on the wound area facilitates the uptake of the exudates into the foam [63]. Its adhesive nature to the surrounding skin helps in keeping the dressing in place thereby preventing excess exudates around the wound which can promote bacteria invasion (Fig. 6) [64]. However, its adhesive nature may cause skin irritation in patients with sensitive or fragile skin. The most commonly used foam is polyurethane and silicone foam is used less frequently. Silicone foams are used as a primary absorbent in wound dressings [62,63]. Foams prepared with film-backing act as a resistant barrier to water and microbial invasion. The permeability of the film-backing varies and influences the foams capacity of water evaporation and gas exchange [63]. Foam-based wound dressings offer several advantages such as good capability to maintain moisture at the wound surface; can be easily removed and protects the skin around the wound; it protects the wound against bacteria; it maintains a temperature suitable for accelerated wound healing; it provides mechanical

90

Modeling and Control of Drug Delivery Systems

FIG 6 Diagram of foam for wound dressing.

protection; they are nontoxic; cost effective with a long shelf life. They are useful for the management of chronic and acute wounds (Table 2) [62,64]. Pyun et al. developed silver-hydroxyapatite-loaded polyurethane foams. In vivo wound healing was evaluated in Sprague-Dawley rat model. Silverhydroxyapatite particles were uniformly dispersed in the foams. The in vitro release of silver from the foam was influenced by time and concentration. The foams displayed good antibacterial activity and were noncytotoxic on L-929 fibroblast cells in vitro. In vivo studies on the foam revealed a scar-free wound healing which was characterized by accelerated re-epithelialization and collagen deposition in the excision wound model [65]. Commercially available foams are Gentian violet and methylene blue (GV/MB) antibacterial dressings in poly(vinyl alcohol) (PVA) foam and polyurethane (PU) foam bonded with GV and MB with a thin film backing. The latter foam does not require hydration or a secondary dressing. They are affordable when compared with other wound dressings. They accelerate wound healing due to features such as easy to use, do not require frequent changes and promote autolytic debridement [66]. Kendall AMD antimicrobial foam dressings are incorporated with an antimicrobial agent, poly(hexamethylene biguanide) (PHMB). They are characterized by high absorption capacity and are designed for the management of acute or chronic wounds which produce moderate to high amount of exudates. They are made of polyurethane foam and provide moisture environment for the wound and inhibit the growth of bacteria. The presence of a polyurethane film prevents the leakage of exudate. The wound contact surface of the foam is nonadherent with an open-cell honeycomb structure that promotes high absorption of exudates into the core of the dressing. The inner core of the foam has a large honeycomb structure which also promotes the retention of exudates. Its good absorption capability prevents the maceration of the surrounding skin around the wound bed. Any bacteria in the

exudates are exposed to the antimicrobial action of the loaded PHMB. PHMB mode of action on bacteria is via binding to bacteria cell’s outer membrane and it disrupts the integrity of the cell membrane; inhibit bacteria cell metabolism and induce cell death [55]. Researchers reported the efficacy of Kendall™ AMD antimicrobial foam dressings in the management of chronic wounds. Sibbald et al. performed a clinical trial on patients with leg and foot ulcers over a period of 5 weeks. The PHMB foam dressing reduced wound superficial bacterial burden significantly at week 4 with (P ¼ .016) when compared with the foam alone. It also reduced pain significantly at week 2 with (P ¼ .0006) and at week 4 with (P ¼ .02). A significant wound reduction by a 35% median by week 4 was also reported when compared with 28% in the control group [54]. Warriner and Spruce performed a clinical trial to investigate the potential application of PHMB foam for the management of overgranulation at gastrostomy sites. PHMB-impregnated foam dressing at 2 weeks resolved the overgranulation tissue in one-third of the patients, at week 4 for three patients and at week 6 for another three patients. Although overgranulation is not life threatening, it is responsible for factors such as odor, bleeding, and exudates which can affect patient quality of life, psychologically [56]. Evans reported the use of Kendall AMD Antimicrobial Foam Dressing with PHMB for the prevention and management of infection of exit sites [57]. PHMB-treated Kendall AMD antimicrobial foam dressing exhibit antimicrobial activity against MRSA suggesting its potential to reduce bacterial contamination of the wound by wound dressings (Internet source: MRSA FD. Efficacy of Kendall; [53]).

3.4. Films Films have been designed as wound dressings (Table 3). Peles and Zilberman developed films from soy protein isolate followed by the loading of gentamicin for controlled release. The films exhibited high tensile strength with suitable Young’s modulus. Factors such as the type

CHAPTER 6

Efficacy of Polymer-Based Wound Dressings in Chronic Wounds

91

TABLE 3

Films Loaded With Antibiotics. Polymers

Antibiotics

Therapeutic outcomes

References

Soy protein isolates

Gentamicin

The drug release profile of was sustained for 4 weeks by diffusion

[67]

Natural rubber

Gentamicin

The films also exhibited good antimicrobial activity against Staphylococcus aureus and Pseudomonas aeruginosa with angiogenic activity

[68]

Poly(vinyl pyrrolidone)

Ciprofloxacin

The films were active against Escherichia coli and Bacillus subtilis with high absorption capability of wound exudates in vivo

[69]

Soy protein isolates

Ciprofloxacin

The antibacterial activity of the films against S. aureus and P. aeruginosa strains was significant

[70]

Soy protein isolates

Gentamicin

The films effectively inhibited S. aureus and S. albus infections over a period of 2 weeks and P. aeruginosa for 3 days

[71]

Chitosan

Ciprofloxacin

The films were capable of detaching from a wound site when swiped with water at a temperature lower than body temperature suggesting that it can relieve pain and injury

[72]

Chitosan

Erythromycin

The drug release was 40.3% and 72.5% for films with and without nanoparticles in 22 h, respectively. The films loaded with drug nanoparticles maintained a moist environment over the wound bed in heavy and moderate exuding wounds with a sustained drug release profile

[73]

The release of gentamicin from the cross-linked films was sustained and effective in inhibiting bacterial proliferation and protecting the wound from infections

[74]

Poly(vinyl alcohol) and chitosan Poly(ethylene oxide)

Streptomycin

The release of the loaded drugs over a period of 72 h was controlled. These drug loaded films exhibited high zones of inhibition against P. aeruginosa, S. aureus, and E. coli

[75]

Chitosan

Ciprofloxacin

The drug release was 15%–20% over a period of 48 h. They were biocompatible and nontoxic

[76]

Chitosan

Gentamicin sulfate

Good biocompatibility with sustained drug release and useful for the protection of the wound from infections

[77]

Chitosan

Silver sulfadiazine

The release profile of silver sulfadiazine was sustained for 4 days with over 82% of drug release. The drug-loaded films were effective against P. aeruginosa and S. aureus strains of bacteria. The films were nontoxic and promoted cell proliferation

[78]

Poly(lactic acid)

Gentamicin sulfate or metronidazole

The gentamicin-loaded films inhibited Proteus mirabilis, S. aureus, and P. aeruginosa growth while metronidazoleloaded film inhibited Bacteroides fragilis for 7 days

[79]

Chitosan and gelatin

Ciprofloxacin

The percentage wound contraction was enhanced for wounds treated with ciprofloxacin-loaded film

[80]

Sodium carboxymethylcellulose

Chlorhexidine

The films showed potent antimicrobial and antibiofilm activity. In vitro cytotoxicity evaluation on human keratinocytes and fibroblasts revealed low cytotoxic effect

[81]

Continued

92

Modeling and Control of Drug Delivery Systems

TABLE 3

Films Loaded With Antibiotics—cont’d Polymers

Antibiotics

Therapeutic outcomes

References

Cellulose

Amikacin and ceftriaxone

The films displayed good antibacterial activity against E. coli, P. aeruginosa, Klebsiella pneumoniae, and S. aureus

[82]

Pluronic F127, pectin and gelatin

Erythromycin

The film’s mechanical properties were superior with extended drug release. The films were nontoxic with good antibacterial activity

[83]

The porous face of the films is suitable for the preloading of bioactive agents and quick drug release with enhanced water vapor permeability

[61,84]

Polydimethylsiloxane

Collagen

Usnic acid

In vivo studies on second-degree burn wounds in 45 Wistar rats showed that the use of usnic acid improved the formation of collagen and rapid epithelization indicating that they are suitable for burn wound healing

[85]

Chitosan

Chlorhexidine

Prolonged drug release with good antimicrobial and antibiofilm activities with no cytotoxic effects

[86]

Poly(vinyl alcohol)

Tetracycline

In vitro release studies showed a slow drug release with good antibacterial activity against S. aureus, E. coli, and Enterococcus faecium

[87]

Chitosan

Norfloxacin

In vitro studies show that the present of norfloxacin in the film reduced the number of human dermal fibroblast cell after 72 h incubation. The antibacterial activities of the films was significant against S. aureus, Bacillus cereus, E. coli and K. pneumoniae strains of bacteria

[88]

Chitosan and poly(vinyl alcohol)

Nitrofurazone

The films exhibited high antibacterial activity against P. aeruginosa

[89]

of plasticizer used, the type of cross-linking agent used and the method of cross-linking influenced the tensile properties of the SPI films, significantly. The water vapor transmission rate of the films was 2300 g/m2/day and it was not influenced by the cross-linking method. The drug release profile of the film was sustained for 4 weeks. The drug release was via diffusion suggesting their potential to protect the wound against bacterial infection [67]. Phaechamud et al. developed porous natural rubber films loaded with gentamicin sulfate. They were prepared by blending glycerine, triethyl citrate, and xanthan gum by solvent casting method. The films exhibited good water sorption capacity with low adhesive property, sustained drug release for 7 days, good water vapor and oxygen permeability. They also displayed good antimicrobial activities against S. aureus and P. aeruginosa with angiogenic activity [68]. Contardi et al. developed poly(vinyl pyrrolidone)based films loaded with ciprofloxacin. The presence of residual PVP-bound acetic acid in the films contributed to the plasticity nature and the softness of the films. The

films were active against E. coli and Bacillus subtilis. In vivo studies on full-thickness excisional skin mice model showed the films capability to resorbed. The absorption of the wound exudates was dependent on the concentration of acetic acid used to prepare the films [69]. Rivadeneira et al. reported the influence of different ratio of soy protein isolate and agar in films blends (3:1, 1:1, and 1:3), respectively. The films were prepared by the gel casting method and loaded with ciprofloxacin hydrochloride. The antibacterial effect of the films against S. aureus and P. aeruginosa strains was significant and not influenced by the ratio. However, the water uptake of the films and the drug release profile was influenced by the ratio of the materials used to prepare the films. The amount of drug released decreased to 80% resulting from a decrease in the ratio of agar content in the films [70]. Peles et al. also prepared soy protein films loaded with gentamicin. The films effectively inhibited S. aureus and Staphylococcus albus infections over a period of 2 weeks and P. aeruginosa for 3 days. The films showed no cytotoxic effect in vitro. However,

CHAPTER 6

Efficacy of Polymer-Based Wound Dressings in Chronic Wounds

a mild inhibition in cell proliferation after 3 days was visible in films containing a high concentration of glycerol. However, the films were biocompatible [71]. Ngadaonye et al. developed thermoresponsive interpenetrating polymer networks by photo-polymerization and cross-linking of DEAAm with chitosan followed by loading of ciprofloxacin. The films were transparent, highly permeable to vapor and displayed temperaturesensitive swelling properties. The Young’s modulus of the films reduced significantly after the incorporation of the DEAAm and glycerol into the film network. The percentage strain for all the dry films increased indicating the good flexibility of the films. The films at 25 °C exhibited high swelling capability which eliminated their adhesive property. At 37 °C, the films were hydrophobic with low swelling capability. These findings indicate that the film dressing can detach itself from the wound site when swiped with water at a temperature lower than body temperature. The ability of the wound dressing to detach can also relieve pain and injury which is often caused by dressing change. The water vapor transmission rate of the films was in the range of 31.2–37.9 g/m2/h at 37 °C. The films containing 30% (w/w) PDEAAm displayed low water vapor permeability suggesting high hydrophobic nature of the film networks. The cumulative percentage drug release from the films was 98.5% in 7 h. The hydrophobic nature of the dressings with high content of DEAAm decreased the diffusion rate of the drugs. Gentamicin and ciprofloxacin-loaded film dressings inhibited the growth of S. aureus and P. aeruginosa [72]. Zhang et al. prepared konjac glucomannan/chitosan film loaded with erythromycin nanoparticles. The equilibrium water content of the film was 99% indicating its capability to prevent excess exudates on the wound bed. The water adsorption of the film was 2362.3% with water vapor transmission rate of 2335
36/g2/day and evaporative water loss of 10% at 1 h and 90% in 6 h. In 22 h, the drug release was 40.3% and 72.5% for films with and without nanoparticles, respectively. The films loaded with drug nanoparticles maintained a moist environment over the wound bed in moderate and heavy exuding wounds with a sustained release of the antibacterial agent on the wound surface [73]. Zhang et al. developed carboxyl-modified poly(vinyl alcohol)-cross-linked chitosan hydrogel films by grafting poly(vinyl alcohol) with succinate acid. The water vapor and oxygen permeability of the crosslinked hydrogel films indicated that they can maintain a moist environment over the wound bed. The films did not exhibit cytotoxic and hemolytic potential. The release of gentamicin from the cross-linked films was sustained

93

and effective in the inhibition of bacterial proliferation and the protection of the wound from infection [74]. Boateng et al. prepared carrageenan-based and polyethylene oxide-based films which were loaded with a combination of an antibiotic, streptomycin with diclofenac for improved healing of chronic wounds. The films were smooth with excellent transparency and homogeneous surface morphology. Their elasticity was high with good mechanical properties. The films capability to absorb simulated wound fluid was high in vitro. The aforementioned feature indicates that they can protect the wound from bacterial invasion resulting from excess exudates on the wound bed. The release of both drugs was controlled for 72 h. Furthermore the films high zones of inhibition on P. aeruginosa, E. coli, and S. aureus was significant when compared with the individual free drugs (streptomycin with diclofenac). The loading of streptomycin was useful for the treatment of chronic wound infections while diclofenac was used to relieve swelling and pain during the inflammatory phase of wound healing [75]. Guan et al. reported a facile and green process for the fabrication of film from quaternized hemicelluloses and chitosan using epichlorohydrin as the cross-linker. The films were loaded with ciprofloxacin. Its tensile strength was over 37 MPa with a roughness of 2–5 nm in the area of 400 nm. The highest drug loading concentration was 18%. The drug release was 15%–20% over a period of 48 h. In vitro cytotoxicity evaluation on 293T cell viability assay indicated good biocompatibility and nontoxicity [76]. Jiang et al. prepared carboxyl-modified hypromellose-crosslinked chitosan films loaded with gentamycin sulfate as a potential wound dressing. Hypromellose was grafted with succinic acid which was reinforced via amide bond formation using a coupling agent. The films displayed good biocompatibility with high swelling ratio, water vapor transmission rate, and oxygen permeability suitable for tissue regeneration. The drug release from drug-loaded films was sustained and useful for the protection of the wound from infection [77]. Fajardo et al. prepared chitosanchondroitin sulfate films loaded with silver sulfadiazine with varied ratio of chitosan and chondroitin sulfate. The swelling behavior of the films was pH-dependent. The amount of chitosan used to prepare the film and the loading of silver sulfadiazine affected the tension of rupture of the films. The Young modulus of the films was influenced by the amount of chitosan in the film network. A sustained release profile of silver sulfadiazine from the film was sustained for 4 days with over 82% of drug release. The drug-loaded films were effective against P. aeruginosa and S. aureus strains of bacteria.

94

Modeling and Control of Drug Delivery Systems

The films were nontoxic and promoted cell proliferation [78]. Chitrattha and Phaechamud prepared poly(lactic acid) porous film by solvent casting method. Gentamicin sulfate or metronidazole was incorporated into the films. The film exhibited a porosity of 55.31% porosity with 20-μm pore size which enhanced the oxygen transmission rate, water vapor transmission rate, degradation rate, and percentage of drug release from the films. The gentamicin-loaded films inhibited S. aureus, P. aeruginosa, and Proteus mirabilis growth while metronidazole-loaded film inhibited Bacteroides fragilis for 7 days [79]. Tvl et al. developed chitosan-gelatin films loaded with ciprofloxacin. The drug-loaded films showed good water absorption capacity, folding endurance and antibacterial activity. The percentage of wound contraction was enhanced for the wounds treated with ciprofloxacin-loaded film. The ciprofloxacin drugloaded films capacity for fast healing and rapid epithelialization of skin is attributed to its antibacterial action which prevents infections of the wound ( [80]). Donnadio et al. prepared sodium carboxymethylcellulose films loaded with chlorhexidine into the layered zirconium phosphate nanoparticles with potential antibiofilm activity. Chlorhexidine was intercalated between the layers of zirconium phosphate nanoparticles for prolonged release. The films showed potent antimicrobial and antibiofilm activity. In vitro cytotoxicity evaluation on human keratinocytes and fibroblasts revealed a low cytotoxic effect [81]. Volova et al. prepared cellulose composites films loaded with silver nanoparticles and antibiotics. The films were loaded with amikacin and ceftriaxone. The films displayed good antibacterial activity against E. coli, P. aeruginosa, Klebsiella pneumoniae, and S. aureus [82]. Alavi et al. developed film dressings made of pluronic F127 and pectin or pluronic F127 and gelatin using glycerol (2.5%) as a plasticizer. Erythromycin (0.1%) was loaded in the films. The optimized film (1:4; pluronic F127/Pectin and pluronic F127/gelatin) displayed a smooth, translucent, and good flexibility. The drugloaded film exhibited smoother surface morphology. The films mechanical properties were superior with extended drug release suggesting good patient compliance by reducing dressing changes. The films were nontoxic with good antibacterial activity [83]. Lee et al. developed a thin soft Janus polydimethylsiloxane (PDMS) film with good porosity, stretchability, and water-wettability [84]. The film can be stretched to 150% of its original length and the nonporous face of the film contributes to its waterproof barrier. The porous face of the films is suitable for the preloading of

bioactive agents and quick drug release with enhanced water vapor permeability. Blood absorbed by the film’s network is distributed thereby preventing the formation of a large clot and hardened scab that can cause a secondary injury to the wound bed upon removal [61,84]. Nunes et al. evaluated collagen-based films loaded with usnic acid as a wound dressing for burn wounds. In vivo studies on second-degree burn wounds in 45 Wistar rats showed that the use of usnic acid improved the formation of collagen with rapid epithelization indicating that they are suitable for burn wound healing. The presence of usnic acid in the liposomes prevented over deposition of collagen fibers [85]. Ambrogi et al. reported chitosan/montmorillonite composite films containing chlorhexidine with prolonged drug release. The drug was intercalated between the layers of montmorillonite and the films and they displayed good antimicrobial and antibiofilm activities with no cytotoxic effects. The results revealed their potential use as a wound dressing material for the prevention of microbial colonization in wounds [86]. Niamlang et al. prepared poly(vinyl alcohol) films loaded with tetracycline hydrochloride encapsulated in quaternized chitosan nanoparticles. The success of the encapsulation was confirmed by Fourier transform infrared (FT-IR) spectroscopy. The encapsulation efficiencies of the drug-loaded chitosan nanoparticles were in the range of 72%–95%. In vitro release studies showed a slow drug release with good antibacterial activity against E. coli, S. aureus, and Enterococcus faecium [87]. Sebri and Amin developed chitosan films loaded with norfloxacin to combat bacterial infection around the wound area. The loading of norfloxacin into the films enhanced the flexibility of the film. In vitro studies show that the present of norfloxacin in the film reduced the number of human dermal fibroblast cell after 72 h incubation. The antibacterial activity of the films was significant against E. coli, Bacillus cereus, S. aureus, and K. pneumoniae strains of bacteria [88]. Kouchak et al. prepared blend films using chitosan and poly(vinyl alcohol) loaded with nitrofurazone by casting evaporating technique [89]. Loading nitrofurazone into the films reduced the tensile strength, swelling ability, oxygen permeability of the films. However, the water vapor transmission rate was increased. The films exhibited high antibacterial activity against P. aeruginosa [89].

3.5. Dermal Patches A transdermal patch is composed of a multilayered structure with an impermeable film loaded with drugs and excipients which is suitable for good skin adhesion and a protective release liner. The liner must be removed

CHAPTER 6

Efficacy of Polymer-Based Wound Dressings in Chronic Wounds

before the application of the patch to the skin [90]. There are different designs of dermal patches depending on their applications. The reservoir/membrane-type patch is characterized by a constant drug release following a zero-order release mechanism. However, this design requires a large patch for increased drug delivery. During storage, highly soluble drugs can diffuse and saturate the membranes of the system resulting in high initial drug delivery. The aforementioned factor reveals the need for a flux moderation for patches loaded with highly soluble drugs [90]. The rupture of the backing membrane can occur accidentally. Matrix patches do not have a liquid reservoir and they are applied to the skin by glueing to the skin. The drug is loaded in the adhesive polymer and the patch exhibit adhesion function and also controls the drug delivery rate. The loading of a drug that is completely in solution results in the rate of the drug release being dependent on the drug concentration in the adhesive, first-order kinetics. The aforementioned factor is a limitation associated with matrix patches indicating that a decrease in the drug release rate with wear time [90,91]. Most drugs are not suitable for patch delivery. Drugs suitable for patches are those that can penetrate the skin. The maximum skin penetration flux for a drug is determined by factors such as solubility and diffusivity in the stratum corneum [92]. Their solubility is influenced by the melting point and the interaction of the drug with the stratum corneum. The diffusivity of the drug is influenced by the molecular weight [93]. Dermal patches have been designed for wound dressings (Table 4). Auda et al. prepared transdermal patches by solvent casting technique from the combination of poly(vinyl pyrrolidones), propylene glycol, and glycerin followed by loading with chlorhexidine gluconate, an antibacterial agent. The % drug content of the patches was in the range of 94%–100%. The tensile strength of the patches was in the range of 0.492–0.681 kg/cm2 revealing their good flexibility and mechanical strength. The use of poly(vinyl pyrrolidones) decreased the percent elongation and tensile strength of the patches. The presence of poly(vinyl pyrrolidones) made the patches permeable to the drug and increased their porosity. The drug release was enhanced due to increased porosity of the patches. The patches exhibited good antibacterial activity [94]. Jiang et al. developed a cost-effective, flexible, and passive system for topical drug delivery in which the wound pH-triggered drug release. The dermal patches swelling at alkaline pH of an infected wound released the drug at a slow rate of ( > 0 > > > Ac ðt Þ ¼ ka  Aa ðt Þ  ke  Ac ðt Þ > < Ac ðt Þ Cc ðt Þ ¼ > Vc > > > > > A ð 0 Þ ¼ F  Dose a > > > : Ac ð0Þ ¼ 0

(1)

where Ac is the amount of drug in the central compartment, Aa is the amount of administered bioavailable drug, ke is the elimination rate constant, ka is the absorption rate, and F is the fraction of bioavailable drug. The concentration in the central compartment, Cc is obtained dividing the amount Ac by the volume of distribution in the central compartment, Vc. Following this model, Cc can be expressed as: Cc ðt Þ ¼

  F  Dose ka   eke  t  eka  t Vc ka  ke

(2)

For the 2-compartment model, the initial conditions are as follows: 8 0 Aa ðt Þ ¼ ka  Aa ðt Þ > > > > > > A0 ðt Þ ¼ ka  Aa ðt Þ  ke  Ac ðt Þ + kr  Ap ðt Þ > > c > > 0 > > > Ap ðt Þ ¼ kd  Ac ðt Þ  kr  Ap ðt Þ > > < Ac Cc ðt Þ ¼ > Vc > > > > > A ð 0 Þ ¼ F  Dose > a > > > > > Ac ð0Þ ¼ 0 > > > : Ap ð0Þ ¼ 0

(3)

where Ap is the amount of drug in the peripheral compartment, kd and kr are the distribution and redistribution rate constants, respectively. The concentration in the central compartment is described by the equation: Cc ðt Þ ¼ α1  eβ1  t + α2  eβ2 t  ðα1 + α2 Þ  eka  t

(4)

where α1, α2, β1, and β2 are auxiliary terms defined as follows: 8 Dose  F ka  ðkr  β1 Þ > > α1 ¼  > > Vc ðka  β1 Þðβ1  β2 Þ > > > > > Dose  F ka  ðkr  β2 Þ > > >  < α2 ¼ V ðka  β2 Þðβ1  β2 Þ c  qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi > 1 > > β1 ¼  ke + kd + kr + ðke + kd + kr Þ2  4kr ke > > 2 > > >  qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi > > 1 > > : β2 ¼  ke + kd + kr  ðke + kd + kr Þ2  4kr ke 2

(5)

In 2-compartment models, the transfer from the central to the peripheral compartment can be quite rapid, leading to a lower Cmax than that of a 1-compartment model (Fig. 3). Usually the elimination in a 2-compartment model ends up being slower, especially if redistribution rate constant (kr) is low because the drug needs to leave the peripheral compartment before being eliminated. To the above equations for 1- and 2-compartment models, t can be replaced by (t  tlag), with tlag being an additional variable that accounts for an eventual delay in the start of drug absorption.

4.3. Application of Pharmacokinetic Models to a Prototypical Polyphenol Epicatechin is a prototypical green tea polyphenol, catechin-type flavonoid and bioactive natural compound widely present in human diet. Epidemiological data and varied laboratory studies reinforce important health beneficial effects of epicatechin, which have been associated to multitarget mechanisms of action [54–56, 65]. Unfortunately, most human trials of tea polyphenols formulations are returning only partially positive outcomes [48, 66], so there is a great interest in developing more effective delivery systems for these compounds [2,8–10,67]. The PK of epicatechin and other catechins has been studied previously in humans by Lee et al. [68]. In this study, 20 mg/kg of green tea solids dissolved in 200 mL of water was orally administered to 8 healthy individuals after overnight fasting, and the plasma concentration of catechins was followed for 24 h [68]. The authors repeated the trial in three separate occasions with the same 8 individuals and each trial was identified as GT1, GT2, and GT3. Then the PK data were analyzed by a 1-compartment model and the study reported the values for Cmax, Tmax, AUC, and t1/2 in Table 2. These are the descriptive parameters routinely reported in PK and bioavailability studies but for guidance in the development of drug delivery systems, more mechanistically relevant parameters as the absorption rate (ka) are needed. To obtain additional PK parameters from the data by Lee et al. [68], we collected the average PK profiles of

CHAPTER 8

135

Development of Delivery Systems and Pharmacokinetics

TABLE 2

Pharmacokinetic parameters of epicatechin in humans obtained by fitting of 1-compartment model to experimental data in Lee et al. [68]. The parameters from the original publication are averages of fits for data from each individual with standard deviations in parenthesis. CMAX (ng/ML)

Lee et al. [68]

Our analysis

TMAX (h)

Lee et al. [68]

Our analysis

AUC (ng*H/ML)

Lee et al. [68]

Our analysis

ka (h21)

ke (h21)

Our analysis

Our analysis

Our analysis

t1/2 (h)

Lee et al. [68]

GT1

120.8 (67.6)

110.5

1.26 (0.41)

1.19

436.5 (157.5)

400.3

1.53 (0.73)

1.43

1.386

0.485

GT2

133.0 (67.6)

127.7

1.35 (0.34)

1.25

593.8 (286.9)

563.8

2.05 (1.01)

2.01

1.616

0.344

GT3

118.3 (73.7)

110.7

1.18 (0.56)

1.19

558.2 (288.8)

509.5

2.52 (1.95)

2.22

1.839

0.312

plasma epicatechin concentration for each trial described in the original publication and carried out our own analysis using 1- and 2-compartment models. The two models were compared using the adjusted coefficient of determination (R2adj) and the 1-compartment model performed better in all situations (GT1, GT2, and GT3), so the 2-compartment model was discarded. The values of the various PK parameters obtained using our 1-compartment model are similar to those obtained by Lee et al. [68] (Table 2). The 1-compartment model was fitted to average PK curves, instead of the individual PK curves used by the authors in their analysis (curves not available), which justifies the small differences in parameter values between the two analysis. Despite this difference, all values we obtained are well within the standard deviation intervals presented by Lee et al. [68]. Using our compartmental analysis, the kinetic values of absorption (ka) and elimination (ke) were additionally obtained (Table 2). The average values for these parameters from the 3 trials are 1.61 h1 and 0.38 h1, for ka and ke rates, respectively, and can be useful to develop delivery systems with release rates in accordance.

5. STUDY MODELS AND APPLICATION IN DERMAL DELIVERY To develop drug delivery systems with suitable release kinetics, relevant study models are necessary for testing and optimization. Franz cells and Transwell systems are setups that can be used to mimic different drug absorption conditions and enable in vitro testing of pharmacological activities under variable concentration-time profiles.

5.1. Studies in Franz Cell 5.1.1. Drug permeation studies A Franz cell is an apparatus that consists of two chambers (donor and receptor), between which a synthetic membrane or an ex vivo biological membrane (e.g., a dermatomed skin sample) can be fixed (Fig. 4A). Placing a liquid or emulsified test compound in the donor chamber allows the study of the permeation through the membrane by sampling and measuring the compound concentration of the receptor fluid over time. A small stir bar keeps the receptor media under constant magnetic agitation. Franz cells are routinely used not only to characterize skin permeability of compounds but also to study drugloaded films and other vehicles intended for mucosal (buccal, intranasal, and intestinal) delivery [2, 7, 8, 18]. In conditions of transdermal and transmucosal delivery, the kinetics of drug absorption will depend on the drug release properties of the delivery system and the permeability properties of the biological barrier [31, 37]. Some synthetic membranes have been used for mimicking skin permeability, namely Strat-M and silicone membranes. Uchida et al. [69] proposed a silicone membrane showing validity as model of human skin permeation for different compounds [69]. We have also been using this membrane and Fig. 4B shows permeation profiles for caffeine. From the experimental data, the permeability coefficient (P), the diffusion parameter (DL2), and partition parameter (KL) can be calculated according to Eq. (6). 8 Flux ¼ P  Cv > > < 1 DL2 ¼ 6  Tlag > > : KL ¼ 6  Tlag  P

(6)

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Modeling and Control of Drug Delivery Systems

FIG. 4 Franz cell setup for drug permeation and release studies. (A) Franz cell and its components;

(B) permeation of caffeine through a silicone membrane from Lintec Co., using donor saturated solutions in a Franz cell with effective permeation area of 0.636 cm2 and a receptor volume of 5 mL saline (NaCl 0.9% w/v).

where the P is calculated from the Flux, which is the slope of the linear phase of the permeation profile divided by the effective permeation area, and from the concentration of the donor solution (Cv). The Tlag can be obtained from the intersection of the linear trend line of the linear phase with the x-axis, and is used to calculate the diffusion (DL2) and partition parameters (KL). Caffeine could permeate the silicone membrane and the acidification of the saturated caffeine solution did not alter the permeation kinetics (Fig. 4B), suggesting that mild acid pH does not influence the permeability of the membrane. Caffeine is used as a model of positive drug in permeation devices [40]. From the kinetic profile the permeability parameters P, DL2, and KL were calculated and are shown in Table 3. The obtained values were similar to the results by Uchida et al. [69] and provide a good reproduction of caffeine permeation (P) through skin, encouraging the use of Franz cells with

appropriate membranes as tool to investigate the transdermal administration and absorption kinetics of bioactive compounds. Ointments, gels, and emulsion systems can be directly tested in this setup.

5.1.2. Drug release studies As indicated in the previous section, Franz cells can be valuable to study film-type or other drug release systems developed for dermal (skin patches) and mucosal delivery. For these applications, we choose alginate as support because this polysaccharide is well accepted for biomaterials and it allows the amenable encapsulation of bioactive compounds in different forms [9, 16, 67]. Moreover, alginate-based materials are already used in dermal applications and are proposed for colon-targeted delivery [2, 13]. Using alginatecontaining carriers, curcumin release was increased by acid pH in the case of chitosan/alginate NPs, and by

TABLE 3

Permeability parameters of caffeine through silicone membrane and skin. Values of permeability coefficient (P), diffusion parameter (DL2), and partition parameter (KL) are compared with the reported in the literature [69]. The literature values of KL and DL2 were estimated as an interval from a graphical representation of the results. Swiss ADME is a free web tool for pharmacokinetic properties of compounds [70] using QSPR model to calculate skin permeation [71]. P (cm∙s21) 7

DL22 (s21 )

KL (cm) 1.6104

Our work

1.810

1.3103

Silicone membrane [69]

2.2107

[1–10]103

1.0104

7

5

Human skin [69]

1.810

[1–10]10

1.1102

Swiss ADME

3.0108

—

—

CHAPTER 8

Development of Delivery Systems and Pharmacokinetics

alkaline pH in nanoemulsion-filled alginate beads [43]. A study with human volunteers showed that alginatecoated gelatin capsules can withstand the stomach environment and then migrate to the ileocecal region of the intestine where they disintegrate [4]. Following the hypothesis that a drug release system enabling kinetics of drug delivery similar to the rate of drug absorption (ka) observed in PK studies will favor therapeutic efficiency, it is important to characterize the release kinetics of the systems under development.

137

In the case of the polyphenol epicatechin, PK analysis after oral administration pointed to an average ka value of 1.61 h1 (Section 4.3). It should be noticed that increased absorption promoted with catechin-rich oral formulations, especially when taken under fasting conditions, may trigger adverse reactions [72]. Fig. 5 displays kinetic curves of epicatechin release by alginate films measured in a Franz cell (as that represented in Fig. 4A). Results for alginate films with and without glycerol (plasticizer) are shown to illustrate

FIG. 5 Release of epicatechin by calcium alginate (A) and glycerol (30% w/w)-supplemented calcium alginate

(B) films in Franz cell (details in Fig. 4). Lines are the nonlinear regression fits of the experimental data to firstorder models, and the parameters from Weibull model are also presented for comparison. In the release assays, alginate films were placed between the donor and receptor chambers, the receptor chamber was filled with 5 mL of saline ensuring contact with the film and 30 μL of the receptor liquid was sampled through the sampling port and replenished with saline at various time intervals. The polyphenol concentration in the samples was measured by HPLC.

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Modeling and Control of Drug Delivery Systems

FIG. 6 Assay of epicatechin delivery by calcium alginate film on ex vivo porcine skin. After 1 h release, skin

surface is sampled and analyzed by HPLC (absorbance signal at 278 nm). The chromatogram of an epicatechin standard with 50 μM concentration is shown for comparison.

how modifications in the delivery system can be used to tune release kinetics. Epicatechin was encapsulated into these films during gelation with calcium chloride, methods detailed in [67]. The experimental data of drug release in Franz cells can be analyzed using diverse modeling approaches, including kinetic models to compare with PK compartmental models. The kinetics of epicatechin release was fitted to a first-order model (Eq. 7) and the Weibull model (Eq. 8):   C ¼ C∞ 1  ekt

(7)

  b C ¼ C∞ 1  eat

(8)

where k is the rate constant, and a and b are the Weibull constants. The first-order kinetics model, while not describing the physiochemical phenomena of release, allows a direct comparison with the transfer rates (absorption and elimination) of PK compartmental models (Section 4). As depicted in Fig. 5, the first-order model fits quite well the epicatechin release data and looking at the a and b values from the Weibull model (contains more independent variables) it converges approximately to a first-order kinetics. Remarkably, the release rate of epicatechin from alginate film (Fig. 5A) is similar to the rate of absorption of the polyphenol in humans (approximately 1.6 h1). Furthermore, the release rate of glycerol-supplemented film (Fig. 5B) is slower than of unsupplemented film

but afforded higher concentrations of epicatechin in the release medium probably because it accomplishes a higher encapsulation efficiency. In both cases, these release systems afforded micromolar concentrations of the tea catechin (Fig. 5), which were further confirmed in assays with ex vivo porcine skin (Fig. 6). Envisaging a dermal application of such carriers, these micromolar levels of tea catechins at skin surface are compatible with simulations of high anti-melanoma bioactivity [31]. Different tools are emerging to simulate drug release and absorption in vitro [29, 32, 39, 40]. In alternative to the standard agitation flask assays for drug release studies, the Franz cell reveals a useful laboratory model to test and optimize drug release systems in conditions that approximate physiological delivery, that is, temporal profile of drug presentation. For bioactive products with the particularities discussed in Section 3, it can be anticipated that the testing conditions become even more critical to successfully develop drug delivery systems with enhanced therapeutic efficiency.

5.2. Studies in Transwell System Transwell is an innovative but relatively easy-to-use setup that allows researchers to add complexity in a cell culture system when compared with normal cell culture methods. Some examples using Transwell culture systems include co-culture models [73], 3D cell printing [74] and in vitro invasion assays [75]. Interestingly,

CHAPTER 8

Development of Delivery Systems and Pharmacokinetics

additional applications include the isolation of cancer stem cell [76] and permeability studies of blood-brain barrier models in the presence of different compounds [77]. When used for drug delivery assays, the Transwell system can be compared with the Franz cell with a donor and a receptor chamber (or compartment) separated by a membrane (Fig. 7). The definition of donor and receptor (or acceptor) chamber will always depend on which compartment the drug will be initially released as the drug diffusion can be studied in both directions, that is, from upper to lower compartment or vice-versa depending on the objective of the study. Transwell systems are available with different membrane’s pore diameter that, combined with other modifications, can be explored to achieve a desired in vitro drug behavior

139

[29]. In this line, the possibility to mimic the PK properties of bioactive compounds in Transwell culture systems is very attractive for investigative delivery studies. This strategy would enable to deepen the pharmacological activities of natural compounds by exposing cells to different concentration-time profiles (Fig. 7). Polysaccharide films loaded with epicatechin (discussed in section 5.1.2) put in a Transwell system slowly released the polyphenol (representative curve in Fig. 8A), with a kinetic rate inferior to the ka absorption rates observed in PK studies (>1 h1). In addition to films, particulates and delivery systems with other geometries can be used in Transwells. We used dry alginate spheres prepared by the method described in [78], loaded with epicatechin and placed in the donor compartment. Monitoring the cumulative

FIG. 7 Schematic representation of the Transwell setup for drug release studies. The delivery system is placed

in the donor chamber (or compartment) and the drug concentration monitored in the receptor (or acceptor) chamber. Cells can be cultured in the system and exposed to a concentration-time profile that mimics the physiological temporal presentation of the drug.

FIG. 8 Time evolution of the epicatechin concentration in the acceptor compartment of a Transwell using

calcium alginate films (A) and spheres (B) loaded with the polyphenol in the donor compartment. A 24-mm Transwell system with a polyethylene terephthalate membrane insert was used. The release medium in the compartments was saline and epicatechin concentration was analyzed by HPLC.

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Modeling and Control of Drug Delivery Systems

epicatechin concentration in the acceptor (Fig. 8B), the loaded spheres enabled a kinetic rate (0.9 h1) more similar to the absorption ka rate observed in humans. The assays described here show the suitability of Transwell systems to screen different delivery systems, including those with different morphologies and identify the ones presenting more suitable release properties. Moreover, since Transwells are already well-established cell culture supports for diverse functional studies, the possibility to generate different concentration-time profiles with varying drug delivery kinetics offers a useful tool to investigate novel features of the compounds’ bioactivity in cells.

A combination of physiology-based modeling, biomaterial design and in vitro screening tools are becoming available for the development of novel drug formulations. An overall goal of future research is to assess the therapeutic value of these promising approaches and the derived products in preclinical and clinical studies.

ACKNOWLEDGMENTS The authors acknowledge the support by “Fundac¸ão para a Ci^encia e Tecnologia” (FCT – Portugal) through the research project PTDC/BIA-MIB/31864/2017. The silicone membrane used in the drug permeation studies was a gift from Lintec Company.

6. CONCLUSIONS Natural bioactive compounds as polyphenol antioxidants ask for more effective delivery approaches that could increase clinical application of new therapies. Epidemiological and laboratory studies give support to the therapeutic potential of several of these compounds, but only in some cases human trials are returning the necessary efficacy for clinical approval, typically with the more specific and potent compounds as alkaloids or capsaicin. Compounds naturally appearing in human diet and metabolism present important pharmacological features that can critically limit their therapeutic outcome. Low bioavailability, multitarget and low-potency action mechanisms make natural compounds-based treatments highly dependent on PK factors, administration dose and time schedule, absorption/excretion rates or cell time-changing exposition. It is proposed PK parameters to be best accounted as possible in the design of delivery systems for natural bioactive compounds. Doses identified by epidemiological studies and existing clinical trials as affording beneficial effects should be analyzed to plan therapeutic approaches. PK data obtained in humans in these favorable conditions offer insights into the dynamics of absorption of the compounds and their course in the human body that may be the key to define successful drug delivery schemes. In addition to usual goals of high loading and controlled release, a PK-guided regulation of temporal drug presentation in the development of delivery systems is hypothesized to lead to more effective treatments by natural compounds. Laboratory testing setups compatible with different delivery systems are presented herein, accompanied by illustrative examples of validation and application procedures.

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[42]

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CHAPTER 9

Nanofiber: An Immerging Novel Drug Delivery System DIPAK KUMAR SAHU • GOUTAM GHOSH • GOUTAM RATH

School of Pharmaceutical Sciences, Siksha ’O’ Anusandhan (Deemed to be University), Bhubaneswar, Odisha, India

1. INTRODUCTION In anticipation of the polymeric biomaterials, as a weapon for the chronic disease’s treatments, progress is continuous as more pharmaceutical investments are pouring in, commensurate with the upward trend of the number of patients. The leading cause of deaths, such as cancer, diabetes, and cardiovascular disorders, as per the WHO, are projected to share half of the mortality by 2035. Hence, as pharmaceutical scientists, effective targeted drug-delivery strategies need to be continuously explored. It is from experimental outcomes known that nano delivery systems have a huge advantage of crossing the intricate physiological and anatomical barrier and can release the actives for months, whereas the initial exposure to the target through a burst release is enough to suppress the progression. But several challenges remain in terms of in vitro correlation to the simulation of auto biodegradation kinetics in the body and reproducibility of the results. The scaling up issues due to lack of identification of optimization parameters common to both industry and laboratory are being looked after to consider an ideal drug-delivery system. An emerging drug delivery system in recent time furnishing the requirements and searched for is nanofiber that has a history of centuries being intermittently studied for fabrications and applications. But their exploitation for drug delivery and tissue engineering for them has gained the impetus, which shows an analysis of the bulk of the literature of 85% having the 23,728 articles got published just from 2006 to 2018. As more institutions are conducting the experiments to gear them up for the industry, there are substantial improvements both in terms of technology of fabrication and delivery optimization because of their flexibility to modulate without losing the inherent properties [1]. Nanofibers have morphological and structural similarity with the extracellular matrix that has categorical

advantages with each attribute such as porosity, mechanical, and biochemical similarity. The porosity is due to the polymeric fabric skeleton that is ultrafine, allows entrapping the drug with high encapsulation efficiency. The endurance strength is due to different physical and chemical interactions such as electrostatic, covalent either present on drug or polymer surface. Sometimes, surface modification through plasma treatment, chemical, or graft polymerization, surface-active reagents introduced artificially help in creating a design construct conducive to degradation in the biological environment. The liberty of mixing the synthetic and natural polymers widens the employability, where the individual demerits can be circumvented for the final construct. The superior fine fiber can accommodate hydrophilic, lipophilic, and amphiphilic drugs alone or in combination. Accordingly, different spinning techniques including uniaxial, biaxial, triaxial, and coaxial variations can be used to prepare nanofibers as per the requirement of specific medical conditions. The drug release profiles without overlapping may be slower or faster as the erosion process starts. Multiple drugs within the polymer matrix each at different layers give the opportunity to bypass unintended interactions among them, and their orderly release helps for synergistic therapeutic activities. In the electrospinning process, the parameters optimized that influence diameter and the interfibrillar distance are the concentration of both drug and polymer, molecular weight, the voltage for the Taylor cone, and, finally, the distance at which the collector plate is placed. The advanced analysis with the response surface methodology (Box Behnken design) for factorization and linking of the thermodynamic parameters, such as enthalpy, entropy, free energy equations with the kinetics data, simplifies the ideal conditions for the electrospinning process [2]. There has been a great amount of effort to avoid toxic solvents during the electrospinning process. In doing so, melt

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electrospinning and solution electrospinning are the other methods that gained popularity, but the melt method stands out due to better fiber structure [3]. A melt electrospinning prototype with 600 nozzles gives an evenly distributed fiber as the output and can be observed with a composite of Poly lactic acid (PLA) with 6% w/w sodium stearate averaging 1000–7000 μm of nanofiber by the above process [4]. The centrifugal spinning production method gives 500 times more productivity, but the issues remaining are the polymer solution elasticity and evaporation rate. Both melt blowing (MB) and melt extrusion can be coupled, then a high-speed gas stream directed toward it can produce industrial throughput nanofiber. Carvedilol with hydrophilic polyvinylpyrrolidone-vinyl acetate copolymer (PVPVA64) and polyethylene glycol (PEG 3000) plasticizer gave an ultra-fast drug delivery system by the MB process, which is further due to the large surface area when compared with the ES process [5]. Similarly, for alternating current electrospinning in the place of direct current electrospinning for hydroxy propyl methyl cellulose (HPMC), solid dispersions give more productivity [2]. Nozzle-free electrospinning method increases the dissolution rate and bioavailability. Niflumic acid with PVP nanofiber in the electrospinning without nozzle increases a 15-fold increase in dissolved amount due to the amorphous form conversion of the drug in the web of the nanofiber [6]. Albendazole in poly vinyl alcohol (PVA) shows a fivefold increase in dissolution and a 3-fold increase in permeation due to the above needle-free process [7]. Other polymers such as HPMC, PVP, sucralose, and mannitol can include different electrospinning process to accelerate the dissolution for poorly soluble drugs [8]. Needle-clogging problem and organic solvents inclusion can be avoided with ultrasound enhanced electrospinning. The waves create heat for the fast evaporation that creates amorphous drug in the network. The nanofibrous mats have high water adsorption potential of 400%–1600% by weight and flexibility of pore diameter as 10–1000 nm; hence, they act as a reservoir for liposome, dendrimers, micelles, antigens, and allergens that are deposited either between the pores or on the surface to give a prolonged-released system. For instance, D-α-tocopheryl PEG 1000 succinate (TPGS) with PF127 solution embedded with cyclosporine A as a micellar system in PVP nanofiber gives a better dissolution profile [9]. Due to the mucoadhesive nature of natural polymers, they can be a sustained delivery system for the sublingual and buccal tissue. The release patterns of the matrix system are best described as immediate, slow,

or prolonged, biphasic patterns. The aligned fiber has more control over the burst release. Biphasic or pulsatile drug release can be designed by the layered form of the drugs-polymer complex. With the thermo-responsive (Ex-cross-linked poly (N-isopropylacrylamide)) temperature and pH (Ex-poly(N-vinylcaprolactum) or ethyl cellulose) and Eudragit L100 or magneto-responsive polymers (Ex-Fe3O4 particles and PVP polymer) give additional benefits of drug release on demand. The current chapter is aimed at giving succinct compilations from various domains in which they have been studied [10–12]. A short compilation of different actives when formulated with nanofiber can show revolutionary therapeutic activities as given in Table 1.

2. NANOFIBER AS PROLONGED-DRUG DELIVERY SYSTEM 2.1. Wound Healing A successful wound-healing material should have biomechanical features that in the natural environment of injuries assert reconstructing the healthy network of the skin layers [25]. But several drug delivery systems such as a hydrogel or a matrix substitute are yet to give the full functionality due to the faulty architecture of them, giving a delayed, uneconomical, and worrisome burden to the patients. Nanofiber, on the other hand, provides a tunable mechanical strength fitting within the range of 15–150 MPa tensile modulus of the human skin, and the required ECM components can be electrospun into them without losing their activities, providing healing results better than the existing models. There are a plenty of therapeutic agents explored with a range of nanofibers for antibacterial, pain management, homeostasis, antioxidant, angiogenesis, reepithelization, and adhesion with proliferating activities. The antibiotics, for example, cefoxitin sodium, gentamicin, ciprofloxacin, pain killers such as lidocaine and mupirocin with poly lactic-co-glycolic acid (PLGA), chitosan, PVA, PLLA, polyurethane (PU), or dextran, were experimented and have shown considerable clinical potentials. Various growth factors such as VEGF, PDGF, and EGF could be combined to enhance the process faster. They could make their ways to the clinics in the form of sutures occlusive or semiocclusive dressing materials. Essential oils such as thymol, rosemary, and geraniol with their antiinflammatory, antifungal, and antibacterial properties can be formulated with alginate and chitosan nanofibrous polymer to promote faster wound closing, fibroblast proliferation, and collagen synthesis [12, 13, 26, 27].

TABLE 1

Polymeric nanofibers for drug delivery applications. Polymer matrix

Drugloading efficiency

Drug release

In vitro or in vivo model

Drug

Category

Highlights

Refs.?

Clotrimazole

Antifungal

PVA or Dextran or alginate

–

30% with burst release and up to 98% within 30 min

Franz’s cell

Controlled release for 30 min with zero-order and no cytotoxicity

[13]

Voriconazole

Antifungal

PVA or sodium alginate with cross-linker GTA

96.45% 5.91%

50% within 30 min and 96% for 8 h

Franz’s diffusion cell and pig dorsal skin

Increased VCZ deposition in the deeper skin layer when compared with 1% w/v VCZ solution in PG

[14]

Griffithsin

HIV-1 and HSV-2

PCL or PEO or mPEG-PLGA GRFT NPs

85.6% 11.0% maximum

Burst release through 72 h and sustained release up to 90 days

Female balb/c mice

Both gel and insert show the GRFT release for 90 days and found to be safe in in-vivo studies

[15]

Dacarbazine

Anticancer

PVA

83.9% 6.5%

Burst release after 30 min (58% in pH 6.8 and 335 in pH 7.4, followed by sustained release for 3 days

–

Prolonged release with improved antitumor effects against the hydrogel

[16]

DOX/PTX/5-FU

Anticancer

PLA or chitosan

Around 95%

Different in single and tri-layer nanofiber up to 700 h

–

High activity of the tri-layer nanofiber for MCF-7 breast cancer cells (94% killed)

[17]

Pilocarpine

Cholinergic

PLGA or PEG

Decreased ADR

[18]

BCNA or Irinotecan or cisplatin

Anticancer

PLGA

Wistar rats

Multiagent release over 8 weeks without any inflammatory reaction

[19]

Vitamin B12

Antianginal

HA or PVA

97%

Burst release of 28% after 30 min and then up to 6 h of steady release

Wistar rats and goat mucosa

No evidence of mucosal ulceration after Nicorandil administration

[20]

Silver sulfadizine (AgSD)

Antibacterial

Zein

–

9.54 ppm to 18.05 ppm Ag ion for 0.3% to 0.5% AgSD in 70 h

–

0.6% AgSD antibacterial for both gram positive and gram negative bacteria

[21]

Methylprednisolone

Hormone

ES100

–

Different formulation ratios up to 11 h possible

–

ES-100 to MP ratios of 10:0.5 have a potential for the delayed release

[22]

Vitamin E

Antioxidant

HPβCD

11% w/w

–

–

Improved water solubility, increased shelf-life, and ultraviolet light stability of vitamin E

[23]

Propolis

Antiinflammatory

Zein

Efficient antinflammatory compound for oral use

[24]

26% in 4.5 h and 46% in 15 days

BCNA, Crusting; ES-100, Eudragit S100; HPβCD, hydroxypropyl-β-cyclodextrin; AgSD, silver sulfadizine.

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2.2. Cancer The usual practice to employ antineoplastic drugs for cancer is by a combination which has different modes of action. But at the same time, the debilitating adverse effects are combined too. To counteract the unwanted effects, the cocktail of drugs can be entrapped in cargos that can deliver slowly to tumorigenic cells with maximum therapeutic effects [28, 29]. Nanofiber produced by the emulsion and coaxial electrospinning have the potential to embrace two or more drugs. The in vitro results are supported by the expansion of the scope of treatments in the laboratory. Through time-lapse videomicroscopy, their penetration can be seen in the tumor cell [30]. The 2D cell culture does not improve the therapeutic paradigm of cancer treatment; hence, 3D scaffolds are designed to bridge the gap between the in vitro and in vivo. For instance, bacterial cellulose nanocrystals having PCL and gelatin help in neurite growth, and the compatibility study with U251 GBM cells shows that they have improved adhesion and proliferation [28]. The PLLA or PF127 nanofiber with Nbutylpyridoquinoxaline1,4-dioxide-NBPQD or 2-amino-3-cyano-6-methylquinoxaline 1,4-dioxide has higher cytotoxic activity when compared with the free drug through up regulation of p53 and p21 apoptotic signaling pathways [31]. Tri-layer nanofibers synthesized through the triaxial electrospinning process have drugs Doxorubicin (DOX), Paclitaxel (PTX) and 5-fluorouracil (5FU) all together at different layers. The inner layer was fabricated with 5FU in CS/PVA. The intermediate layer as PLA/CS and outer layer as DOX and PTX molecules in initial loading the g-C3N4 nanosheets, followed by incorporation in PLA/CS [17]. The investigations as a comparison in a single layer of the individual drug when compared with the trilayer reveal that the triple-layered drug was having more efficiency against MCF-7 breast cancer cells due to the sustained release. The lower stability and poor pharmacokinetic profile of soluble tumor necrosis factor-related apoptosis-inducing ligand (sTRAIL) can be improved by combining genetic engineering with coaxial electrospinning. sTRAIL was modified with adenovirus knobless fiber than in PLGA nanofiber that give excellent biological activity by inhibiting the breast tumor cancer cells [32]. Aligned fibers have similarities with the brain ECM, and chemotherapeutics, with effective concentration, can overcome the BBB. For brain tumor, PTX in PLGA against glioma C6 cells release for 80 days and TMZ in PLGA, PLA, and PCL show an increase in median survival time, greater in 4 monthsreleasing nanofiber than 7 days survival time nanofiber [33, 34]. A magnetic fibrous membrane made up of PU

nanofibers studded with superparamagnetic γ-Fe2O3 nanoparticles can have the advantages of heat generation in response to the alternating magnetic field that can be used as magnetic hyperthermia of cancer therapy [35].

2.3. Microbial Diseases Vaginal drug delivery systems in the form of a gel giving a messy consistency and the tablet giving a grainy residue are the problems to encounter; hence, Azidothymidine, Acyclovir, and Maraviroc in a blend of PLLA/PEO give the 100% release in 1 h [36]. Vulvovaginal candidiasis occurs in pregnant women and immunocompromized patients. Mucoadhesive formulation designed with dextran and alginate for the drug clotrimazole obviates the need for adding a plasticizer, gives faster, and double antifungal properties [13]. A concentric cylindrical capsule with an inner diameter of 6 mm and length of 15 mm can be designed with the ES technique having PVP middle layer and PCL outer layers that can separate to two layers after the PVP dissolution [37]. Nanofiber mucoadhesive patches of clotrimazole for oral candidiasis can be designed in three layers with outer HA or catechol, middle PVP with HPβCD, and CT that show enhanced antifungal activity when compared with CT powder. Voriconazole can be formulated with SA and PVA that show slow release with the crosslinker GTA described as the Higuchi’s square root kinetic model [38]. Similarly, ungeremine (alkaloid) in a blend of PLA/PEG and cinnamon oil in PVA as a core-sheath structure by emulsion electrospinning gives enhanced antifungal properties [39, 40]. Rifampicin in PCL/HA nanofiber cause 3 log reduction in bacteria S. aureus, methicillin-resistant S. aureus (MRSA), and S. epidermis and 2 log reduction in P. aeroginosa bacteria [41]. Through physical contact only, ESPAN nanofibers have both antiyeast and antifungal activities [42]. Gold nanorods in PVA/CS hybrid nanofiber have the functionalities of inhibiting the ovarian cancer cell without interfering with the structural integrity of the nanofibers [43]. Curcumin in poly 2-hydroxy ethyl methacrylate nanofiber against MRSA and ESBL can be useful for the MDR bacteria. For the brain infection, as a postoperative procedure, 1–2 months of parental antibiotic treatment is required causing a high amount of toxicity [44]. A biodegradable nanofibrous membrane can be made up of PLGA in which vancomycin can be loaded, and after placing in the cerebral cavity of the rat, the release was found to be up to 80 days [45].

2.4. Cardiovascular Diseases Cardiovascular injuries have the highest mortality, and transplantation is the least invasive method; hence,

CHAPTER 9 Nanofiber: An Immerging Novel Drug Delivery System regenerative cardiology has a great impact on the healthcare system. Self-assembling peptide sequences of various lengths are used for the delivery of the therapeutics to the myocardium in case of the infarction. The advantages are the specific molecular interactions that may be electrostatic, hydrophobic, and van der Waal or because of hydrogen bonding. The sequences evolve into a 3D assembly that can deliver small molecules such as NO. Nanofibrous scaffold formed by the addition of the catalyst such as monovalent salt, where a change in pH and adaptation of the β sheet, β hairpin structure, accommodate VEGF, heparin-binding domain, PDGF, etc., as to regenerate the myocytes. The injectable form of them targeted to the infarction has shown improved vascularization, diminishing size of the dead zone with reperfusion, and improved function. Oral administration of Nicorandil for angina lead to mucosal ulceration; hence, a composite nanofiber construct comprising of HA, PVA, and Vitamin B12 can be designed to counteract the adverse consequences. The pharmacological profile of the drug Carvedilol increases with the change in crystalline to amorphous form, during the electro spraying process of nanofiber formulation of it along with Eudragit® RS100 [1, 20, 46, 47].

2.5. Macromolecules for Miscellaneous Applications The intrinsic properties of therapeutics that are proteinaceous make them suitably integrated into the biochemical environment; hence, the maximum therapeutic efficiency is achieved. But the formulations are facing extreme challenges in terms of stability due to the fragile nature. Polymers are vastly explored to maintain the bioactivity to the endpoint user. Proteins conceal themselves by entrapping in the polymeric matrix through weak interacting bonds to get immobilized or to get bioconjugated. Different nano polymeric delivery systems alternate to oral or intravenous, such as implants and microneedle, are at the clinical bench now. Still, the invasiveness and patient compliance are at the stress. Both plant- and animal-derived natural or synthetic peptides can be mixed with PLA, PEO, and Eudragit® with the help of cross-linkers to make protein-based nanofibers. There are several advantages, such as stability, less prone to degradation, biological function enhancement, and facilitation of electrospinning as desired in the process. For instance, soy protein, zein, silk fibroin, gelatin, keratin, fibrinogen-based nanofiber with cross-linkers such as glutaraldehyde, genipin, and glyoxalin have given significant results at the preclinical level. The delivery of plasmid DNA, siRNA, and

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autologous stem cells with nanofibers can be customized easily. Mesoporous silica nanoparticle (MSN) for the bolus delivery of siRNA on surface adsorbed as siRNA-MSN complex on poly (etherimide) nanofiber give a sustained release up to 5 months as gene-silencing effects [48]. Highly effective vaccines are lacking for deadly diseases such as cancer, influenza, HIV, and tuberculosis due to different immune responses. Through nanofiber, an effective host response can be predicted. Peptide nanofiber that is conjugated to a short peptide may function as an epitope that further can be stabilized with adjuvants for a modulated immune response [49, 50]. For example, the hepatitis B vaccine in PLA or PLGA gives a robust and specific immune response. Other natural polymers, such as chitosan, cellulose, hyaluronic acid, and zein were studied with quercetin, vitamin E, curcumin, rutin, and porphyrin and have shown an increased bioavailability in in vivo tests. Amorphous lyoprotectant trehalose for the long-term viability at room temperature for 24 weeks of L. plantarum ATCC 8014 with a loading of 7.6  108 with PEO nanofiber can give a complete release over 30 min. This can be useful for the reestablishment of microbiota balance in the vagina [51]. Nanofibers by taking the base of Aloe vera can preserve the woundhealing properties for the diabetic ulcer [52]. A fusion of HPV16 E744-62 vaccine to self-assembling peptide Q11 in nanofiber matrix can be used for the immunization for human papilloma virus by inciting greater immunogenic response. Xerostomia is the clinical condition raised by polypharmacy in elderly people. The frequent systemic Pilocarpine has many ADRs, and also the efficacy is limited. However, Pilocarpine with PLGA or PEG nanofibers support a sustained delivery and the salivary gland growth. MP can cause the ulcer to the gastric mucosal lining so as to prevent the damage. Enteric polymers, such as Eudragit L100-55, Eudragit S100, and Kollicoat MAE100-55 can be taken for nanofibers preparation out of which the selected polymer Eudragit S100 with the drug as 10:0.5 ratio showed that it was the best-optimized candidate for delayed release profile. The Weibull’s model was established as a complex release phenomenon. Recombinant growth hormone needs frequent injections for short in vivo life span. By stabilizing with sugar glass nanoparticle, then electrospun in polyester urea, nanofibers give a sustained release over 6 weeks [53]. Through electrohydrodynamic techniques, bioactive molecules such as DNA and growth factors can be delivered and that is also a suitable process to get nanofibers with multiple agents [54]. This has a high loading efficacy with the maintenance of topographical features when compared with

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plain electrospinning. Patient-oriented customized medications are achievable through the above method. An endogenous chronobiotic sleep and wake cycle that is disturbed in certain clinical conditions can be restored with a three-layered tablet of melatonin having lactose monohydrate in the top and bottom layer and having CA and PVP mixed with melatonin in the middle layer. The failure of the stem therapy is due to the low retention that could be solved by a core-shell nanofiber comprising of PCL (10%–20% w/v), 10% w/v Gel-MA, and 1% w/v alginate at the core and hydrogel as the cell for injectability. They support the 3D growth of hMSCs, and viability was around 80% after the delivery. Thus other stem cells such as embryonic stem cells and hematopoietic stem cells can be used for the regenerative medicine [55].

adherence. Being nanofibers accepted to intervene in other sectors such as energy, water, and environment remediation, the healthcare sector is not at the bay for exploiting them as efficient drug-release systems. The maturation to translate them as clinical candidates remain unresolved at certain aspects of industrial challenges and advancement in in-vivo and human trials. The industrial challenges can be taken up by the industrial friendly process of fabrication at the laboratory. More scientific experiments need to be focused on biodegradable natural nanofibers with a vast array of a rational combination of drugs to understand the early challenges. As more biological revelations are occurring, robust practices need to be followed in every methodology that will be helpful for a less struggle in the laboratory to clinic transfer (Fig. 1).

3. CLINICAL APPLICABILITY CHALLENGES

4. FUTURE PERSPECTIVE

The annual cost of caring for chronic diseases is more than the US $100 billion. For some of them, long curative or palliative treatment remains without success due to the therapeutic failure or loss of tenacity to patient

As transformations are underway in academics to industry, there have been active pursuits for a noble candidate. The high amount of discoveries in the realm nanofibers is constantly progressive, and expectations

FIG. 1 Nanofiber utility in various domains.

CHAPTER 9 Nanofiber: An Immerging Novel Drug Delivery System are on the upbeat. Hence, drug delivery domain that has emerged in the recent decade does not seem to be halted with unprecedented challenges. The sense of closing the gap is continuous, and commercialization could be realized soon.

5. CONCLUSION Nanofiber has met vibrant results in all aspects of drug delivery, coming out of entrapping fragile to the rugged molecule, with no regard for the drug solubility and penetration in oral, topical, and transmucosal delivery systems with conspicuous merits. Now, it is worth saying that as a prolonged drug delivery system for cancer, wound, organ damage, or microbial resistance, their effectuality is in tandem with the advancement. Hence, considering the huge potentials, scientists from different disciplines must act in coordination to face the challenges.

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CHAPTER 10

Molecular Dynamics Simulations on Drug Delivery Systems ZAHRA SHARIATINIA

Department of Chemistry, Amirkabir University of Technology (Tehran Polytechnic), Tehran, Iran

1. INTRODUCTION Presently, there are diverse kinds of dermatological, oral, and injectable drug formulations that are used for the treatment of various diseases [1–5]. However, such therapeutics cannot control the dug release rate so that the drug is not released at desired site and in appropriate time period [6–9]. Therefore the burst drug release leads to its low bioavailability and fast removal by the renal clearance system resulting in administration of higher doses of drugs [10–12]. Hence, it is necessary to use drug delivery systems (DDSs) to meet several requirements to increase the drug bioavailability and to release the drug at favorite time in anticipated tissue [13–15]. The drug carriers exhibit some advantages such as preserving the drug level for a specific time period, delivery of multiple drugs in one formulation, delivery of drugs to a particular site in the body (targeted drug delivery) and controlling the drug release rate (controlled drug delivery) [16–20]. Such systems enhance the therapeutic efficacy and decrease toxicity and side effects of drugs toward diseased and healthy tissues or organs through transporting the pharmaceuticals to the preferred site. Moreover, the biological, physical, and chemical properties of the drugs are conserved in these systems till they reach their targeted tissues [21–23]. If these systems are prepared in nanosized dimension, they can penetrate into their targeted sites [24, 25]. The drug carriers used to achieve such nanosized systems are diverse materials such as polymers and polymer composites/nanocomposites, graphene, carbon nanotubes, fullerenes, DNA, peptides, proteins, nanoparticles, microparticles, micelles, and liposomes [26–29]. Polymers are widely used to prepare nanomaterials as DDSs so that currently some of them are approved by the food and drug administration, FDA, and under clinical trials [30–32]. Some anticancer drugencapsulated polymeric micelles have entered the market such as paclitaxel-loaded poly(ethylene oxide)co-poly(D,L-lactide) micelles called Genexol-PM [33]. Among different polymers, the most fruitful polymers

include poly(lactic-co-glycolic acid), PLGA, chitosan, and polyethylene glycol (PEG) [34–36]. Indeed, biopolymers are used as important candidates in drug delivery due to their valuable characteristics [37]. Nanocarriers can be substituted by conventional DDSs as numerous studies are accomplished experimentally and computationally to develop efficient nanomaterials as drug delivery vehicles [38–40]. These nanomaterials can control the targeted drug release and lead to lower frequency of drug administration hence they decrease adverse side effects. Modern nanobiotechnology endeavors to resolve the drugs mechanisms of actions at their target sites through optimizing formulations of drug carriers [41]. Drugs indicating low absorption, solubility, and bioavailability have undesirable pharmacokinetics behaviors. In these cases, effective drug carrier must be used in clinical applications. Nanotechnology adopts numerous formulations and approaches to solve the issues related to diverse kinds of drugs particularly the hydrophobic drugs demonstrating poor pharmacokinetics characteristics [42]. Therefore the efficacy of the treatment method is enhanced for specific targets and lower adverse side effects are observed. Smart DDSs are developed through different polymers and biocompatible materials to enhance the targeted and controlled drug delivery abilities of the carriers [43, 44]. Computational molecular dynamics (MD) simulations are beneficial tools to explore the mechanisms of drug delivery at the molecular levels [45]. It is known that achieving effective and optimized DDSs is usually along with challenges such as spending lots of cost and time. Hence, computational methods that are able to estimate the optimized conditions for the designed systems are widely used as alternatives to experimental procedures [46]. Numerous molecular models are used to perform the simulations and modeling of drug carriers to study the properties of molecules in the systems [47]. Also, these computations examine molecular

Modeling and Control of Drug Delivery Systems. https://doi.org/10.1016/B978-0-12-821185-4.00013-0 © 2021 Elsevier Inc. All rights reserved.

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structures and their interactions without needing advanced tools and investigate structural variations without requiring carry out chemical processes [48, 49]. Hence, molecular modeling affords valuable information about the nature of diverse systems that are experimentally formed by interactions of high energies. Most of the computational simulation techniques are low-cost, simple, and rapid that have important effects on identifying mechanisms of intermolecular interactions [50]. However, the ab initio quantum chemical methods consider electrons and protons through quantum equations [51–53]. Consequently, even if several electrons are not considered, for example, in semi-empirical approaches, still many particles must be considered which result in very long, timeconsuming, and costly calculations [54–56]. On the other hand, methods based on force fields, called molecular mechanics procedures, take into account the energies of systems only by considering the nuclei positions. Consequently, such methods are economically and simply applied with low computational costs and satisfactory accuracies for systems including many particles such as biological materials [57]. These methods are commonly used to design drug nanocarriers. Numerous investigations have found that computational data are well consistent with the experimental results [58]. As a result, such evidences indicate that these simulations can be utilized to design DDSs and to predict and optimize the systems properties. As well, the simulations are applied as initial screening approaches to estimate various situations and reduce costs of experiments [59]. Herein, the most recent MD simulations accomplished on numerous materials as drug carriers are reviewed. It is shown that diverse DDSs have so far been used such as polymers and their composites/nanocomposites, graphene and its derivatives, carbon nanotubes and their derivatives, fullerenes, DNAs, peptides, proteins, nanoparticles, liposomes, and micelles.

2. POLYMER COMPOSITES/ NANOCOMPOSITES AS DRUG DELIVERY SYSTEMS Nowadays, polymeric composites/nanocomposites as DDSs are extensively applied as drug carriers using both experimental and computational approaches to investigate the interactions occurred among these systems therapeutics [38]. Such studies indicate that development of new effective drug vehicles is necessary and afford significant information about the biological behaviors of drug-encapsulated formulations [23, 24].

Curcumin-encapsulated starch carrier was used to prevent the Streptococcus mutans bacterial activity by avoiding the formation of plaque and biofilm on teeth [60]. Starch nanoparticles were used as the antiinflammatory natural polymer and an effective antioxidant material to decrease the dental cavities. The molecular conformational changes and the interactions among all molecules were studied by MD simulations. The cell size, energy, density, distribution function (RDF), and temperature radial established that after five steps, starch was stable bound to curcumin in bacteria presence. Nanospherical morphology was achieved for the starch because it released a high energy indicating strong interactions were happened within this system. Also, decreasing the density of the system exhibited its effective antibacterial activity. Salicylic acid-grafted chitosan oligosaccharide (COSSA) was used as DDS and loaded by paclitaxel (PTX) because it revealed outstanding biodegradability and biocompatibility as well as low cytotoxicity [61]. The COS-SA molecular aggregation was examined through MD simulations, and it was demonstrated that the most important interactions occurred in the PTX encapsulation were hydrophobic and van der Waals. Furthermore the hydrogen bond and electrostatic interactions helped aggregation of the COS-SA. The RDFs and solvent available surface areas specified that the COS-SA nanoparticles had high water solubility and they considerably enhanced the water solubility of hydrophobic drugs. Diverse drug-loaded systems were considered and the optimum drug loading was 10 w/w% (Fig. 1). It is known that transporting polynucleotide therapeutics to targeted cells occurs through various interactions by glycosaminoglycan chains situated over cell membranes. Mechanisms occurred during variations of complex nanoparticles of polynucleotide/polymer assisted with glycosaminoglycan are not known but they can completely be explained by obtaining detailed information on intermolecular interactions at atomic levels. Thus nanoparticles containing short complexes of polyethylenimine, PEI, and interfering RNA, small interfering RNA (siRNA), were generated to perform MD simulations using glycosaminoglycan heparin prototype (Fig. 2) [62]. The complex components were bound onto the heparin and revealed important properties related to the siRNA NPs attached to the heparin. Three major metastable situations were seen as the siRNA NPs by adding heparin that was transformed to diverse functional products. Dissipative particle dynamics (DPD) simulations were carried out on the loading and release of doxorubicin, DOX, from a pH-sensitive self-assembled amphiphilic tri-block copolymer formed using

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FIG. 1 Last snapshots of the system after 20 ns for drug loadings of (A) 7.3%, (B) 10%, and (C) 12% [61].

poly(ε-caprolactone)-b-poly(diethylaminoethyl methacrylate)-b-poly(ethylene glycol methacrylate) or poly(sulfobetainemethacrylate) (PCL-PDEA-PEGMA/ PSBMA) (Fig. 3) [63]. It was indicated that the two copolymers were self-assembled in water as core-shell corona micelles but the corona assemblies were very dissimilar for these copolymer micelles. PCL-PDEA-PSBMA and PEGMA micelles created homogenous and heterogeneous shell layers, respectively, that were mostly associated with the lower hydrophilic nature of PEGMA compared with that of the PSBMA (Fig. 4). Increasing the copolymer concentration from 10% to 50%, the PCL-PDEA-PSBMA microstructures were remained as spherical micelles while the spherical structure of PCL-PDEA-PEGMA was changed to the cylindrical shape and at last to a lamellar micelle (Figs. 5 and 6). All micelles revealed pH-responsive drug release properties because the PDEA block could be protonated (Fig. 7). The

drug release was occurred through swelling and demicellization. Follicle-stimulating hormone (FSH) is extensively used in advanced ovarian stimulation procedures but it is daily administered as its half-life is short. The chitosan (CS) nanogels modified by cholesterol (CST) are favorable protein carriers with controlled release features that can diminish the proteins irreversible denaturation and aggregation. MD simulations were done by GROMACS software up to 200 ns to study the mechanisms and interactions happened in the FSHencapsulated CST/CS nanogels (Fig. 8) [64]. The hydrophobic interactions of the CS/CS chains were found to be the foremost driving forces to form the CST/CS nanogels in water. Furthermore the hydrogen bonds along with the hydrophobic interactions could lead to the formation of CST/CS nanogels (Fig. 9). The FSH was

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FIG. 2 Initial (left: top view; and middle: side view) and final (right: top view) configurations of the simulated

complexes; (A) C1, (B) C2, (C) C3, and (D) C4. siRNAs are given in cyan, 568 Da PEIs are in gray, and 1874 Da PEIs are in orange. Water and ions are removed for clarity [61].

gradually encapsulated in the CST/CS nanogel through the hydrophobic interactions of the nanogel hydrophobic domains and the FSH hydrophobic patch (Fig. 10). The FSH flexibility decreased when the nanogel was added except for its hydrophobic patch domain. Therefore the FSH-nanogel interactions could be investigated by such molecular level simulations to design ideal CST/ CS nanogels as the protein carriers.

3. GRAPHENE AND ITS DERIVATIVES AS DRUG DELIVERY SYSTEMS When graphene was prepared in 2004, it found numerous applications as it illustrated incredible low toxicity, biocompatibility, biodegradability as well as electronic and structural properties [27, 28]. Graphene is a material of one atom thickness thus it has been used in various applications including drug carriers, chemical and

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FIG. 3 Chemical structures and corresponding coarse-grained models of simulated systems [63].

biomedical sensors, display screens, solar cells, and electric circuits [29, 39]. MD simulations were carried out on the interactions of various functionalized graphene nanocarriers highly loaded with doxorubicin (DOX), anticancer drug [65]. The properties of graphene sheets functionalized by – COOH (carboxyl), –OH (hydroxyl), –NH2 (amine), and –CH3 (methyl) groups. The MD simulation data demonstrated that the most effective adsorption of DOX was happened on the –COOH-functionalized graphene among all functionalized graphene carriers and this was related to the highest binding energy measured for this system. Moreover, the influences of surface porosity, temperature and hydrogen bonding were assessed on the drug delivery efficacy. It was shown that the solubility parameters and binding energies were dependent on the temperature so that the best results were achieved at 35°C as it was close to the temperature in the human body. MD simulations were performed on the temperature influence on the adsorption of carmustine anticancer drug onto the graphene nanosheet [66]. Root mean

square deviation analyzed at different temperatures revealed that the graphene-drug conjugate was very stable even at high temperature. At higher temperature, as anticipated, greater drug diffusion was observed. Density functional theory (DFT) calculations were also performed to determine binding energies of the active species produced by the drug to alkylate the DNA. The DFT energies exhibited that the N7 site in the guanine base in the DNA major groove favorably bound the drug. Thus the release mechanism and the chemotherapeutic carmustine drug influence on this process could be predicted using these calculations. MD simulations were accomplished on the interactions of graphene nanosheet and β-cyclodextrins in absence and presence of water-indicating β-cyclodextrin was placed over the graphene surface so that strong physisorptions were not occurred on single sorption sites [67]. Instead, they simply placed on the graphene surface and formed a two-dimensional (2D) β-cyclodextrin layer. The β-cyclodextrin movement in the z-direction and perpendicular to the graphene sheet was highly limited. The

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FIG. 4 Configurations of the blank micelles at different block-lengths and different pH values in aqueous

solution. Water beads are not shown for clarity (the same below) [63].

FIG. 5 Configurations of PCL20-PDEA20-PSBMA20/PEGMA20at different concentrations [63].

β-cyclodextrin/graphene complex was hydrophilic which was appropriate for its biomedical applications. The viral protein R (Vpr), Vpr13-33 fragment highly affects the regulation of nuclear import in human immunodeficiency virus (HIV) genes by formation of channels as it forms an alpha-helical leucine-zipper like conformation [68]. However, the helical Vpr13-33 transforms into random coil or β-sheet structures that are aggregated over

the graphene oxide or graphene surfaces by hydrophobic contacts (Fig. 11). To study conformational transitions in viral protein R (Vpr), that is, Vpr13-33, MD simulations were done by confirming Vpr13-33 in solution-preserved the α-helical structure, however, it was converted to the β-sheet structure once it was adsorbed on the graphene surface preferentially through hydrophobic contacts. It was indicated that the structural conversion from α-helical

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FIG. 6 Comparison of self-assembly morphologies of PCL20-PDEA20-PSBMA20/PEGMA20(I) Overall views; (II) Sectional views, and (III) Density profiles of different beads [63].

FIG. 7 pH-responsive drug release behavior of PCL20-PDEA20-PSBMA20: (A) Dynamics process of DOX

release behavior and (B) Radial distribution functions of different beads [63].

to β-sheet was occurred at first nevertheless β-sheet was not completely formed. This was supported with the free energy data on the peptide conformational transformations. The computational results were in agreement with the experimental ones confirming lower Vpr13-33 cytotoxicity after its conjugation with graphene.

To clearly understand the reason for the formation of aggregates by several types of lipids over the graphene surface, all-atom MD simulations were done in both vacuum and water environments to assess the dynamical and equilibrium characteristics of the lipids adsorbed onto the graphene sheets [69]. Different lipid

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FIG. 8 The morphology and structure of the CST/CS nanogel. The snapshot shows that there is more than one

hydrophobic nanodomain in the nanogel structure. In addition, the complexation process was primarily driven through the hydrophobic interactions between the hydrophobic patch of FSH and the cholesterol groups of the hydrophobic nanodomains of the nanogel [64].

aggregates were formed onto the graphene sheets that were dependent on the water presence, quantity of lipid layers, and their initial orientations. Lipid layers were reoriented and self-organized to decrease the hydrophobic mismatches at the interfaces of water/lipid, lipid/ lipid, and graphene/lipid. Several structures were formed such as inverted micelle-resembling assemblies, uniform layers, or weakly bound cylinder micelles over the monolayers that were approved by the experimental data. It was found that graphene could strongly order the lipid molecules directly contacted to its surface and located at 0.35- and 0.85-nm distances. Graphene quantum dots (GQDs) exhibit exceptional mechanical and structural characteristics that make them valuable materials for application as drug carriers, bioimaging agents, and biosensors. Recently, MD simulations were carried out on protein HP35 villin headpiece that was adsorbed onto GQDs of diverse sizes [70]. It was displayed that the π-π stacking interactions among the GQDs and HP35 aromatic residues significantly affected the protein binding onto the GQDs. Also,

increasing the GQD size led to enhancement of the binding strength and number of adsorbed residues that increased the structural change by the adsorbed protein and this was confirmed through several protein structural analyses. Thus these simulations help to understand the GQDs biosafety and toxicity in designing biomedical devices based on GQDs.

4. CARBON NANOTUBES AND THEIR DERIVATIVES AS DRUG DELIVERY SYSTEMS Carbon nanotubes (CNTs) and their derivatives such as boron nitride nanotubes (BNNTs) are interesting nanomaterials for application as drug carriers because they have high loading capacities and able to control the drug release rate [71]. Moreover, they indicate important and valuable physical and chemical characteristics. They are highly chemically stable and reveal compatibility with biomolecules such as proteins and nucleic acids and do not exhibit oxidative DNA destruction. The hollow tubular spaces existing in their structures lead to

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FIG. 9 The solvent exposed hydrophobic patch in FSH [64].

FIG. 10 Snapshots of the FSH-CST/CS system during the nanogel formation indicating the CST/CS nanogels interact with FSH mainly through the FSH hydrophobic patch [64].

simultaneously encapsulation of various therapeutics [72]. In addition, their surfaces can chemically be functionalized by different drugs. Both ends in the structures of CNTs can be removed (uncapping) to facilitate the existing cargos release to the desired cells. Up to now, they have commonly been applied to encapsulate numerous materials and biomolecules including anticancer drugs, metallic nanoparticles, RNA and DNA oligonucleotides, enzymes, and peptides to decrease adverse side effects of drugs and enhance their effectiveness [73]. Estrogen receptor alpha (ERα) is one of the most significant receptors in reproductive system of humans and the ER binding to CNTs could lead to the CNTs toxicity to the reproductive system. MD simulations were performed to understand key problems in SWCNT binding to the ERα ligand-binding domain (LBD) [74]. MD simulations and fluorescence spectra together confirmed the SWCNT binding onto the ERα because a conformational alteration happened to the tertiary structure for which the driving force was the hydrophobic interactions. An in vivo test indicated that the SWCNT exposure improved the ERα expression from 26.43 to 259.01 pg/mL which demonstrated possible estrogen interference effect by the SWCNT. Hence, the modeling results were useful to precisely assess the probable SWCNT health risk. MD simulations were done on gemcitabine, GEM, encapsulated BNNTs incorporated with Au clusters to

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FIG. 11 Snapshots at several time points when Vpr approaching the GA surface in which proteins are revealed in cartoons with purple helix, yellow sheet, and green loop. The GA is shown in cyan. The aromatic residues that form the π-π stacking interactions are exhibited in red, basic amino acid (Lys, Arg, and His) in blue, and Ala amino acid in orange [68].

investigate the drug release from the carriers [75]. It was shown that the size of BNNT highly affected the interactions among BNNT and Au clusters and the Au clusters stability within the BNNT. The Au clusters embedded in BNNT (17,0) exhibited the greatest interactions with the internal wall of the BNNT. The single Au cluster compared with multiple clusters revealed a different tendency to release the drug from the nanotube. Also the van der Waals interactions were the most important forces occurred in the GEM release process (Fig. 12). Configurations and arrangements of diverse DOXloaded SWCNTs as drug delivery vehicles were investigated by the MD simulations [76]. It was displayed that the arrangement and orientation by the DOX molecules were affected by the drug concentration and the SWCNT diameter. In a SWCNT of relatively small diameter with a great confinement like SWCNT (10,10), the DOX molecules preferably formed a single helix in the SWCNT suggesting the drug molecules could be loaded and released in a controlled manner through using an

accurate SWCNTs with a suitable diameter. Consequently, the disadvantages of DOX accumulation in solution could be overcome to decrease the chemotherapy dosage frequency. MD simulations were accomplished on the zwitterionic and neutral ciprofloxacin (CFX) adsorbed onto a SWCNT in both water and gas phases [77]. The CFX was remained adsorbed onto the SWCNT in the two phases throughout the simulation runs. Highly negative interaction energies indicated preferred adsorption onto the SWCNT interior wall. Most of the CFX molecules were adsorbed parallel to the SWCNT surface by the π
π stacking interactions but they were aggregated by the formation of intermolecular hydrogen bonds in vacuum. Another sandwich-like structure was noticed in vacuum once four CFX or more existed. The CFX/CFX hydrogen bonds showed that the CFX molecules were further distributed over the SWCNT in water. However, if eight CFX molecules were added into the SWCNT, somewhat aggregation occurred. Free energies revealed

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FIG. 12 The drug release from nanotube (17, 0) using Au clusters after (A) 0, (B) 1 and (C) 5 ns [75].

that the adsorption process spontaneously happened and both CFX forms were more favorably adsorbed onto the SWCNT internal walls. Thus zCFX was absorbed specially onto the SWCNT interior wall in neutral pH. For the SWCNT/CFX complex, much more negative interaction energies were calculated compared with their free energies. Therefore desolvation was strongly happened during the adsorption that was proved through decreasing the amount of water molecules hydrated the CFX adsorbed onto the SWCNT compared with the free CFX in pure water. The computational data on the zCFX adsorption were confirmed by the experimental results that suggested adsorption was a thermodynamically favorable interaction accompanied by positive entropy. MD simulations were accomplished on the SWCNTs and multi-walled CNTs (MWCNTs) to load and release the DOX drug in different pH media [58]. Also, SWCNT and double-walled CNTs were compared to find which

carrier was superior for the DOX delivery. It was exhibited that the DOX was weaker adsorbed onto the MWCNT and SWCNT in acidic pH compared with the neutral pH and this was because of the electrostatic interactions occurred among the DOX and the CNT carboxyl groups. This result along with the diffusion coefficients, hydrogen bonds, and others proved that the drug was suitably released in acidic pH simulating body environment. As the pH values of cancer tissues are acidic, CNTs were appropriate for delivery and release of DOX drug to the cancerous tissues. Furthermore, the blood pH 7 was favorable to load the DOX molecules onto the CNTs as the DOX/CNT bonds were durable in such pH value. Consequently the CNTs were well capable of DOX delivery within the blood to release the drug into the cancer tissues confirming CNTs were auspicious carriers for the cancer treatment (Fig. 13). As the MWCNT-DOX bonds were stronger, the carrier

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FIG. 13 The DOX loading and release using the MWCNT [58].

displayed a slower DOX release into the sarcoma tissues than the SWCNT and this was the advantageous of the MWCNT for the transport and release of DOX.

5. FULLERENES AS DRUG DELIVERY SYSTEMS Fullerenes are promising nanomaterials that are used in numerous biomedical areas such as drug delivery, gene delivery, bioimaging, and sensors [78]. Recently, fullerene derivatives and particularly the fullerene C60 have attracted substantial attention as DDSs [26]. They can also be utilized in skin care formulations because they reveal high antioxidant abilities that efficiently inhibit radical oxygen species [79]. Fullerene conjugates with various biomolecule including oligonucleotides, peptides, amino acids, esters, and sugars can be used in cosmetics and drug carriers [78, 79]. The mechanism of fullerene C60 permeation to the skin lipid layer was investigated by constrained and

unconstrained coarse grained (CG) MD simulations [80]. The skin layer composed of ceramides, free fatty acid, and cholesterol in equal molar ratios. At low concentrations in water, small fullerenes clusters, 3 and 5, were formed which were spontaneously permeated into the bilayer and dispersed in its interior. In contrast, at higher concentrations in water, fullerenes were aggregated and penetrated in their aggregated form into the interior of the bilayer. Lower fullerenes concentration did not lead to major alterations in the bilayer structure but higher concentration could change the bilayer. The fullerene permeability was dependent on its concentration that was thermodynamically related to the permeation free energy and dynamically to the diffusion ability. Considering the dispersion and aggregation of fullerene, the optimum fullerene concentration was attained that could be used in cosmetics and drug transport formulations. Systematic coarse-grained MD modeling was carried out on the fullerene interactions with cell membranes

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[81]. Bilayers containing lipid species of diverse unsaturation degrees and various cholesterol fractions were selected. Fullerene nanoparticles were added into the phase-isolated ternary membranes by lipid raft model to organize the cell membranes. Addition of fullerene to the lipid membranes improved structural features of membranes such as area, thickness, and inner organization of lipids along with dynamics characteristics such as molecular diffusivity and flip-flop of cholesterol. The phase segregated ternary lipid membranes accumulated the fullerene species especially in disordered liquid domains that promoted domain arrangement and phase segregation within the membranes. Lipid membranes and particularly membranes of little internal organization could dissolve the fullerene molecules. Thus preferential dissolution of fullerenes within the higher disordered hydrophobic domains of lipid bilayers and phase changes might affect the cell membranes activities in organisms.

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MD and molecular mechanics Poisson-Boltzmann surface area (MM-PBSA) simulations were accomplished to study various interactions occurred among different fullerenes and DNA and structural changes that could exhibit their toxicity effects [82]. Hydroxylated and pristine fullerenes did not affect the DNA structure. The hydroxylated fullerenes preferentially bound the major groove of DNA through hydrogen bonding formation and intermediation of water molecules. Fluorinated derivatives penetrated into the DNA structure and formed intercalation complexes of great binding affinities (Fig. 14). The fullerene C60 is extensively used in different biomedical and biomedicine areas such as targeted and diagnostic drug transport systems, tumor growth inhibition, reactive oxygen species scavenging, and bacteria and viruses inactivation. Nevertheless, this nanocage is hydrophobic hindering its practical application. As a result, it is required to functionalize fullerenes to prepare amphiphilic materials. Recently, a carboxylated

FIG. 14 The iso-chemical shielding surfaces (ICSS) for 12mer complexes of (C) C60F10[S] and (D) C60F18[S]. Anti-aromatic and aromatic regions are colored red and yellow, respectively. Fullerene rings facing 12mer nucleobases are also displayed and their corresponding nucleus-independent chemical shift, NICS, values are depicted in the 2D Schlegel diagrams. Fluorine-bonded carbon atoms are colored gray [82].

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water soluble fullerene C60[C(COOH)2]3 was synthesized and its biomedical properties was extensively studied including photodynamic features, radical scavenging capacity through its reaction by diphenylpicrylhydrazyl radical, binding to human serum albumin, hemolysis of erythrocytes, cytotoxicity to HEK293 human embryonic kidney cells, genotoxicity to human peripheral mononuclear cells, and platelet aggregation [83]. Furthermore the MD simulations were done on the structural and dynamical properties of H2O/C60[C (COOH)2]3 to measure the size distribution for the C60[C(COOH)2]3 aggregates. Also the simulation results indicated that the carboxylated C60 more strongly attracted H2O molecules than the pure C60. Several computationally and experimentally and bis and mono adducts of C60 fullerene derivatives were used to evaluate their interactions as inhibitors with HIV type I aspartic protease (HIV-1 PR) by docking methods [84]. Also, MD simulations on free and the inhibitor attached HIV-1 PR were done to complement

the docking results that afforded suitable input HIV-1 PR structure for the docking runs. β-hairpin flaps exhibited diverse orientations in the two models. The HIV-1 PR bound to the inhibitor revealed that the enzyme flaps were pulled into the active site bottom, closed system, but the free HIV-1 PR flaps were moved far from the Asp25 dual catalytic site to form a semiopen system. The MD simulations on the HIV-1 PR flexible and catalytic flap sites helped understand structural changes in these areas and revealed alignments of fullerene compounds in the enzyme active site. Five various combined stereo-electronic groups of three-dimensional (3D) models were achieved using quantitative structureactivity relationships (QSAR) and comparative molecular similarity indices analysis (CoMSIA) for the fullerene analogues to introduce new materials with enhanced HIV-1 PR inhibitory influences (Fig. 15). The optimum QSAR/CoMSIA 3D model generated noncross and cross-validated r2 values equal to 0.993 and 0.739, respectively. This model demonstrated that the of

FIG. 15 (i) CoMSIA steric/electrostatic contour maps of template compound 23 (template compound; has best binding affinity in training set, left on the figure) and compound 36 (has worst binding affinity in training set, right on the figure). Sterically favored areas are shown in green color (contribution level of 80%). Sterically unfavored areas are shown in yellow color (contribution level of 20%). Positive potential favored areas are shown in blue color (contribution level of 80%). Positive potential unfavored areas are shown in red color (contribution level of 20%). (ii) CoMSIA H-bond donor/H-bond acceptor contour maps of compounds 23 and 36 (on the left and right of the figure, correspondingly). The individual contributions from the H-bond donor and H-bond acceptor favored and disfavored levels are fixed at 80% and 20%, respectively. The contours for H-bond donor favored fields are shown in cyan color while its disfavored fields are indicated in purple color. H-bond acceptor favored fields have orange color while its disfavored fields have white color [84].

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FIG. 16 The binding interactions of 1 and 23 at the active site of the HIV-1 PR [84].

significant contributed factors were hydrogen bond acceptor by 28.0%, hydrogen bond donor by 16.7%, electrostatic by 12.7%, and steric by 42.6%. The resultant contour plots were utilized to propose new inhibitor compounds of higher bioactivities and binding affinities toward the HIV-1 PR (Fig. 16).

6. DNAS AS DRUG DELIVERY SYSTEMS DNA nanomaterials and their self-assembled forms have been emerged as effective drug carrier biomaterials with particularly interesting versatile structures and functionalities. Self-assembled DNA nanotubes (DNTs) illustrate exactly controlled biocompatible nanostructures that among the most auspicious drug carriers [85, 86]. DNTs are widely investigated in delivery of various pharmaceutics. It was revealed that folic acid labeled DNTs targeted and internalized the receptors existing on the surfaces of cancer cells [87]. Also, DNTs loaded cargos along their lengths and triggered the cargo release to respond external stimulants [88]. DNTs were used as auspicious nanocarriers that targeted macrophages of tissues [89]. The DNTs cellular uptake was examined using siRNA nanomaterials. Such results could help to recognize the DNTs capacities and benefits in delivery of medicines. However, few MD simulation studies have so far been accomplished on DNAs as DDSs [90–92]. The DNA origami technique causes folding long and single strand DNA as 3D complex arrangements in subnanometer sizes. MD simulations were done to examine microscopic mechanical and structural features of DNA origami substances [90]. MD simulations exhibited that DNTs underwent substantial nanometric structural variations. In aqueous medium the DNA origami structures leaved their perfect targets due to electrostatic, steric,

and solvent-mediated interactions. Holliday junctions in DNA origami materials adopted an antiparallel lefthanded conformation. Type of lattice in these substances significantly affected global mechanical characteristics like bending rigidity. MD simulations were accomplished on selfassembled 384 base pair small origami that were created from staple single and strands of oxDNA that was a nucleotide DNA model in solution [91]. It was observed that new staple strands were attached in parallel; however, the second staple domain was bound when the neighboring bond was partially generated. The system only containing one copy of every staple strand, full assembly happened at intermediate temperatures so that complete assembly was not occurred at low temperatures upon misbonding, whereas at higher temperature extremely large free energy barriers were measured for the assembly. At high concentrations using extra staple strand, full assembly was not seen as two copies of identical staples were attached onto the scaffold and created a kinetic trap which prevented each staple from full binding. In real organizations, such staple blockage could form partially aggregated origamis that could lead to design origami structures. To achieve targeted and smart drug nanocarriers using DNTs, their interactions with anticancer drugs were explored by MD simulations [93]. It was revealed that the DNT drug carriers highly absorbed anticancer drugs through π-π interactions particularly using high drug concentration. Consequently the drug aggregation was significantly decreased in water. Furthermore the DNTs stability was enhanced on drug absorption. This study proved that DNTs were favorable drug vehicles as they strongly absorbed anticancer drugs.

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RNA nanoclusters were constructed using some nanorings with sizes of up to 20 nm and their structural characteristics were studied by MD simulations under physiological human body conditions used to deliver drugs [94]. The variations of ion concentrations about RNA nanoclusters were examined by changing temperature and time. It was revealed that raising the temperature increased the number of ions at a distance from the nanotube but the number of ions decreased at this distance nearby the nanoclusters by quenching the simulations. Thus it was found that some ions evaporated on decreasing the temperature and this happened for the nanorings. The RDF data demonstrated a result comparable with the temperature that was obtained for the RNA nanoring. MD simulations were accomplished using empirical force fields in solution to get insights about the dynamical and structural characteristics of RNA and DNA [95]. One RNA and three DNA sequences were simulated by CHARMM27 all-atom force field designed for the nucleic acids to examine their dynamical, structural, and hydration features such as helical distributions, sugar puckering, and dihedral angles. As well, canonical B and A forms of a DNA hexamer in 75% and 0% ethanol were simulated. The differences in root mean squares of the canonical B and A forms of DNA showed that they had extremely different behaviors confirming such force field could examine the equilibrium state for the two DNA forms. High stability of A form in 75% ethanol but B form in water displayed that the equilibrium was changed through factors such as solvent. Therefore this force field could successfully reproduce various experimental results for the RNA and DNA duplex approving it was very useful for application in computational studies on nucleic acids and their interactions by lipids and proteins. MicroRNAs are noncoding RNAs that can regulate gene expressions within biological systems. In many diseases such as cancer extensive deregulation of miRNAs happened so that some miRNAs become oncogenes and/or tumor-suppressive materials by concurrently targeting various mRNAs. Accordingly, miRNAs are used as auspicious therapeutics in cancer therapies. In this context, peptide nucleic acids (PNAs) were designed that could target 30 UTR over the MYCN mRNA and sis not contain complete complementary base pair sequence, similar to miRNAs. In these experiments, miRNA-34a was chosen as the model that could suppress tumors in numerous cancer cells such as neuroblastoma. Particularly, the miRNA-34a could directly regulate the MYCN oncogene as its overexpression was used as a beneficial biomarker in extremely aggressive phenotype of neuroblastoma. Three oligomers of PNA with diverse lengths

were designed, synthesized and their interactions by two binding domains onto the MYCN mRNA target were examined through MD simulations, ultraviolet-visible and circular dichroism spectra [96]. Uptake in vitro tests and confocal microscopy images for the PNA sequences were acquired using neuroblastoma Kelly cells. Interestingly, although the RNA/PNA hetero duplexes had several mismatches, they highly retained their cellular uptake, affinity, and stability.

7. PEPTIDES AND CELL PENETRATING PEPTIDES AS DRUG DELIVERY SYSTEMS Peptides are short chains formed from amino acids that are bound through peptide bonds [97]. They involve in numerous key biological processes and form precise secondary structures and nanomaterials indicating controlled properties. Nanostructured peptides exhibit outstanding characteristics as they are biodegradable, versatile, bioactive, and biocompatible [98]. Hence, they are used in biomedicine particularly in tissue engineering, drug carriers, and antimicrobial agents. Efficient therapeutic transport in the plasma membrane of cells can be a problem mainly for therapeutics that are large, ionized, or strongly linked onto the proteins in plasma [18]. Cell penetrating peptides (CPPs) were introduced in 1994 as drug carriers for application in intracellular transport of drugs [99]. It was found that the CPPs interactions with cell membranes occurred through electrostatic contacts with proteoglycans. Also the cellular uptakes of CPPs were influenced by factors such as cell type, the secondary and the primary CPPs structures, membrane composition, cargo nature/concentration, concentrations of salts, and CPPs. As CPPcargo conjugates and CPPs can internalize by endocytosis pathway, an effective way for the delivery of cargoCPP is the endosomal escape [100]. Both of the covalent and noncovalent bonds can be formed between CPPs and pharmaceuticals to form DDSs. MD simulations were carried out on the formation of polypeptide membranes from surfactant resembling peptides containing 15 amino acids that created a hydrophobic domain including three valines (V), three isoleucines (I), three glycines (G), three alanines (A), and a hydrophilic area generated through three lysines (K) [101]. Figs. 17 and 18 illustrate the initial and final structures of the K3G3A3V3I3 and I3V3A3G3K3 polypeptides. The density values, hydrogen bonds amounts, van der Waals and Coulombic interactions among peptide-peptide residues and peptide-water specified that although much water was infiltrated, the membranes maintained their structures.

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FIG. 17 The YZ, XZ, and XY planes and water boxes for I3V3A3G3K3 (top) and K3G3A3V3I3 (bottom) polypeptide initial structure. Isoleucine, valine, alanine, glycine, lysine and water are in purple, ochre, gray, yellow, red, and blue, respectively. The water boxes (YZ plane) with two water layers are also revealed. The polypeptide membrane for the simulation is inserted in the empty region, forming a double surface with direct water interaction [101].

FIG. 18 The YZ, XZ, and XY and water boxes for I3V3A3G3K3 (top) and K3G3A3V3I3 (bottom) polypeptide final structures. Isoleucine, valine, alanine, glycine, lysine and water are in purple, ochre, gray, yellow, red, and blue, respectively [101].

To get insight about the microscopic self-assembly mechanism of a drug-binding peptide, MD simulations were accomplished at both body and room temperatures [102]. The peptide was used as the carrier for the chemotherapeutic drug DOX. The DDS was composed of five DOX molecules and one peptide in water. In all systems, multiple DOX molecules were spontaneously attached onto the peptide that was in

agreement with the experimental data on the drug affinity to the peptide. It was revealed that the existence of tyrosine and tryptophan were substantial in the peptide binding and the prevailing DOX-peptide interactions were the π-π stacking contacts. Therefore this peptide was a valuable medicine carrier that showed the ability to bind the π-conjugated therapeutics.

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MD simulations and atomic force spectroscopy (AFS) were applied to rationally design inhibitors which indicated blocking of amyloid-amyloid bonds formation that was the initial phase in the formation of toxic amyloid oligomer [103]. The pseudo-peptide amyloid-β (Aβ) inhibitors (which were bound onto the Aβ peptide) were created that efficiently prevented formation of the amyloid-amyloid bonds. The binding affinities for the Aβ and inhibitors besides the inhibitors to themselves were achieved by Umbrella Sampling computations. AFS experimentally examined some inhibitors to measure their capacities in blocking the formation of amyloid-amyloid bonds. It was found that the AFS results were consistent with the MD simulations as three pseudo-peptides bound to the amyloid fragment by diverse affinities that successfully prevented the Aβ-Aβ binding. Hence, these pseudo-peptides were proposed as promising drugs that illustrated the ability to hinder the toxic effect of the Aβ in the Alzheimer’s syndrome. MD simulations were done on the ibuprofen (IBU) load and release using (AF)6H5K15 amphiphilic peptide, FA32, and its analogues F16H5K15 and F12H5K15 [104]. After the IBU is loaded into the FA32, core-shell spherelike micelles were created. The IBU drug was primarily positioned inside the hydrophobic core that was enclosed with the alanine and phenylalanine residues but lysine was within the hydrophilic shell. Increasing the IBU concentration enlarged the micelles upon increasing the hydrophobic contacts. The IBU loading into the FA32 analogs afforded various morphologies especially using F16H5K15 formed a nanofiber structure. When pH was changed, the IBU release from the F16H5K15 nanofiber was faster compared with that of the from FA32 micelles signifying the latter was a superior controlled release system. Also, it was demonstrated that the IBU-encapsulated morphology was varied upon altering the peptide type that significantly affected the IBU release. Accordingly, such bottom-up method was beneficial to rationally design drug vehicles that could efficiently load and release drugs.

8. PROTEINS AS DRUG DELIVERY SYSTEMS Proteins are also utilized as DDSs particularly human serum albumin (HSA) is frequently applied in transporting various pharmaceutics [105]. HSA is comprised of 585 amino acids that can transport numerous exogenous and endogenous substances such as nutrients, fatty acids, metal ions, steroids, drugs, and hormones [106]. The HSA structure contains three homologous areas of I, II, and III and each domain is divided to two subdomains A and B so that merely one tryptophan residue

(Trp-214) exists within the subdomain IIA [107]. Different materials are usually bound to the two main HSA areas located inside the hydrophobic cavities in subdomain IIA and IIIA called Sudlow’s sites I and II, respectively. Site I is situated within the hydrophobic cavity in the subdomain IIA that can bind diverse heterocyclic and neutral materials through strong hydrophobic contacts but site II existing in the subdomain IIIA is attached onto several aromatic carboxylic acids via hydrogen bonds and van der Waals forces [107]. The C-1027 aromatic chromophore (Chr) that is able to selectively cleave DNA is delivered and stabilized in vitro using apoprotein (apo) but it is released when the holoprotein, apo + Chr, is penetrated to the cultured sarcoma cells [108]. The holoprotein is used as an attractive DDS in clinical trials whereas the mechanism of Chr release is unclear. Hence, MD simulations were done to find the release paths indicating they were dependent on local movements by 3 loops including Asn97–Leu100 (L9), Thr75–Thr79 (L7), and Val39– Gln42 (L3). Major problems in the Chr release were hydrophobic interactions, direct hydrogen bonds, and steric hindrance happened through the three loops. As well, Ser98 was a significant residue throughout the release course. The interactions of HSA and phthalic acid esters (PAEs) as endocrine disruptor were examined using MD simulations and fluorescence spectra to evaluate the HAS-PAEs distances energy transfer between them [109]. It was revealed that all four types of PAEs quenched the inherent HSA fluorescence through nonradiative transfer of energy and static quenching mechanisms. Thermodynamics tests and molecular docking proved that binding was mostly controlled by hydrophobic forces. Also, four PAEs were primarily bound onto the subdomain IIIA of HSA demonstrating good agreements between the computational and experimental data. MD simulations illustrated that HSA conformation was a slightly changed on its binding to the PAEs. Moreover, PAEs-HSA complexes had higher stabilities than the native HSA protein. The interactions among HSA and aflatoxin G1 and B1 were studied by MD simulations molecular docking and fluorescence spectra [110]. The fluorescence spectra verified that the HSA fluorescence emission was substantially quenched by adding aflatoxin G1 and B1 by mechanism of static quenching. The calculated thermodynamic factors specified that the spontaneous nature of the interactions so that the van der Waals and hydrogen bond forces significantly affected the HSA binding to aflatoxin G1 and B1. The binding constants for the aflatoxin G1 and B1 linking onto the HSA were

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measured. The molecular docking data exhibited that aflatoxin G1 and B1 were bound onto IB subdomain in HSA predominantly through hydrogen bonds and hydrophobic interactions and these results were well matched to the fluorescence spectra. The interaction constants attained by the docking studies were matched to those of the experimental values measured from the fluorescence spectra compared with those achieved by the MD simulations. Overall, the results obtained from the MD simulations molecular docking and fluorescence spectra confirmed each other. It is found that the crystallization capacities of proteins are enhanced on substituting lysine residues in their surfaces by other residues. Matrix-assisted laser desorption and ionization time-of-flight mass spectra were acquired experimentally for the PH1033 protein of Pyrococcus horikoshii that was chemically modified by NHS-biotin to assess the surface lysine residues [111]. The biotinylation of protein using 1:1 molar ratio indicated that merely 7 of 22 lysine residues existing within the protein containing 144 residues were biotinylated. MD simulations were done to mimic and analyze the experimental results confirming the biotinylation was considerably affected by four parameters related to the local surroundings of lysine residues including pKa values, solvent accessibility, number of hydrogen bonds, and electrostatic energy. Therefore the biotin functionalization avoids using lysine residues having high intramolecular interactions that can decrease the proteins crystallinities.

9. NANOPARTICLES AS DRUG DELIVERY SYSTEMS Nanoparticles with small diameters in the 10–100 nm range can be served as effective drug carriers that are freely circulated even in capillaries thus they are naturally superior compared with bigger drug transporter materials in crossing some biological barriers [112]. It is worth mentioning that, nevertheless, nanomedicines pass longer and more difficult FDA approval processes relative to parent unimolecular drugs [113]. The reason is that it is required to investigate the effectiveness, side effects, in vivo aggregation, drug release, and other properties of nanomedicines each components of their formulations to be FDA-approved as nanocarriers. FDA regulations, nanocarriers complexity and performance changes by reformulation of small drugs into the nanocarriers/nanomedicines are among important issues for the nanocarriers commercialization by the pharmaceutical companies. Fortunately, as the number of FDAapproved drugs and those currently used in clinics is

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increasing [114, 115] and FDA presents more obvious nanomedicine approval rules, there is a growing interest in clinical nanomedicine development [113]. The PEG amount and temperature effects on the PEGylated NPs translocation through the asymmetric plasma membrane in eukaryotic cells were examined by coarse-grained MD simulations [116]. It was exhibited that NPs translocation in the membranes was enhanced by raising the temperature and this was related to formations of more disordered lipids and quicker diffusion. In contrast, steric hindrance influence by PEG inhibited the NPs translocation and promoted flip-flop of lipids. The PEG chains were rearranged to decrease the interactions of lipid tails and PEG throughout the translocation and this looked like the snorkeling. Besides, flip-flops of lipids were changed with the PEGylation degree and the translocation direction by the NPs. Greater amount of PEGylation promoted the flip-flops of lipids whereas it inhibited the extraction of lipids from bilayers. More symmetric membranes were created on lipids extraction and their flip-flops. Oral chemotherapy method is favored over injection and other methods nevertheless its usage is limited as anticancer drugs reveal little bioavailability [117]. Selfassembled nanoparticles can be applied as promising nanocarriers to solve this issue however designing suitable nanocarriers is challenging. Effective DDSs of HA (hyaluronic acid) copolymers were designed for the DOX drug using MD simulations and the chain length effect in the fatty glyceride of HA was estimated on peroral DOX absorption. MD simulations were done to assess the DOX compatibility with HGS (HA-g-glyceryl monostearate), HGL (HA-g-glyceryl monolaurate), and HGC (HA-g-glyceryl monocaprylate). The HA copolymers were also synthesized to confirm the predicted results. The HGS, among all copolymers, exhibited the most compatibility by DOX and then HGL and HGC. The stability and physicochemical features of all nanoparticles were dependent on the structures of copolymers so that the HGS/DOX nanoparticles displayed superior characteristics and subsequently HGL/DOX and HGC/DOX nanoparticles. This trend was observed for the epithelial transport, cellular uptake, and in vivo absorption tests in rats as HGS/DOX NPs revealed seven times greater absorption when perorally administered compared with the intravenous DOX injection. Hence the MD simulations effectively used to rationally design nanoparticles for the oral transport of DOX. Drug carriers based on lipids are auspicious materials for hydrophobic drugs. Distribution of lipids in droplets changes their loading capacities. Consequently, MD simulations were performed on diverse kinds of lipid

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nanoparticles with diameters of around 12 nm, that is, NLC (nanostructured lipid carriers), NE (nanoemulsion), and SLN (solid lipid NPs) to recognize lipids organizations in the carriers [118]. These lipid nanoparticles contained oleic acid and stearic acid lipids plus sodium dodecyl sulfate surfactant dispersed in water (Fig. 19). Additionally, the influence of solid/liquid ratio on distribution of NLC lipids in the carriers was examined. At equilibrium, it was found that vesicle-like structures were formed in all carriers in which the hydrophilic fragments of surfactant and lipids were arranged as a semi-bilayer that was folded to a droplet with their hydrophobic tails were accumulated amongst them. Though it was expected that the harsh sodium dodecyl sulfate surfactant was located on the droplet surface, it was entered the lipids. Furthermore, high amount of water beads were entrapped within the droplets as one or more cavities alongside the inner layers of head groups that were enclosed with the head groups of lipids. Also, for the SLN and nanoemulsion structures, in the droplets centers, denser lipids were formed compared with that of the NLC. Additionally, crystallization did not take place in the lipid carriers. The lipid distribution in the NLC carrier was not affected by the solid/liquid lipid mass ratio.

10. LIPOSOMES AS DRUG DELIVERY SYSTEMS Liposome drug carriers exhibit valuable advantages like nontoxicity and biodegradability that are generally formed from amphiphatic phospholipid compounds [119]. So far, important advances have been attained on using liposome drug delivery formulations but there are little approved liposome drugs for clinical usage and primarily as antitumor and antifungal agents. Indeed, the clinical applications of liposomal DDSs especially include cancer chemotherapy and acute fungal toxicities [120]. The anticancer drugs encapsulated in liposomes illustrate highly different pharmacokinetics and biodistributions which cause decreased cytotoxicity and enhanced targeted drug release into anticipated tissues. Liposomes have spherical and enclosed shapes that create lipid bilayers or membranes. Phospholipids reveal asymmetric structures containing hydrophobic chains of fatty acids and hydrophilic choline and phosphate head groups. Hence, the structural and physicochemical characteristics of liposomal bilayers are not isotropic. In fact, liposomes form single membranes known as unilamellar liposomes that are large or small unilamellar vesicles or they form multilamellar membranes representing matryoshka doll shaped concentric

structures called multilamellar vesicles [121]. In multilamellar liposomes, the number of concentric membranes indicates the liposome lamellarity. The liposome DDSs are exceptionally miscellaneous. They are able to transport hydrophobic cargos in the lipid nonpolar core of their membranes or hydrophilic cargos within their inner aqueous pocket. Moreover, the head groups of phospholipids can be conjugated with polymers to protect them. Furthermore the lipids functionalized by polymers can be conjugated with targeting ligands for application in targeted drug carriage [122]. Sheared polymers grafted to flat surfaces were created in the MD simulations as liposomes functionalized by PEG brushes that could be used as drug carriers for the topical therapy of human vasculature diseases [123]. In such application, particular rheological features must be met like low viscosities at large shear rates to enhance the liposomal drug carriage. Consequently, MD simulations were done on polymeric PEG brushes with different lengths and shear rates were applied to achieve average viscosities and friction coefficients that were affected by polymerization degrees and shear rates in theta solvents. It was exhibited that high shear rates led to substantial shear thinning of the PEG brushes. The simulation data were in agreement with the measured viscosity values at high shear rates for red blood cells within a solution containing liposomes. Flavonoids and cathechins display various advantageous that help to be healthy. They contain two substances that exhibit, among others, therapeutic and antioxidant properties. These compounds include morelloflavone (MF) and epigallocatechin 3-gallate (EGCG) extracted from Garcinia dulcis and green tea, respectively. MD simulations were carried out to study the interactions of MF and EGCG by the lipid bilayers indicating addition of salts affected the encapsulation degree of the neutral liposomes [124]. As expected, EGCGs were bound onto the hydrophilic phospholipids groups to be mainly placed at the lipid-water interface. The salt concentration and formula influenced absorption of the EGCG. Moreover, aggregated MFs were highly stable in water that greatly penalized the interactions of flavonoid by the polar lipid head groups. The MF penetration into the liposome was affected to the salts presence although they are cannot entirely assist its absorption into the bilayer. The penetrations of both substances were increased on adding magnesium chloride but calcium chloride indicated a reverse influence. Liposomal drug carriers are used to be adhered onto specific tissues and sites. Nevertheless, information is little about precise drug delivery mechanisms and drug

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FIG. 19 Final distributions of lipid moieties inside droplets at different solid to liquid lipid fractions. The snapshots are captured from the cross-section of droplets at last time frames of simulation. Stearic acid head group are rendered in green, oleic acid head group in red, SDS head group in yellow, water beads in blue, and all the hydrophobic tails in gray [117].

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distributions within lipid carriers [125]. Hence, coarsegrained MD simulations were done on a liposome containing more than 2500 lipids indicating various drug loading degrees. Hypericin was used as the drug molecule (commonly applied for the photodynamic treatment) to penetrate the membrane bilayers. Simulations of 10 μs were run on liposomes including 21
84 hypericin molecules. The distributions and orientations of hypericin molecules in the lipid bilayers and average force potentials needed to transfer them from aqueous inner droplets across the liposome bilayers were estimated.

11. MICELLES AS DRUG DELIVERY SYSTEMS Polymeric micelles are amphiphilic block copolymers that are self-assembled structures in aqueous solutions exhibit exceptional characteristics such as large drug loading, biocompatibility, and high in vivo stability which make them beneficial drug carriers [126]. Micelles applied to deliver anticancer drugs illustrated extraordinary properties such as improved targeted drug release, high chemotherapeutic effectiveness, and lower undesirable drug side effects. Hydrophilic and hydrophobic regions in amphiphilic copolymers form diverse micelles for application in delivery of genes, proteins, therapeutics, and drugs [127]. Such copolymers form core-shell nanostructures by inter- and intra-molecular interactions. Polymeric core-shell micelles contain hydrophobic cores covered by hydrophilic shells that are widely used nanobiotechnology and pharmaceutical areas. Micelles can be derived from both natural and synthetic polymers. Usually, natural polymers are more advantageous as they are nontoxic, biodegradable and biocompatible [128]. Polymeric micelles are broadly examined as drug vehicles but it is necessary to understand their detailed morphological changes on drug loading [128]. It has been shown that rods, bilayers, spheres, vesicles and cylinders are created by changing the composition of block copolymer, solvent, interactions of blocks, ionized blocks, pH and temperature [129]. Consequently, investigating mechanisms occurred during formation of micelles as drug carriers are vital because such information assists to recognize their structural and morphological variations at microscopic levels. To explore structural and dynamical features of polymer micelles, several techniques can be utilized including UV-visible and fluorescence spectra, dynamic light scattering, transmission and scanning electron microscopies [129]. Nevertheless, it is hard to get comprehensive data on the self-assembly and transformation

mechanisms of polymer micelles using such analyses as the micelles are formed at nanoscale. Hence, computational simulations are complementary tools to the experimental methods that afford more information to understand the morphology variation, distribution and dynamics of the systems. For example, the mechanism of drug carrier formation within aqueous media is explored with more details by mesoscopic simulations instead of using experimental approaches [130]. However, among typical methods of mesoscopic simulations, the MD simulations provide time scales that are too short to allow micelle formation and simulations at the atomic level is highly costly [131]. Thus DPD allowing very larger length scales and time steps is favorably applied to simulate very complex coarse-grained systems. In fact, the DPD simulations are used as an effective and systematic technique to study the formation mechanisms and microstructures of different polymer micelles [132]. MD simulations were accomplished on the solubilities of hydrophobic drug compounds Cucurbitacin I (CuI) and Cucurbitacin B (CuB) within poly(ethylene oxide)-b-poly(α-benzyl carboxylate ε-caprolactone) (PEO-b-PBCL) block copolymers indicating diverse tacticities [133]. Particularly, a di-block copolymer of various three tacticities, that is, PEO-baPBCL, PEO-b-sPBCL, and PEO-b-iPBCL was utilized. Binary random mixtures containing 10 wt% of drugs were used to calculate the solubility values. The solubilities of the two drugs were highly dependent on the di-block copolymer tacticity. MD simulations exhibited that only PEO-b-sPBCL was soluble but the two others did not reveal solubility. As the drugs experimentally indicated solubility into the PEO-b-PBCL, the experimentally synthesized di-block copolymer was expected to display syndiotactic tacticity. Such prediction was confirmed by the results obtained from the ring opening polymerizations of cyclic lactones dominantly yielded syndiotactic polymers using stannous octoate catalyst to synthesize PEO-b-PBCL block copolymers. MD simulations revealed that the drugs solubilities within the PEO-b-sPBCL was dependent on the intramolecular and intermolecular interactions of drugs and di-block copolymer molecules that were investigated by the RDF data. MD simulations were done on a DDS containing one PEO-b-3PCL block copolymer composed of three blocks of poly(ε-caprolactone) (PCL) with identical lengths attached onto one end of poly(ethylene oxide) (PEO) block that encapsulated two groups of hydrophobic drugs having diverse structures [134]. The first group of drugs was two CuI and CuB cucurbitacin

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drugs containing several hydrogen bonding acceptors and donors uniformly located on these structures but another group of drugs were nimodipine and fenofibrate principally only containing clustered hydrogen bonding acceptors. The PEO-b-3PCL including cucurbitacin drugs exhibited substantially improved drug solubility than the PEO-b-PCL di-block linear copolymer having a PCL:PEO ratio of 1. Nevertheless, a contrast result was achieved for the nimodipine and fenofibrate drugs. The intermolecular interactions confirmed that much more hydrogen bonds were formed among the cucurbitacin drugs and 3 PCL blocks compared with those of the di-block linear copolymer. Also, as the hydrogen bond donors are absent and the hydrogen bond acceptors are clustered onto the nimodipine and fenofibrate drugs, the hydrogen bonds created within the multi-PCL block milieu were considerably decreased and afforded undesirable solubility values. Thus copolymers with multiple hydrophobic blocks highly increased loading of hydrophobic drug molecules on which several hydrogen bond acceptors and donors were uniformly dispersed. MD simulations were carried out in water at 1 atm and 298.15 K on a spherical micelle formed using N-acetylated poly(ethylene glycol)-poly(γ-benzyl L-glutamate) (PEG-PBLG-Ac) amphiphilic block copolymers, containing 9 BLG and 11 EG units, that were used as drug carriers [135]. Calculations were done on the spherically arranged copolymers and reached the equilibrium indicating somewhat elliptical micelle composed of a PBLG hydrophobic interior core with a PEG hydrophilic external shell. PEG blocks in the micelle showed dense helical conformations and the PBLG blocks revealed R-helical forms. Numerous hydrogen bonds formed by the water solvent molecules led to stabilization of the helix conformations for the PEG blocks so that they became hydrated and this was proved with extended residence times for the water molecules around the oxygen atoms of PEG ether than that of the pure water. Several water molecules were dispersed in the hydrophobic core that displayed constant exchanges by the pure water throughout the simulations. The molecules generally formed clusters in places among the copolymers that created hydrogen bonds with themselves and the hydrophobic core by hydrophilic amides and esters groups. The micelle formed hydrogen bonds with water molecules that greatly stabilized its structure. Hybrid polymer micelles were designed as drug carriers containing combined corona chains of polyethylene glycol (PEG) and poly(L-glutamic acid) (PLGA) (Fig. 20) [136]. The water-soluble PLGA random coils in acidic solution were changed to water insoluble

175

α-helices that caused micro-phase separated micelle coronas and creating PEG channels. The channels connected the interior core and the external shell that enhanced the drug diffusivity from the micelles into the solution. The PEG-b-PPO and PLGA-b-PPO-b-PLGA formed micelles in water, where PPO was poly(propylene oxide). The PPO blocks in the two block copolymers were aggregated throughout their selfassembly to cores enclosed with corona mixed chains of PEG and PLGA blocks. Such structures were characterized by zeta potential, nuclear magnetic resonance spectra, dynamic light scattering, and circular dichroism spectra. As well, MD simulations were accomplished to investigate the structures of hybrid micelles as DDSs indicating preserved colloidal stability along with adjustable release rates. The release rates were controlled with micelle structures, mixture compositions and other factors like pH.

12. CONCLUSION Computational modeling methods provide valuable tools to design and develop diverse drug carriers with improved features including nanoparticles, polymers and polymeric nanocomposites, graphene and its analogues, fullerenes, carbon nanotubes and its derivatives, DNAs, proteins, peptides, micelles, and liposomes. The efficacy of these systems is mostly dependent on their drug loading capacities, drug release rates, and blood stability. MD simulations afford very important full data on carriers’ structures and interactions occurred among them and with drug molecules under different physical and chemical environments. MD simulations are complimentary to the experimental results because they provide quantitative and microscopic information about the mechanisms of experimentally happened phenomena. Also, then are able to recognize limiting elements in different systems that aid to find the optimized formulations. So far, DPD and CG approaches have been used in MD simulations to investigate the properties of various drug carriers. Atomistic MD procedures have restricted length and time scales of some nanometers and microseconds but the DPD and CG approaches are more appropriate to simulate submicron scales for several hundreds of microseconds that neglect atomistic information for both of carrier and drug. Hybrid CG and atomistic approaches gives a favorable method for the scales problem as they take benefits of the two procedures. Although the carriers’ stability and loading are central parameters for their development, another fundamental factor is carriers’ interactions with the biological membranes. These investigations offer in-depth

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FIG. 20 Snapshots of A10B7A10 triblock and B1C3 diblock copolymers in (A) basic and neutral as well as (B) acidic conditions. The red, blue, and green lines represent C, A, and B blocks, respectively. Arrows indicate the C block aggregation areas [136].

information on mechanisms of drug transport through cell membranes. MD simulations are also able to satisfactorily predict the tumor microenvironments in tissue scales. The heterogeneous nature of solid tumors varies from one patient to another because of the genetic substances contributing in the tumor progress. The biomedical imaging techniques afford biological and physical properties of tumors. Nevertheless, most of currently used imaging methods of high-resolution are not appropriate for in vivo experiments as they cannot give full images indicating microscopic interactions occurred within the tumor microenvironments. Hence, more efforts are needed to develop effective multiscale approaches to couple macroscopic and microscopic transport routes to envisage the effects of tumor microenvironments on drug distributions inside real tumors.

MD simulations can be used in patient-specific applications. For this purpose, it is required that patients’ data are gathered at early diagnosis stage and after several follow-ups which include elementary biochemical information, clinical data, and medical high-resolution images. Adding such information to computational models allows performing simulations for each person that are useful in designing efficient treatments for the patients. It is expected that future clinical treatment methods and postoperation cares take advantage of personalized simulation. To decrease the costs of computational simulations, complex models must not be used and instead reduced models are more attractive containing specific patient information as inputs to initially predict the effectiveness of treatment. Nonetheless, such systems are very non-linear and they can show several resistances into the treatments by drugs. Consequently,

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it is required to gather enormous experimental data to improve the models of adequately high accuracies. Finally, considering all attempts have been made in this intriguing field of study, it is anticipated that the MD simulations find their way in the near future to the clinics by predicting the effectiveness of various DDSs. This is reasonable as the MD computational methods consider very much larger systems than ab initio and DFT methods. Moreover, a combination of ab initio, DFT and MD simulations can be used to predict various physicochemical properties of the designed drug carriers. This allows selection of the most suitable DDS before its trial on the patients. Hence the MD simulations can reveal promising clinical applications in future.

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ACKNOWLEDGMENTS This work was financially supported by the Research Office of Amirkabir University of Technology (Tehran Polytechnic), Tehran, Iran. Author is gratefully appreciated.

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CHAPTER 10

Molecular Dynamics Simulations on Drug Delivery Systems

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CHAPTER 11

Nanoparticle Drug Delivery: An Advanced Approach for Highly Competent and Multifunctional Therapeutic Treatment SAIMA AMJAD • M. SERAJUDDIN

Department of Zoology, University of Lucknow, Lucknow, India

1. INTRODUCTION Nanoscience and nanotechnology is the most emerging multidisciplinary scientific field at the beginning of the 21st century, which has given better opportunities using engineering and manufacturing principles at the molecular scale to design and develop complete, high performance products [1, 2]. Nanotechnology is revolutionizing the medical field by using nanomaterials because these nanomaterials have an ability to cure the disease by targeted drug delivery at the cellular and molecular level [3]. The prefix “nano” is derived from Greek (Latin nanus) which means dwarf. In the International System of Units, a nanometer is equal to onebillionth of a meter (10
9) [4]. National Nanotechnology Initiative (NNI) defines nanotechnology as the manipulation of matter with all three dimensions at least one dimension range approximately 1–100 nm [5]. Richard Adolf Zsigmondy gave the first accurate observation and size measurement of the nanoparticle by using dark field ultramicroscopy and first to coin the term “nanoparticles” [6]. In the year 1959, physicist Richard Feynman expressed the concept of nanotechnology in his lecture “There’s plenty of room at the bottom” and suggested the possibility of direct manipulation to make nanoscale machines that is assembling matter on an atomic scale [7]. Feynman further explained in terms of medicine that the use of tiny machines would become interesting in surgery if the patient swallows the surgeon, that is, the tiny machines or surgeons move inside the blood vessel and find out which valve is the faulty one and takes the knife and removes it out. Moreover, the other tiny machines might be permanently residing in the body to aid some improperly functioning organ [8].

The term “nanotechnology” was coined by Norio Taniguchi at the University of Tokyo (1974). Kroto’s and Smalley’s research team discovered fullerene C60 in the year 1985 and Saumio Lijima discovered carbon nanotubes in 1991 [9]. In 2000 the United States launched the NNI to make the way for the future advancement of nanotechnology. The first nanoparticles of 100-nm diameter of poly(methylmethacrylate) as a new adjuvant were made by Kreuter and Speiserin (1976) [10] in the drug delivery area. Nanomedicine is the branch of nanotechnology in which the procedure of treatment, diagnosis [11], and prevention of diseases [12] using the novel methods of drug delivery [13] such as biocompatible nanoparticles [14] and nanorobots [15] generate more efficient and effective therapy [16]. According to Emerich and Thanos [17], nanomedicine application for drug development depends on various molecular technologies which broadly included three classes which are represented in Fig. 1. Conventional drugs which are used by the traditional medical practitioner are not very effective because of their poor solubility and have limited bioavailability after oral and intravenous intake. Although these limitations of conventional drugs could be diminished by the application of nanotechnology approaches by the drug delivery method. The targeted nanoscale drug delivery system which has the potential to revolutionize drug delivery systems used nanomaterials such as nanocapsule, nanoparticles, nanopores, nanoliposomes, nanoshells, dendrimers, fullerenes, nanotubes, quantum dots, nanosphere, nanovaccines, and nanocrystals. Thus nanodrug formulation can be used for the intentional development of new drug delivery systems and reinvent existing drugs to enhance efficiency, patent protection,

Modeling and Control of Drug Delivery Systems. https://doi.org/10.1016/B978-0-12-821185-4.00008-7 © 2021 Elsevier Inc. All rights reserved.

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FIG. 1 Classification of nanomedicine application.

patient-compliance, risks of toxicity, and decreasing the cost of health care [18,19]. The encapsulation of medicinal drugs increases the specificity, efficacy, tolerability, and therapeutic index of subsequent drugs [20–25]. Nanomedicines have several advantages for the protection of premature degradation of drug and interaction with the biological component, enhancement of absorption into a selected tissue, bioavailability, retention time, and improvement of control delivery [26–29]. Some nanomedicines are used in different trial phase of testing for the diseases such as diabetes [30], cancer [31], AIDS [32], malaria [33], prion disease [34], and tuberculosis [35]. Furthermore, nanochips, magnetic nanoparticles for chemotherapy attached to the specific antibody, and nanorobotics are new dimensions of their use in drug delivery. Nanoparticle delivery systems are engineered technologies that are designed and tested for the use of nanoparticles for developing clinically useful therapeutic agents for the targeted delivery and controlled release [36]. Nanoparticles include a variety of materials such as lipids (liposomes), viruses (viral nanoparticles), polymers (polymeric nanoparticles, dendrimers, or micelles), and metallic nanoparticles. The advancement of nanodevices which are synthesized by the incorporation of biocompatible/biodegradable polymers has therefore rapidly emerged with the discoveries of albumin [37], polyalkylcyanoacrylate [38], polylactate-coglycolate [39], and afterward, solid lipid [40] or chitosan [41] nanoparticles. The other advantageous property of nanoparticles is to improve the solubility of orally taken poorly soluble drugs [42]. There are several anticancerous drugs that are used in chemotherapy treatment are poorly water soluble such as paclitaxel, doxorubicin. The nanoparticle formulations of these drug increases

the bioavailability without using undesirable excipients, such as polysorbate or Cremophor EL which are used in Taxoteres and Taxols formulations, respectively [36].

2. BACKGROUND OF NANOPARTICLES IN HUMAN HISTORY AND DRUG DEVELOPMENT Humans are always surrounded by nanoparticles and their existence in the environment for a long time and are not necessarily produced by modern techniques. The Ancient Egyptians were using lead sulfide (PbS) nanoparticles (NPs) ( 5-nm diameter) for hair dye around more than 4000 years ago [43]. Similarly, in the 3rd century BC, “Egyptian blue” synthetic pigment was first prepared by Egyptians by using a mixture of nanometer-sized glass and quartz [44]. Metallic nanoparticles were also used in ancient times as color pigments in luster and glass technology [45,46]. During the 9th century, metallic luster decorations of glazed earthenware were found in Mesopotamia [47]. At late 1960s Peter Paul Speiser synthesized the first nanoparticles which can be used for targeted drug therapy [48] and the first research paper on nanoparticles “a pioneer in the conception of nanoparticles” was published by Speiser (1976) which focused on the advancement of nanoparticles for vaccination purposes, targeting for a slow release of the antigen, leading to a better immunity [19,49]. Georges Jean Franz K€ ohler and Cesar Milstein in the 1970s succeeded in constructing monoclonal antibodies [50]. The first nanoparticles were modified as a carrier for the transport of fragments and genes and were reached into cells with the help of antibodies at the beginning of the 1990s [48,51]. Paul Ehrlich in the starting of the 20th century attempted

CHAPTER 11 to create “magic bullets” on which drugs were loaded and would kill all pathogens after only a particular treatment [48]. In the late 1970s, targeted drug delivery by using nanoparticles was still exploring and facing lots of limitation of the nonbiodegradability of the polymers which are used for their synthesis such as polyacrylamide [52] or polymethylmethacrylate [49]. Hence, at that time, the applications of nanoparticles in medicine for systemic therapeutic treatment for humans remained an unapproachable dream. Robert A Freitas gave the term “Nanomedicine” and has established it with the publication of a book and since then it has been used in the technical literature [12].

3. TYPES OF NANOPARTICLES USED FOR THERAPEUTIC TREATMENT Nanoparticles are used for drug delivery systems are submicronsized particles size ranges from 3 to 200 nm or devices that are designed by using various materials including lipids (liposomes), polymers (polymeric nanoparticles, micelles, or dendrimers), and even organometallic substances. Nanoparticles-based drugs used for targeted delivery reduces the toxicity and side effects to improve the therapeutic index of the targeted drug. The nanodrugs are beneficial because they have similar size as biomolecules such as receptors, antibodies, and nucleic acids [53]. The approach of targeted drug delivery by nanosizing of drugs has numerous advantages as reported by McNeil [54] which has shown in Fig. 2.

Nanoparticle Drug Delivery

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3.1. Metal and Metal Oxide Nanoparticles Metal nanoparticles have a wide application not only in the medical field for drug delivery but also in electronics, optics, fluorescent materials, biosensors, and catalysts. Metal nanoparticles are used alone or as a carrier with drugs and bioactive herbal extracts and represented as a promising candidate in drug delivery applications because of their size, biocompatibility, targeted, and controlled drug release [55,56]. The nanoparticles which are commonly used for therapy are magnetic nanoparticles (iron oxide), gold and silver nanoparticles, nanoshells, and nanocages have been continuously used and modified to enable their use as a diagnostic and therapeutic agent. Iron oxide (FeO) is an inorganic compound and occurs naturally as the mineral magnetite which is superparamagnetic in nature. Superparamagnetic iron oxide nanoparticles (SPION) are using for several biomedical applications because of their ultrafine size and magnetic properties. The medical application of SPIONs such as to enhance resolution contrast agents for magnetic resolution imaging (MRI), as a drug carrier and imaging, gene therapy, stem cell tracking, molecular/cellular tracking, magnetic separation technologies (e.g., rapid DNA sequencing) hyperthermia for cancer treatment, early detection of inflammatory, cancer, diabetes, and atherosclerosis [57–66]. Polyethyleneimine (PEI)-modified magnetic nanoparticles (GPEI) are used as a potential carrier targeted delivery of vascular drug/ gene to brain tumors [67]. According to the studies by Alexiou et al., complete tumor reduction in tumorbearing rabbits have been observed by using magnetic

FIG. 2 Diagrammatic representations of advantages of nanosizing of drug.

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nanoparticle without any negative side effects. Additionally the applied dose of drug decreased to 20% of the regular systemic dose [68,69]. Lubbe et al. studied the treatment of Phase I human clinical trial which showed positive results of the physiological tolerance of magnetic drug (4-epidoxorubicin) by patients [70]. Gold and silver nanoparticles are ideal for various pharmaceutical applications and drug delivery because of their stability, noncytotoxicity, inert nature, high disparity, and biocompatibility [71]. Previous studies showed the treatment of intracellular infections with conjugates of gold nanoparticles and antibiotics provide promising results [72,73]. The association of gold nanoparticles in antibiotic treatment enhances the efficiency of drug delivery to target cells [74,75]. The amount of dose required in this therapy is higher than the actual amount required for pathogenic treatment. The overdose of antibiotics can cause adverse effects [76] because of this gold nanoparticle with antibiotics conjugates helps targeted delivery and reduces the amount of dose with improving antibiotic efficacy. Chen et al. studied methotrexate drug conjugated to 13-nm colloidal gold. Methotrexate is an anticancerous drug an analog of folic acid that can destroy cells folate metabolism [77]. Li et al. demonstrated the study of functionalized gold nanoparticles (AuNPs) which showed an important role in efficient drug delivery and biomarking of drugresistant leukemia K562/ADM cells. It also indicated that the AuNPs interaction with biologically active molecules on the leukemia cell surface may contribute to the observed improvement in cellular drug uptake [78]. The conjugation of gold nanorods and small interfering RNA (siRNA) by electrostatic binding has used for targeted delivery of siRNA to specific cells or tissues [79]. Silver nanoparticles (AgNPs) are used for the treatment of diseases by targeting specific cells, such as interacting with the HIV-1 virus and inhibiting its ability to bind host cells in vitro [80]. Bhattacharya and Mukherjee studied nanocrystalline Ag and Ag sulfadiazines that are used for wound healings to treat ulcers and to treat burn wounds, respectively, in the form of pastes or creams [81]. Some metallic nanoparticles are used in the form of metal oxides NPs such as zinc oxide nanoparticles (ZnONPs) and titanium dioxide nanoparticles (TiO2NPs) as a skin protector in sunscreen [82]. The ultraviolet radiations (UV) have the ability to damage DNA in human skin cells by inducing oxidative stress and it also plays an important role in the cause of pathogenesis of melanoma and nonmelanoma skin cancer [83]. Therefore for the protection of the skin from cancer development against both UVA and UVB radiation, some metallic nanoparticles are used in the form of metal oxides NPs such as ZnONPs and TiO2NPs as skin protectors [82].

Cobalt oxide nanoparticles (CoONPs) have different physical and chemical properties such as catalytic, magnetic, and optical properties. Additionally, CoONPs have anticancerous properties and these nanoparticles have excellent uptake of cancerous cells after surface modification when attached with amide. Chattopadhyay et al. [84] showed doxorubicin and methotrexate-attached folic acid PMIDA-coated CoONPs used as a carrier for targeted anticancer drug delivery.

3.2. Chitosan Nanoparticles Chitosan is a naturally founded polysaccharide, highly basic, cationic, mucoadhesive biocompatible polymer extracted from crustacean shells of crabs or prawns and cell walls of fungi and used as drug carrier and in tissue engineering approved by US Food and Drug Administration (FDA) [85]. Chitosan NPs are useful for slow/controlled drug release and they can cross biological barriers in vivo and delivering drugs to enhance the efficacy of the targeted site which improves drug stability and solubility, enhances efficiency, and reduces toxicity [86]. The function for drug delivery of chitosan NPs and poly(D,L-lactide-co-glycolide) (PLGA) nanoparticles [87] is similar but the chitosan NPs drug delivery is more pH dependent [88]. Mohammed et al. studied exendin-4 loaded PLGA and chitosan-PLGA NPs for the treatment of type-2 diabetes [85]. Tamoxifen-loaded lecithin-chitosan NPs increased the solubility and permeation of drug across the intestinal epithelium and it is useful for oral cancer drug delivery [89]. Liu et al. showed that Carbamazepine drug which is used for epilepsy treatment, its brain-to-plasma exposure ratio reached to 150% and also enhances the bioavailability of drug to cross the blood brain barrier (BBB) when loaded with chitosan NPs [90].

3.3. Solid Lipid Nanoparticles Solid lipid nanoparticles (SLNPs) are emerging pharmaceutical delivery system submicron size range 50–1000 nm are used as alternative carriers to colloidal systems, for controlled and precise drug delivery [91,92]. SLNPs are synthesized from biocompatible and biodegradable materials which have an ability to carry or localize the lipophilic and hydrophilic drugs in the solid lipid matrix [93,94]. There are several drugs which have been incorporated into SLNPs [95] for therapeutic purposes such as desoxycortisone [96], timolol [96,97], idarubicin, doxorubicin [98], [D-Trp-6]LHRH [99], pilocarpine [100], thymopentin [101], oxazepam, diazepam, cortisone, betamethasone valerate, retinol, prednisolone, retinol, menadione, ubidecarenone [102], 30 -azido-30 -deoxythymidine

CHAPTER 11 palmitate [103,104], acyclovir [105,106], azidothymidine palmitate [104], gadolinium (III) complexes [107], cyclosporine [108], etomidate, tetracaine [109], vitamin E palmitate [110], progesterone, hydrocortisone [111], and camptothecin [112].

3.4. Mesoporous Silica Nanoparticle In 2011 the US FDA approved silica nanoparticles for Phase I human clinical trials, it was an important initiative toward clinical acceptance of silica nanoparticles [113]. Mesoporous silica materials which consist of porous structure with hundreds of empty channels such as honeycomb (mesopores) structure used for drug delivery such as MCM-41 (Mobil Composition of Matter) and SBA-15 (Santa Barbara University mesoporous silica material). This gives the possibility of using these nanomaterials in a combined drug delivery therapy. These porous structures can encapsulate large amount of drug molecules [114], and it has made silica nanoparticles (SiNPs) highly attractive carrier because of wide applications of nanotechnology to enhance the characteristics of nanoparticles such as adsorption, sensing, catalysis, and separation [115–120]. Mesoporous silica nanoparticles (MSiNPs) structure enables adsorption of DNA and gene transfer [114].

3.5. Liposome Nanocarrier Liposomes vesicles made up of single or numerous lamellae layers (lipid bilayers) consist of the aqueous core inside suitable for encapsulation [121]. Liposomes are clinically developed nanocarriers to deliver genes, cytotoxic and antifungal drugs, vaccines, and imaging dye [122] and it is also used to enhance the functionality of several types of nanoparticles such as hydrophilicity, stability in plasma, controlled delivery, and improved biocompatibility. Additionally, the cationic liposomes can interact with oppositely charged molecules that are attached to them on the surface. This property makes liposomes suitable to conjugate with ligands or antibodies for targeted delivery [123]. Some liposomebased drugs such as anthracyclines doxorubicin (Doxil and Myocet), liposome-encapsulated curcumin, and albumin-paclitaxel nanoparticles were suitable for the treatment of cancer and liposomal daunorubicin (DaunoXome) for AIDS-related Kaposi’s sarcoma [124–127].

3.6. Polymeric Nanocarriers Polymeric nanoparticles are biodegradable and biocompatible size ranges of 10–1000 nm. Drugs can be physically adsorbed on the surface of the polymer or chemically linked to the surface and encapsulated in the core of polymer nanocarriers [128]. Biodegradable

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polymers that are used for encapsulation of a variety of therapeutic compounds are poly(D,L-lactic-co-glycolic acid) (PLGA), poly(ε-caprolactone), poly(D,L-lactic acid) (PLA), and poly(ethylene glycol) (PEG). In South Korea for Genexol-PMTM, a PLGA-b-methoxy PEG NP encapsulating paclitaxel has received regulatory approval for cancer treatment and is undergoing phase II clinical trials in the United States [129]. Moreover, these polymer nanoparticles are not only suitable for delivery of small molecule drugs but with the array of polymer and surface modification techniques, it can also be used to deliver proteins [130], diagnostic agents [131], and nucleic acids [132].

3.7. Dendrimer Dendrimers are roughly large spherical, threedimensional branched structures that have a typically symmetric core, size ranges of 10 nm due to which drug incorporation into dendrimers can be limiting [133–135]. Besides drugs, it is also used to deliver genes, sensors, and killing the bacterial cell [135–137]. Poly(amidoamine) or PAMAM (polypropylenimine dendrimers) are the commonly used dendrimers. PAMAM dendrimers conjugated with anticancerous drug cisplatin showed controlled release with minimum toxicity and high accumulation in solid tumors as compared with free cisplatin [138]. According to Chauhan et al., [139] PAMAM dendrimer-loaded indomethacin enhances the bioavailability of nonsteroidal antiinflammatory drug indomethacin in transdermal delivery applications.

3.8. Polymeric Micelles Nanoparticles Micelles are lipid molecules that organize themselves in a circular form in aqueous solutions and exhibit a core and shell structure. Its morphology has significantly enhanced the impact of pharmacokinetic properties [140]. The core is hydrophobic in nature which is used for the encapsulation of drug whereas the hydrophilic shell provides aqueous solubility and steric stability to the micellar structure [141]. Depending on the solvent environment and relative length of hydrophobic/hydrophilic blocks micelles form a variety of shapes such as vesicles, spheres, tubules, rods, and lamellae [142–144]. Polymeric micelles nanoparticles are broadly used in experimental studies as a carrier for the delivery of poorly water-soluble drugs. Micelle carrier enhances the solubility and reduces early degradation of anticancerous drugs and accumulates at the tumor site. Micelles nanocarrier which are used in clinical trials or in clinical use are mPEG-PLA (paclitaxel) for breast cancer [145], Pluronic L61 and F 127 polymeric micelle (doxorubicin) for lung cancer [146], PEG-PGA polymeric micelle (Cisplatin) for

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Phase II/III solid tumors, gastrointestinal, and genitourinary cancers [147].

4. TOXICOLOGICAL PROFILE OF NANOPARTICLES The application of nanoparticles in pharmaceuticals is increasing in the last few years to reduce drug doses and their side effects. However, these nanoparticles themselves may eventually cause the risks of toxicity in patients. Nanoparticles sizes are similar to cellular organelles therefore nanoparticles can enter inside the human body through dermal, oral, inhalation, intravenous, and intraperitoneal routes [148]. It can induce cytotoxicity, oxidative stress, genotoxicity, and inflammatory responses inside the body therefore before application of nanoparticles as a drug carrier it should make biocompatible specifically metallic nanoparticles [148–152]. Agglomeration of nanoparticles may also occur during the drug delivery in the vascular system which might lead to blockage and further could induce toxicity [153,154]. Some previous studies reported that the accumulation of cationic liposomes in vital organs can also lead to toxicity by disrupting cell membrane function and attaching to serum proteins [155,156]. Therefore, it is necessary to methodically study the toxicological profile of nanoparticles in human before their applications used in medicine.

5. CONCLUSION AND FUTURE DEVELOPMENT The recent advancement in nanomedicine uses a wide range of nanoparticles for drug delivery, imaging and biosensing (nanosensor) to cure incurable diseases. These multidirectional approaches of nanoparticles have revolutionized the nanomedicine. However, nanoparticles applications for drug delivery have both beneficial and harmful effects on human health. Several nanoparticles have been evaluated clinically, but still, some limitations are there which needs more experimental studies. Standard methods should be established to encourage the benchmark materials, analyze both the short-term and long-term effects and quality management for the preclinical characterization of nanoparticles used for drug delivery.

ACKNOWLEDGMENTS The first author expresses her gratitude to the Department of Science and Technology, India, for funding scholarship (INSPIRE-IF140412). The authors are thankful to the Head, Department of Zoology, for the facilities and administrative support.

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CHAPTER 12

Targeted Drug Delivery: Advancements, Applications, and Challenges

HOSSEIN RAHIMIa • SOODABEH DAVARANb • HAMED NOSRATIc • HOSSEIN DANAFARc a

Department of Medical Biotechnology, School of Medicine, Zanjan University of Medical Science, Zanjan, Iran, bDrug Applied Research Center, Tabriz University of Medical Sciences, Tabriz, Iran, c Department of Pharmaceutical Biomaterials, School of Pharmacy, Zanjan University of Medical Sciences, Zanjan, Iran

1. INTRODUCTION Specialized and controlled delivery of the drugs to the body is one of the most important issues in the treatment of various diseases such as cancer. In classical drug delivery methods, the drug is distributed throughout the body due to the nonspecific targeting ability of drugs, which affects both healthy and diseased sites of the body. Actually, in most cases, the drug-injected site is far from the diseased area, which results in widespread drug distribution throughout the body. Therefore they cause many side effects and unwanted effects. Although in classical drug delivery methods, a high dose of a drug is entered into the body, there is less therapeutic effect than the taken dose. This challenge in the treatment of cancer with chemotherapy also causes severe side effects, damage to normal cells and tissues around the cancer cells, and weakening the body. The main goals in drug delivery are efficient and specific drug delivery from the administered site to the diseased site, minimizing side effects, reducing costs and affecting on diseased cells without the least damage to adjacent normal cells. Given the challenges in the conventional drug delivery methods, new techniques need to be developed for specific and targeted delivery of drugs to the diseased area in the body. The advent of medical nanotechnology and targeted drug delivery techniques has revolutionized in drug delivery and treatment of various diseases. Nanotechnology-based drug carriers have received a great deal of attention in recent years due to the unique properties of nanoparticles in the treatment of various diseases. A variety of nanoparticles (gold nanoparticles, iron oxide nanoparticles, protein nanoparticles, polymeric nanoparticles, nanogels, and liposomes) are used as drug carriers in targeted drug delivery [1]. Targeted drug delivery methods have many advantages against

classical drug delivery methods, including specific and targeted delivery of the drug to the diseased site without any effect on normal cells, which in nontargeted methods, normal cells also affected by administrated drugs. A further advantage of targeted drug delivery is the need for a lower dose of the drug, which uses a much lower dose than the nontargeted methods with a greater therapeutic effect. Nanoparticles provide a longer treatment effect. Nanoparticles can also release the drug in a controlled manner, for example, under sensitive conditions to specific stimuli such as pH, temperature, or other conditions [1, 2]. Generally, active targeting and passive targeting are the main targeting strategies for targeted drug delivery. In active targeting the drug can be specifically delivered into target cells by binding ligands (which have overexpressed receptors on the surface of the target cells) on the drug formulations surface. In passive targeting, drug-containing nanoparticles accumulate in the surrounding area of the diseased site due to enhanced permeability and retention (EPR) because of the different characteristics of the tumor tissue environment than normal areas [3]. This chapter is organized into four main sections. The first section is an introductory to the classical drug delivery methods, their challenges, the emergence of targeted drug delivery methods and its importance, as well as the role in the development of various therapeutic modalities. The second section introduces active targeting strategy; we learn how to deliver a drug formulation specifically to the target area. Also, focus on targeting mediated highly expressed receptors on the surface of defined cells. The second section reviews targeting strategies by peptide ligands, folic acid (FA), and aptamers. Section 3 discusses passive targeting and a number of studies on passive targeting. Section 4 provides a

Modeling and Control of Drug Delivery Systems. https://doi.org/10.1016/B978-0-12-821185-4.00011-7 © 2021 Elsevier Inc. All rights reserved.

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number of studies that compared the efficiency of active and passive targeting.

2. ACTIVE TARGETING It is very important that in the treatment of cancer or other diseases, drugs or drug formulations reach the target tissues and cells in a very specific way and do not have any adverse effect on the surrounding normal cells. In active drug targeting, drug-containing nanoparticles reach specifically to the target tissue or cells by attaching ligands that have over-expressed receptors on the target cells surface. Cancer cell receptors that are overexpressed on cancer cells to receive nutrients are different from healthy cell receptors [4, 5]. Many of these receptors have been identified and their associated ligands have been made [3]. Interaction between attached ligands to the nanoparticles surface and related receptors on the cell surface facilitates the uptake and internalization process of drug-containing nanoparticles into the cells [6]. Actually, using this targeting strategy increases the specificity of drug delivery and avoiding the side effects.

2.1. Receptor-Mediated Active Targeting The development of targeted drug delivery methods has led to increased drug uptake by the target cells [7]. One amazing strategy is to bind ligands to the drug carrier that interacts specifically with the target region (receptors that are overexpressed on the diseased cells) and thus the drug is specifically delivered to the target region. Drug delivery through targeting of receptors on the cell surface is a wonderful way to specific drug delivery and also to accumulate drug in the diseased site [8]. A key issue that ensures the safety of the targeted drug is the high expression of the targeted receptor in the target cells relative to the surrounding normal cells. Target receptor expression levels must be more than three times higher than a normal condition to deliver the drug formulation specifically to the target region [9]. In selecting the appropriate ligand for targeting a receptor, the ligand size is an important factor because this feature may lead to a good or bad event in conjunction with other factors [10–12]. In receptor-mediated drug delivery, a thorough understanding of the structure and biochemistry of the target receptor is essential. A number of receptors that can be used as therapeutic targets on different types of cancer cells including FA receptor, integrin αvβ3 and epidermal growth factor receptor (EGFR), which have high expression on the cancer cells surface, such as lung, breast, ovary, brain, and colon cancer.

2.1.1. Folic acid receptor The FA receptor (40-40 kDa) is a member of the glycoprotein family and has three isoforms including alpha (α), beta (β), and gamma (γ). The first two isoforms are attached to the cell membrane but the third isoform is found in hematopoietic cells [9, 13, 14]. Actually, the FA receptor is a cysteine-rich glycoprotein and its normal function is to bind to the FA and internalize it into the cell. FA is essential for the synthesis, methylation, and repair of DNA molecules. The alpha isoform of the FA receptor is overexpressed in a number of cancers on the surface of cells [15, 16]. As will be discussed in this chapter, FA receptor on cancer cells has been studied as a therapeutic target in a number of studies.

2.1.2. Integrin αvβ3 Integrins are cell surface proteins that are involved in cell attachment to the extracellular matrix (ECM). Moreover, integrin receptors play an important role in sending messages to cells, regulating cell morphology, cell migration and also metastasis of cancer cells. Integrins are heterodimers of alpha and beta subunits. The αVβ3 integrin which expressed by platelets is composed of two parts: integrin alpha V and integrin beta 3 and is a receptor for vitronectin. Abnormal expression of v3 is associated with the prevalence of many diseases so it can be used as a therapeutic target in the treatment of various diseases [17–20].

2.1.3. Epidermal growth factor-receptor EGFR, also known in human as human epidermal growth factor receptor 1 (HER1) or ErbB-1, is a transmembrane protein that is a member of the ErbB family of tyrosine kinase receptors. This receptor is often highly expressed on the surface of epithelial cancer cells. Binding of ligands such as EGF to EGFR leads to EGFR dimerization and tyrosine autophosphorylation, which activates intracellular pathways that are critical for the maintenance of malignant phenotypes. Defective signaling of tyrosine kinase receptors, especially EGFR, leads to diseases such as Alzheimer’s disease, whereas overexpression of these receptors is associated with cancer [21–26]

2.2. Peptides Peptides are either synthetic or originate from nature and have wide applications in drug delivery, cancer treatment, and diagnosis. Peptides can be used in a variety of roles, such as antibiotics, inhibitors, and hormones [27]. Peptides have attracted much attention because of their superiorities such as small size, lower immunogenicity, stability, easy synthesis and low cost, and most importantly easy attachment on the surface

CHAPTER 12 Targeted Drug Delivery of nanoparticles for targeting [28]. The RGD peptide is one of the most common peptides used in targeting nanoparticles and other drug formulations. RGD (arginine-glycine-aspartic acid) is one of the most common motifs found in extracellular matrix proteins including fibronectin, fibrinogen, vitronectin, osteopontin, and others, also a cell adhesion protein called integrin binds to this peptide [29–31]. Multifunctional systems established by combining liposomal nanoparticles with magnetic nanoparticles called magnetoliposomes have various applications such as drug delivery and imaging. In a study conducted by Belderbos et al., the ability and targeting property of magnetoliposomes was used to target αvβ3 integrins that are overexpressed on the different cancer cells surface. First, Texas red fluorophore and cyclin RGD to target αvβ3 integrin were loaded onto synthesized magnetoliposomes. To create a mice model for evaluating the cRGD-MLS, SKOV-3 ovarian cancer cells were injected subcutaneously into mice. Three weeks after injection and tumor growth, the prepared nanosystem biodistribution at certain times was monitored by T7 magnetic resonance imaging (MRI) scanner and fluorescence imaging (FLI). Tumoral uptake of designed system was confirmed by FLI, electron paramagnetic resonance (EPR) spectroscopy, and histopathology. It was observed that the uptake of cRGD-MLs was higher and the highest uptake was 4 h after administration. Overall the results of this study demonstrate the ability of the cRGD-MLs system to target the SKOV-3 xenograft [32]. In another study, Lim et al. prepared the L121/ F127 hybrid micellar system and loaded docetaxel onto the system and investigated the anti-cancer effects of the prepared system. They linked the RGD peptide to the prepared system to increase targeting specificity and therapeutic efficacy of the system. Binding of the cRGD ligand to the prepared system resulted in increased cellular uptake and increased anti-cancer efficacy on U87MG cancer cells expressing αvβ3 integrin. Also, the cRGD ligand increased the accumulation of drugcontaining system in the tumor area [33]. Theranostics application of gadolinium-based small rigid particles (SRPs) bound to the cRGDfK peptide sequence (containing the RGD motif ) investigated in vitro on HEK293 (β3) and U87MG cells and in vivo in mice bearing U87MG tumor. The importance of the RGD motif in the synthesized particles was observed in targeting cells with high expression of αvβ3 integrin [34]. Liu et al. developed a new series of dual-ring RGD peptides to enhance αvβ3 integrin targeting efficiency [35]. In another work, for specific delivery of the drug to monocytes/neutrophils in the brain, the RGD peptide was bound to ferulic acid liposomes [36].

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Liposome-based drug delivery systems have attracted much attention in recent years as high-capacity drug carriers [37, 38]. RGD-conjugated drug-loaded nanoliposomes have been widely used to target cells with high expression of αvβ3 integrin. However, one of the major challenges for this group of drug carriers is the trapping by the reticuloendothelial system (RES) system as well as uptake by other cells, which leads to unwanted effects and toxicity to other cells. Therefore it is necessary to design new systems of nanoliposomes bound to the RGD peptide that can overcome these challenges. To this end, Amin et al. in a study focused on the design and synthesizing of a liposome-based system linked to the RGD peptide to enhancing target specificity and tumor uptake. They prepared the sterically stabilized liposomal doxorubicin complex (SSLD) and then bound the cRGDfK and RGDyC peptide sequences to the prepared complex (RGD-SSLD). It was found that increasing hydrophobicity of the peptides resulted in the increased therapeutic efficacy of the RGD-SSLD complex in C-26 tumor-bearing mice and also reduced the trapping by RES system and adverse effects. Following, the ability of the N-methylated RGD peptide to targeting cells with high expression of αvβ3 integrin was evaluated in both in vitro and in vivo conditions. The ability of RGDf [N-methyl] C-liposome to target and affinity to B16F0 and BLM tumor cells was confirmed both in vitro and in vivo (Fig. 1) [39]. It has also been shown that gold nanoparticles if linked to the RGD peptide by the poly(ethylene glycol) (PEG) chain and oligolysine linker can efficiently target colorectal cancer cells (Fig. 2) [40]. Attaching targeting agents to the drug carrier is an important and valuable strategy to increase therapeutic efficacy and reduce side effects. Hepsin is highly expressed in a number of cancer cells, and it has been shown that the IPLVVPL peptide is a high-affinity ligand for cancer cells expressing hepsin. In an attempt, the IPLVVPL peptide was bound to the pH-redoxsensitive and containing doxorubicin hydrogels. Conjugation of this peptide onto hydrogels resulted in high uptake efficiency by cells expressing hepsin (Fig. 3) [41]. Drug delivery to the brain is one of the major challenges in drug delivery. The brain is protected by an efficient shield called the blood-brain barrier (BBB). BBB is able to protect the brain cells against the blood contents and toxic compounds in it. But this shield is a barrier to the entry of drugs into the brain, which remains a major challenge. However, molecular carriers known as BBB shuttles are promising in drug delivery to the brain. Peptide-based shuttles have attracted a great deal of

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FIG. 1 Schematic illustration of RGDf [N-methyl] C-liposomes targeting and biodistribution of RGDf[N-met] C-SSLD and Plain-SSLD at different time points in serum and tumor. The upper curve shows the rate of tumor growth in mice carrying B16F0 tumor after IV injection of either sucrose 10% or 15 mg/kg doxorubicin encapsulated in RGDf[N-met]C-SSLD or Plain-SSLD, and the bottom curve shows the tumor growth rate of female BALB/c mice bearing C-26 tumor after an IV injection of 15 mg/kg doxorubicin in RGDf[N-met]C-SSLD or Plain-SSLD or empty RGDf[N-met]C-SSL (8.1 μmol/kg lipid) or sucrose 10%. (Reprinted from M. Amin, et al., Development of a novel cyclic RGD peptide for multiple targeting approaches of liposomes to tumor region, J. Control. Release 220 (2015) 308–315 with permission from Elsevier.)

FIG. 2 Preparation of gold NPS-PEG-RGD complex, and their targeting activity toward SW620 cells.

(Reprinted from F. Biscaglia, et al., Gold nanoparticle aggregates functionalized with cyclic RGD peptides for targeting and imaging of colorectal cancer cells, ACS Appl. Nano Mater. 2 (10) (2019) 6436–6444 with permission from American Chemical Society.)

CHAPTER 12 Targeted Drug Delivery

FIG. 3 Schematic illustration of hepsin-targeting hydrogels

and cellular uptake. (Reprinted from B. Xue, et al., Peptidefunctionalized hydrogel cubes for active tumor cell targeting, Biomacromolecules 19 (10) (2018) 4084–4097 with permission from American Chemical Society.)

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without GSH leads to increased brain uptake of this peptide [45]. In addition, specific uptake of GSHconjugated PEGylated liposomes by the receptor revealed the ability of GSH in brain targeting [46]. PEG-modified polyamidoamine dendrimer (PAMAM) conjugated to the HAIYPRH peptide (T7) and loaded with doxorubicin targeted transferrin receptors which are highly expressed on tumor cells. It was found that the presence of T7 ligand increased the cellular uptake of nanoformulation mediated by transferrin receptor and also increased the accumulation of doxorubicin in the tumor site compared to the T7-free nanocomplex. The T7-conjugated nanocomplex also showed more tumor growth inhibition than the T7-free nanocomplex, indicating that the T7 peptide could be used as an effective ligand in active targeting (Fig. 4) [47].

2.3. Folic Acid attention in the delivery of drugs to the brain because of their unique features such as cheapness and low immunogenicity. Glutathione (GSH) is a shuttle peptide that crosses the BBB has reached clinical application stages [42]. In this regard, Nosrati et al. developed an efficient system for the delivery of paclitaxel (PTX) from the BBB by synthesis of magnetic nanoparticles and the conjugation of GSH on the surface of synthesized nanoparticles, and also monitored the delivery by MRI [43]. Also in a series of studies, GSH-PEGylated liposomes were developed and studied to increase the delivery efficiency of the opioid peptide DAMGO (H-Tyr-d-Ala-Gly-MePheGly-ol) [44]. Another work showed that encapsulation of DAMGO peptide in PEGylated liposomes with or

Li

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FA (vitamin B) is a water-soluble vitamin with a molecular weight of 441 Da. This biomolecule is essential for the synthesis, methylation and repair of DNA molecules. FA is stable at various pH and temperatures and also retains its ability to binding to its receptor on the cell, when bound to other molecules and drugs [48]. The FA receptor, because of its high expression on cancer cells, can be used to specifically target in cancer treatment. Actually, the FA receptor on cancer cells is an important biomarker known for cancer cells. This receptor enables the detection of cancer cells through specific imaging and the treatment of cancer cells (targeted drug delivery). Actually, by binding FA to a formulation, it can be targeted specifically to cancer cells that have a

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Control DOX 5 mg/kg PAMAM-PEG/DOX 2 mg/kg PAMAM-PEG-T7/DOX 2 mg/kg

low

FIG. 4 The biodistribution of DOX in tumor-bearing mice after injection of saline as a control, free DOX,

PAMAM-PEG/DOX, and PAMAM-PEG-T7/DOX. Br, brain; H, heart; Li, liver; Lu, lung; K, kidney; Sp, spleen; Tu, tumor. (Reprinted from L. Han, et al., Peptide-conjugated PAMAM for targeted doxorubicin delivery to transferrin receptor overexpressed tumors, Mol. Pharm. 7 (6) (2010) 2156–2165 with permission from American Chemical Society.)

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high expression of the FA receptor. Superparamagnetic iron oxide nanoparticles (SPIONs) coated with PEG and PEI and conjugated with FA (for specific targeting of cancer cells) and loaded with Doxorubicin (DOX) showed efficient uptake by MCF-7 cells. Intravenous administration of these nanoparticles to tumor-bearing mice was found to inhibit tumor growth very effectively. It was also observed that the use of magnetic field to direct nanoparticles increased the efficiency of growth inhibition of MCF-7 cancer cells in both in vitro and in vivo (Fig. 5) [49]. Iron oxide nanoparticles (IONPs) conjugated with FA (as a nanoparticles coating and specific targeting agent for prostate cancer cells) were prepared. It was observed that FA-conjugated nanoparticles

were specifically uptake by the target cells. In addition, these nanoparticles showed a high ability to detecting by MRI and by hyperthermia in treatment [50]. Mohapatra et al., in a study by synthesizing magnetic nanoparticles and conjugating FA to synthesized nanoparticles, observed the FA receptor-mediated uptake in target cancer cells [51]. Gold nanoparticles conjugated with FA through GSH also showed specific uptake by cancer cells that had high expression of FA receptor, whereas uptake was not observed in cells without FA receptor expression [52]. Specific uptake of FAconjugated cellulose nanocrystals (CNCs) by target cancer cells with high expression of the FA receptor was also observed in another study (Fig. 6) [53].

FIG. 5 Schematic illustration of preparation of SPIONs-PEG-PEI-DOX system and the DOX-loaded system

for targeted drug delivery to the tumor and the MRI. (Reprinted from Y. Huang, et al., Superparamagnetic iron oxide nanoparticles conjugated with folic acid for dual target-specific drug delivery and MRI in cancer theranostics, Mater. Sci. Eng. C 70 (2017) 763–771 with permission from Elsevier.)

FIG. 6 The chemical structure of folic acid-conjugated cellulose nanocrystals (CNCs), and cellular image

of treated cell with folic acid-conjugated CNCs. (Reprinted from S. Dong, et al., Synthesis and cellular uptake of folic acid-conjugated cellulose nanocrystals for cancer targeting, Biomacromolecules 15 (5) (2014) 1560–1567 with permission from American Chemical Society.)

CHAPTER 12 Targeted Drug Delivery Jiang et al. targeted the KB cells (with high expression of folate receptor) by FA-conjugated magnetic Fe3O4 nanoparticles (as MRI contrast enhancer) for hyperthermia and MRI aims [54]. Theranostics applications of InP/ZnS quantum dots conjugated with FA, D-glucosamine, or both of them have also been investigated [55]. It has also been observed that the presence of FA in the carboxymethyl chitosan nanoparticles loaded with the anticancer drug doxorubicin resulted in specific uptake by HeLa and B16F1 cells, thereby enhancing the therapeutic efficacy of the synthesized nanoparticles [56]. Evaluation of the ability of FA as a targeting factor on curcumincontaining Fe3O4 magnetic nanoparticles in tumorbearing mice indicates that FA-conjugated nanoparticles have better therapeutic efficacy than non-FA-targeting particles [57]. Nosrati et al. synthesized FA-conjugated, and curcumin-loaded Bi2S3@BSA nanoparticles (Bi2S3@ BSA-FA-CUR) and investigated the therapeutic efficacy of the established nanocomplex both in vitro and in vivo conditions. In this study bismuth-based nanoparticles have been used as an enhancer for X-Ray therapeutic efficacy as well as curcumin carrier due to their high ability to increase X-Ray sensitivity. It was observed that approximately 3 weeks after the injection of Bi2S3@ BSA-FA-CUR complex into tumor-bearing mice, the tumor was completely removed, indicating the high therapeutic efficacy of the synthesized nanocomplex and the accumulation of particles in the tumor (Fig. 7) [58].

2.4. Aptamer Aptamers are small fragments of peptide or oligonucleotide sequences that are capable of specific binding to the target molecule by creating three-dimensional structures such as antibodies [59, 60]. Generally, aptamers can be divided into two classes: (1) oligonucleotide

201

aptamers (RNA, DNA, or XNA) consisting of short oligonucleotide strands; (2) peptide aptamers are synthetic proteins that consist of one or more peptide domains. Peptide aptamers bind to cellular protein targets and perform biological functions in vivo. Peptide aptamers can also identify and bind to specific targets in vitro. A number of aptamers exist naturally and some are produced for a specific purpose [61]. Because of their small size and unique properties, aptamers have superiorities over antibodies, including (1) aptamers show faster and more efficient penetration into the tissues because of their low molecular weight (8–25 kDa); (2) since aptamers are oligonucleotides, they are therefore not recognized by the immune system and do not result in the immune response; (3) aptamers are stable at higher temperatures due to their oligonucleotide property; and (4) low cost of producing aptamers compared with antibodies. In general, aptamers can identify and bind to a wide range of targets such as proteins, viruses, cells, and drugs [61–64]. Tetrahedral DNA nanostructures (TDNs) are programmable and controllable structures that are used in a variety of fields including drug delivery. These DNA nanostructures have been considered promising carriers in drug delivery because of their properties (biocompatibility, biodegradability, penetrable the cell membrane, and capable of functionalizing with different groups). To increase the efficiency and capability of these DNA nanostructures for drug delivery, the AS1411 aptamer was attached to TDNs, and the efficiency of the developed system was compared with TDNs for drug delivery in different cells. AS1411 has the ability to bind specifically to nucleolin, which is highly expressed on tumor cells. It has been found that AS1411 containing TDN is more efficient in tumor cells targeting for drug delivery than free TDN and also capable of inhibiting tumor cells growth (Fig. 8) [65].

FIG. 7 Schematic illustration of Bi2S3@ BSA-FA-CUR nanocomplex therapeutic mechanism. (Reprinted from H. Nosrati, et al., Tumor targeted albumin coated bismuth sulfide nanoparticles (Bi2S3) as radiosensitizers and carriers of curcumin for enhanced chemoradiation therapy, ACS Biomater. Sci. Eng. 5 (9) (2019) 4416–4424 with permission from American Chemical Society.)

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FIG. 8 Comparison of the effect of TDNs and AS1411-TDNs on the growth of MCF-7 and L929 cancer cells. (Reprinted from Q. Li, et al., Aptamer-modified tetrahedral DNA nanostructure for tumor-targeted drug delivery, ACS Appl. Mater. Interfaces 9 (42) (2017) 36695–36701 with permission from American Chemical Society.)

AS1411 was also conjugated to porphyrin derivatives TMPyP4 (AS1411-TMP) to enhance the delivery efficiency to MCF7 breast cancer cells. It was observed that the AS1411-TMP complex resulted in the highest level of TMPyP4 accumulation in the target cancer cells compared with normal cells, indicating the importance of AS1411 in specific targeting of nucleolin on the surface of the cancer cells [66]. AS1411 in an attempt used in complexed with acridine-based G-quadruplex ligand, C8, to deliver C8 to HeLa cancer cells [67]. AS1411 in combination with DNA pyramids also resulted in

increased uptake by target cells and specific inhibition of the target cancer cells growth [68]. The complex consists of gold nanoparticles (GNPs) and prostate-specific membrane antigen (PSMA) RNA aptamer (which binds specifically to PSMA) and doxorubicin for targeting prostate cancer cells as well as the ability of the developed complex as a CT imaging contrast agent was evaluated. It was found that the developed complex resulted in a fourfold increase in CT intensity and showed significant potency against the targeted cancer cells (Fig. 9) [69].

FIG. 9 Schematic illustration of the gold NPs and PSMA RNA aptamer and DOX complex. (Reprinted from

D. Kim, Y.Y. Jeong, S. Jon, A drug-loaded aptamer gold nanoparticle bioconjugate for combined CT imaging and therapy of prostate cancer, ACS Nano 4 (7) (2010) 3689–3696 with permission from American Chemical Society.)

CHAPTER 12 Targeted Drug Delivery

3. PASSIVE TARGETING In the passive targeting method, drug-containing nanoparticles (drug formulation) accumulate in the tumor environment by EPR phenomenon due to different characteristics around the tumor tissue compared with normal cells and tissues [70]. Actually, this targeting method is used when the target tissue has different properties than other cells and tissues [70]. The EPR effect occurred mainly due to (1) high vascular permeability of malignant tumor tissues and (2) lack of lymphatic drainage within tumors. The enhanced permeability of the tumor vasculature allows macromolecules, lipids, and nanoparticles circulating in the blood to extravasate through the leaky tumor blood vessel, then enter the tumor interstitial space. Therefore by binding the drug to a suitable carrier, the drug accumulation in the target area can be increased up to 100-fold [4, 71]. The delivery of the nanoparticles and drug to the target region is related to factors such as tumor microvasculature, size, shape, and surface charge of the nanoparticles [70]. One of the most important issues in the passive targeting method that guarantees success in drug delivery and therapeutic efficacy is the long-term circulation of the drug. Nanoparticles are usually eliminated in the bloodstream by the RES, and therefore, modification of the nanoparticle surface is essential for the long-term circulation of the drug-containing nanoparticles [72]. This is possible by coating the nanoparticles with substances such as PEG that could decrease hydrophobicity of the nanoparticles surface. The hydrophobic surface of the nanoparticles forms a hydrogen bond between the oxygen molecules of the PEG coating with the water molecules, which creates a hydration film around the nanoparticles that prevents the drug-containing nanoparticles from being removed by the phagocytic system. One example of PEG coating using to prolong the circulation of drug-containing nanoparticles was doxorubicin liposomal PEGylation, which lasted from minutes to hours. Other materials have also been used for coating nanoparticles, including polaxamer, polyvinyl alcohol, poly (amino acid) s and polysaccharides [73–76]. Albumin nanoparticles are an attractive and efficient drug carrier for delivery of hydrophobic drugs to inflamed joints. In this regard, the targeting and antiinflammatory properties of tacrolimus-loaded albumin nanoparticles (TAC) were studied by Thao et al. Evaluation of the therapeutic effects of established nanocomplex on splenocytes extracted from the spleen of inflammatory model mice showed the antiproliferative effect of this nanocomplex on activated T cells. It was also observed that the established nanocomplex showed a significant antiinflammatory effect, indicating

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the efficient targeting of the inflamed region by the albumin nanoparticles as well as the accumulation of the nanocomplex in the target region (Fig. 10) [77]. The multifunctional nanosystem based on magnetic iron oxide nanoclusters containing polymer polypyrrole (PPy) and functionalized with PEG as carrier of doxorubicin was investigated for synergistic cancer treatment under NIR light. This nanosystem enjoyed from iron oxide for controlling drug delivery through the magnetic field as well as contrast imaging agent for T2-weighted MRI. PPy also has a strong photothermal effect due to its high NIR absorption, which showed high efficacy in killing cancer cells (Fig. 11) [78]. The properties of the nanoparticles can be modified by attaching polysaccharides to the particles surface. For example, by binding positively charged chitosan to the nanoparticle surface, it gives positive charge to the nanoparticles, which through electrostatic interactions can bind to negatively charged cargoes such as DNA sequences. Actually, one of the important uses of polysaccharides in giving new features to nanoparticles is to prevent the rapid clearance of the nanoparticles by the mononuclear phagocyte system (MPS) [79–82]. It was shown that lipochitosan-modified lipid nanocapsules showed more uptake by HEK293 cells. These lipid nanocapsules attached to lipochitosan and lipochitosan polysaccharides look promising for targeting ligands such as peptides or proteins as well as molecules such as siRNA [83]. Betulinic acid conjugated N-(2-hydroxypropyl) methacrylamide (HPMA) polymer was studied for tumor targeting and controlled release of betulinic acid derivatives in tumor cells [84]. In another attempt, to enhance the EPR effect, nitric oxide (NO) producing nanoparticles were developed for specific delivery of NO. In addition, doxorubicin was loaded onto the synthesized nanoparticles, and the therapeutic effect of the drug-containing nanoparticles was observed with the significant accumulation of doxorubicin in the tumor, indicating the remarkable ability of these nanoparticles to enhance the EPR effect [85]. Denis et al. developed a pH-sensitive delivery system for the specific release of Vorinostat in mesothelioma tumors as well as histone reacetylation [86]. Hoffmann et al. synthesized dualfluorescent HPMA copolymers for passive targeting of tumor and drug release [87]. THCPSi nanoparticles encapsulated in solid lipids (SLN) were examined for passive tumor targeting and drug delivery [88]. Im et al. also showed that re-injection of nanocomplex composed of reduced graphene oxide and iron oxide nanoparticles, then coated with PEG leads to the phenomenon of accelerated blood clearance (ABC), which

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Modeling and Control of Drug Delivery Systems

Normal Joint

Rheumatoid arthritis Mouse hind paw

Synoviocyte accumulation & angiogenesis

Inflamed synovial membrane

Albumin permeation  & Targeting  HSA

Bone & cartilage erosion

TAC Size = ~ 180 nm

Blood

l vesse

Inflammatory cytokines Angiogenic factors

FIG. 10 Schematic showing the injection of HSA-TAC nanocomplex to mice for targeting inflammation region and effects of nanocomplex on target tissues. (Reprinted from H.J. Byeon, et al., Pharmaceutical potential of tacrolimus-loaded albumin nanoparticles having targetability to rheumatoid arthritis tissues, Int. J. Pharm. 497 (1–2) (2016) 268–276 with permission from Elsevier.)

Emulsifier

Pyrrole

Ultrasonic dispersion

Stir at 0~5⬚C

Fe3+, Polymerization

C18PMH-PEG Ultrasonic dispersion

Triggering Drug Release NIR Light

Enhanced Cancer Cell Killing

Promoting Cellular Uptake SDS/PVA

DOX loading

® ÑB Magnetic Field Pyrrole

Fe3O4

C18PMH-PEG

DOX

FIG. 11 Preparation of Fe3O4@PPy-PEG-DOX nanosystem and drug loading and finally remotely controlled

cancer cell killing under dual physical stimuli. (Reprinted from C. Wang, et al., Iron oxide@ polypyrrole nanoparticles as a multifunctional drug carrier for remotely controlled cancer therapy with synergistic antitumor effect, ACS Nano 7 (8) (2013) 6782–6795 with permission from American Chemical Society.)

CHAPTER 12 Targeted Drug Delivery reduces the ability of passive targeting as well as shortening the shelf life of nanoparticles in the bloodstream [89]. Comparison of passive targeting of small molecules IRDye 800CW and GSH-coated luminescent gold nanoparticles in vivo showed that the nanoparticles are similar to dye molecules in early physiological stability and clearance from kidney. However, GSH-coated gold nanoparticles showed much longer tumor retention time than dye molecules. It was observed that the nanoparticles were cleared from healthy tissues approximately threefold faster than the dye molecules (Fig. 12) [90]. A number of active and passive targeting studies discussed in this chapter are summarized in Table 1.

4. COMPARISON OF ACTIVE AND PASSIVE TARGETING It has been proven that the difference between targeting strategies determined the outcomes and therapeutic efficacy of the established formulation. In this regard, two polymeric nanoparticles were synthesized and then peptide GE11 (for active epidermal growth factor receptor targeting) was attached to one of them and no agent was attached to the other and studied as passive targeting. It was shown that nanoparticles functionalized with GE11 peptide that specifically enter the target cells resulted in better therapeutic outcomes than passive targeting, indicating the high efficacy of active targeting compared to the passive targeting (Fig. 13) [91].

205

In an attempt, the efficiency of active targeting [by Arg-Gly-Asp (RGD) and Asn-Gly-Arg (NGR) peptides] was compared with passive targeting (by EPR effect). It was observed that after injection into tumor-bearing mice, passive targeting resulted in more accumulation in the tumor than active targeting. The results indicate that active targeting should not always be considered superior and more efficient than passive targeting (Fig. 14) [92]. Also in another attempt, passive targeting (by EPR effect), active targeting (by RGD peptide for αvβ3 integrin targeting), magnetic targeting by magnetic field placed in the tumor region, and also combined of active targeting with magnetic targeting were compared in terms of therapeutic efficacy, efficacy as contrast agent for MRI imaging, as well as ex vivo biodistribution using PLGA-based nanoparticles loaded with PTX and superparamagnetic iron oxide (SPIO) was evaluated. It was found that the combination of active targeting with magnetic targeting resulted in an eightfold increase in drug accumulation in the tumor area compared with passive targeting. It was observed that the therapeutic efficacy and the MRI contrast ability in combination method were increased compared with passive targeting or single targeting mode (Fig. 15) [93]. HPMA copolymer-docetaxel and HPMA copolymerdocetaxel-RGDfK complexes were also evaluated to target cancer cells [94]. Hui et al. synthesized liquid-filled silica nanocapsules with a range of hardnesses and found that there is a relation between the hardness of

FIG. 12 (A) Schematic illustration of tumor targeting and (B) IRDye 800CW structure and glutathione-coated

luminescent gold nanoparticles. (Reprinted from J. Liu, et al., Passive tumor targeting of renal-clearable luminescent gold nanoparticles: long tumor retention and fast normal tissue clearance, J. Am. Chem. Soc. 135 (13) (2013) 4978–4981 with permission from American Chemical Society.)

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TABLE 1

Summary of Active and Passive Targeting Studies.

Targeting Strategy

Targeting Ligand/ Surface Coating

Drug

Purpose

Magnetoliposomes (MLs)

Active

RGD

–

-

αvβ3 integrin targeting on different cancer cells

[32]

Pluronic blending micellar system

Active

RGD

Docetaxel

-

αvβ3 integrin targeting on different cancer cells

[33]

SRP-cRGDfK

Active

cRGDfK

–

-

αvβ3 integrin targeting on different cancer cells and tumor-imaging

[34]

RGDf [N-methyl] C-liposomes

Active

RGD

–

-

αvβ3 integrin targeting on different cancer cells and tumor

[39]

Gold NPS-PEGRGD

Active

RGD

–

-

αvβ3 integrin targeting on colorectal cancer cells and tumor-imaging

[40]

IPLVVPL-PMAA hydrogel cubes

Active

IPLVVPL

Doxorubicin

-

Targeting of hepsin

[41]

Magnetic NPs

Active

GSH

Paclitaxel

-

Brain delivery

[43]

GSH-PEGylated liposomes

Active

GSH

DAMGO peptide

-

Brain delivery

[45]

PAMAM-HAIYPRH

Active

HAIYPRH

Doxorubicin

-

Targeting of transferrin receptor

[47]

SPIONs-PEG-PEIFA-DOX

Active

FA

Doxorubicin

Drug delivery and imaging

[49]

IONPs-FA

Active

FA

–

Targeting and imaging of lymph node metastases of prostate cancer

[50]

Magnetic NPs-FA

Active

FA

–

Targeting of different cancer cells

[51]

FA-GSH-GNPs

Active

FA

–

Targeting and detecting of cancer cells

[52]

CNCs-FA

Active

FA

–

Targeting of different cancer cells

[53]

Fe3O4-FA NPs

Active

FA

–

Fe3O4-FA NPs as MRI contrast enhancement agent for detection of KB cells

[54]

QD-FA-GA-DOX

Active

FA

Doxorubicin

Targeting and imaging of cancer cells

[55]

Carboxymethyl chitosan NPs-FADox

Active

FA

Doxorubicin

Targeting of cancer cells

[56]

Fe3O4 NPs-FACurcumin

Active

FA

Curcumin

Targeting of cancer cells

[57]

Bi2S3@BSA-FACurcumin

Active

FA

Curcumin

-

[58]

Formulation

-

Targeting of cancer cells and tumor Chemoradiation therapy

References?

CHAPTER 12 Targeted Drug Delivery

207

TABLE 1

Summary of Active and Passive Targeting Studies—cont’d

Formulation

Targeting Strategy

Targeting Ligand/ Surface Coating

Drug

Purpose

References?

TDNs-AS1411

Active

AS1411

–

Targeting of cancer cells

[65]

AS1411-TMP

Active

AS1411

–

Targeting of cancer cells

[66]

AS1411-DNA pyramids

Active

AS1411

–

Targeting of cancer cells

[68]

THCPSi-SLNCs

Passive

–

–

Tumor targeting

[88]

Dual fluorescent HPMA copolymers

Passive

–

Doxorubicin

Tumor Targeting

[87]

Vorinostat-polymer NPs

Passive

–

Vorinostat

Tumor targeting

[86]

NO-NPs

Passive

–

Doxorubicin

Tumor targeting

[85]

HPMA-BA

Passive

–

BA

Targeting of cancer cells and tumor

[84]

Fe3O4@PPy-DoxPEG

Passive

PEG

Doxorubicin

Remote control of cancer treatment

[78]

HSA-TAC

Passive

–

TAC

Targeting of rheumatoid arthritis tissues

[77]

Cell Surface EGFR Receptors

EGFR Targeted NP

sive

Pas

1.0

Absorbance (A.U.)

Act ive

–4

0.8

Non-Targeted NPs Targeted NPs

0.6 0.4 0.2 –2

0 2 Log[DTX]

4

6

Non Targeted NP

FIG. 13 Schematic illustration of difference between active targeting and passive targeting. (Reprinted from

T.D. Clemons, et al., Distinction between active and passive targeting of nanoparticles dictate their overall therapeutic efficacy, Langmuir 34 (50) (2018) 15343–15349 with permission from American Chemical Society.)

the nanocapsules and the efficiency of passive and active targeting as well as immune evasion. They observed that soft nanocapsules were swallowed by macrophages three times lower than hard nanocapsules. It has also

been shown that FA-conjugated nanocapsules have high cellular uptake [95]. Hydrophilic quantum dots trapped in liposomes were developed for active delivery (by monocytes) and passive delivery (by EPR effect) to

FIG. 14 Potential benefit of active targeting over passive targeting. Schematic illustration of hybrid CT-FMT

imaging, in which the anatomical information obtained using CT is used to properly allocate (and quantify) the tumor accumulation of the fluorophore-labeled nanocarriers. (Reprinted from S. Kunjachan, et al., Passive versus active tumor targeting using RGD-and NGR-modified polymeric nanomedicines, Nano Lett. 14 (2) (2014) 972–981 with permission from American Chemical Society.)

FIG. 15 Schematic illustration of the passive targeting through the EPR effect, active targeting through the RGD grafting, magnetic targeting through a magnet of 1.1 T placed on the tumor (MT) and combination of active, and magnetic targeting comparison. (Reprinted from N. Schleich, et al., Comparison of active, passive and magnetic targeting to tumors of multifunctional paclitaxel/SPIO-loaded nanoparticles for tumor imaging and therapy, J. Control. Release 194 (2014) 82–91 with permission from Elsevier.)

CHAPTER 12 Targeted Drug Delivery

Targeted Cell

Cell Uptake

High

Short PEG

Long PEG

Targeted Cell

Low Gold Nanoparticles Integrin Receptor

209

Targeted Cell

Low

RGD

Proteins

FIG. 16 Schematic represents the interactions between nanoparticles with different surface chemistry and target cells. (Reprinted from G. Su, et al., Effects of protein corona on active and passive targeting of cyclic RGD peptide-functionalized PEGylation nanoparticles, Mol. Pharm. 15 (11) (2018) 5019–5030 with permission from American Chemical Society.)

breast tumors bearing mice [96]. Targeting, functionalization, and physiological responses of nanoparticles can be altered by protein corona. In this regard, Su et al. investigated the effect of corona protein on the efficiency and targeting functions of cyclic RGD peptidefunctionalized nanoparticles as well as their relationship with PEG length and density of targeting ligands bound to the nanoparticles. In this study, they synthesized 20 types of nanoparticles with different surface chemistry and observed that nanoparticles containing short PEG and medium RGD showed efficient targeting. It was shown that due to the presence of PEG on the surface of nanoparticles, protein corona could be useful for passive targeting by reducing macrophage cellular uptake (Fig. 16) [97].

5. CONCLUSION As discussed in this chapter, the design and development of targeted drug delivery systems have played a significant role in advancing the treatment of various diseases. In conventional drug delivery methods, any drug that is introduced into the body is distributed throughout the body and affects the normal areas in

addition to the diseased site, resulting in many side effects. Although in conventional therapies, a high dose of a drug is introduced into the body, but it is often insufficient for complete ablation of tumors, yet. The advent of medical nanotechnology and targeted delivery methods has revolutionized drug delivery. Advantages of targeted drug delivery over conventional drug delivery methods include (1) use the lowest dose of the drug; (2) specific delivery of the drug to the diseased site and not to normal areas of the body; (3) very low side effects compared to conventional drug delivery methods; (4) increase the therapeutic efficacy; (5) increase drug accumulation in the diseased site; and so on.

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CHAPTER 13

Strategies-Based Intrathecal Targeted Drug Delivery System for Effective Therapy, Modeling, and Controlled Release Action PRAVIN SHENDE • SHARAYU GOVARDHANE

Shobhaben Pratapbhai Patel School of Pharmacy and Technology Management, SVKM’S NMIMS, Mumbai, India

1. INTRODUCTION According to WHO, approximately one in six of the world’s population, suffer from central nervous system (CNS) disorders such as depression, seizures, locked-in syndrome, meningitis, and migraine. Despite multiple delivery routes for the targeted action to CNS, brain, and spinal are considered to be the most effective ways for emergency treatment. The presence of unique protective semipermeable layer composed of endothelial cell known as blood-brain barrier (BBB) acts as a challenging aspect for the delivery of actives to CNS. Intrathecal drug delivery has emerged as an alternative method for transportation to CNS, due to the direct administration to brain and spinal cord leading to the penetration through BBB and avoidance of side effect to other parts of the body. Intrathecal is the fluid-filled space present between the thin layers of tissue in the brain and spinal cord. Intrathecal drug therapy (ITDT) consists of a catheter connected with the drug reservoir for effective action from chronic intrathecal pain associated with osteoarthritis, fibromyalgia, shingles, and cancer [1]. ITDT aims to administer the drug directly to its receptor sites in a sufficient quantity for effective treatment and avoidance of adverse effects. The majority of patients for noncancer-related pain usually show spine pathologies as the cause of their pain. In cancer patient cases, intrathecal delivery of opioids are considered as an active mode for providing analgesic action with lesser side effects and higher quality of life. US FDA (United States Food and Drug Administration) approved the first battery-powered and programmable intrathecal pump for cancer-related pain in 1988 [2]. Table 1 shows the FDA-approved intrathecal preparations. The conventional dosage forms such as Rexulti, Brintellix,

Trintellix, and Cipralex Meltz undergo first-pass metabolism leading to low amount of actives available for the therapeutic effect. Although ITDT provides the advantage of direct delivery to the targeted site, the procedure to deliver the active therapeutics remains complex. Nanotechnology is considered as a foremost strategy to overcome the drawbacks associated with conventional dosage form. Recent trends in nanotechnology show higher lipid solubility, physical stability of nanoparticle, drug loading, and controlled rate of release for enhancing the permeation of drug into lipid layers of brain fluid. Recent trends in ITDT include novel drug delivery systems such as peptide-based drug delivery, magnetic drug delivery system, and hydrogel-based drug delivery system. In peptide-based drug delivery, the drug is chemically conjugated by a linker to the targeted moiety by a specific receptor and undergoes endocytosis to improve the delivery of drug. A novel approach of intrathecal magnetic drug targeting is developed in a combination of traditional intrathecal drug administration with magnetic drug targeting for highly localized treatment of neurological disorders such as Arnold-Chiari malformation and non-complex regional pain syndrome (CRPS) neuropathic pain. Thus nanotechnology in ITDT shows effective therapy in pain management for patients suffering from chronic intractable pain conditions. So the objective of this article is to provide the insights of new strategies for the delivery of drug by intrathecal route and the new trends in this field of medicine.

1.1. Outline of the Chapter This chapter is organized in four sections: Section 1 consists of the Introduction to ITDT and the difficulties

Modeling and Control of Drug Delivery Systems. https://doi.org/10.1016/B978-0-12-821185-4.00004-X © 2021 Elsevier Inc. All rights reserved.

213

214

Modeling and Control of Drug Delivery Systems

TABLE 1

Marketed Preparation of the Intrathecal Drug Delivery System.

S. No.

Marketed Preparation of the Intrathecal Infusion

1. 2.

Drug Content

Application?

Infumorph

Morphine sulfate

Chronic pain treatment

Prialt

Ziconotide

Pain reliever

associated with the conventional delivery. The Section 2 represents the strategies to overcome the barriers associated with ITDT, to ease the delivery and enhance the bioavailability for the actives. The Section 3 signifies the emerging trends in ITDT such as nanoparticulate drug carrier system, hydrogels-mediated drug delivery, microbubble-assisted ultrasound-based drug delivery, intranasal drug delivery, receptor-mediated opening, and carbon-nanotubes. The Section 4 consists of the newer development and the futuristic approach for the researchers in the field of ITDT

2. STRATEGIES FOR INTRATHECAL DRUG DELIVERY BBB is the main obstacle for the delivery of the drug across the brain. To overcome this barrier, few strategies allow the transfer of the actives through this wall (Fig. 1.) (See Fig. 2.)

2.1. Blood-Brain Barrier Disruption by Ultrasound BBB is a brain’s first-line of defense from harmful substances in the blood stream and composed of endothelial cells which are segmented extremely close to each other forming tight junctions. This tight joint of the tissue acts as main barrier for the penetration of drug and therapeutic effect in case of CNS disorders. The traditional method for drug delivery in the brain was the administration of small molecule drugs, transcranial drug delivery by an invasive catheter. Most small molecule drugs (400–500 Da) are not able to cross the BBB and few neurological conditions such as Parkinsonism and Alzheimer respond to small-molecule drugs [3]. Ultrasounds are mechanical waves with frequency greater than 20 kHz, frequency above the human hearing range. Ultrasound waves cause the disruption of BBB through widening the tight junctions and activating transcellular mechanisms, with little effect on the

surrounding parenchyma. Furthermore the opening occurs at acoustic power level orders of magnitude lower than was previous use, making method substantially easier to apply through the intact skull. Ultrasoundinduced effects are generated by two major mechanisms: thermal and nonthermal physical and biological effects. Thermal effects of ultrasound include increased blood flow, reduction in muscle spasm, increased ductility of collagen fibers, and a pro-inflammatory response [4]. The nonthermal effects of ultrasound, including cavitation and acoustic microstreaming, affects more in the penetration of actives against the tight conjugate. Recent development in image-guided focused ultrasound clinical system helped ultrasound to the targeted regions in the brain through the intact skull and the animal experiments. The administration of the small- and largemolecule drug to the brain can be delivered using FUS-induced-targeted BBB disruption [5].

2.2. BBB Disruption by Osmotic Mechanism Osmotic shock is obtained by sudden change in salt concentration leading to disruption of a number of cell types. Sudden osmotic shock to the endothelial cell causes the water withdrawal from the cell. This leads to the shrinkage of the endothelial cell, leading to stretching of the cell. As shown in the Fig. 3, the expansion of cell causes the net flow of water out of the brain cell, leading to the active transportation [6]. In 1972, Rapoport et al. reported CNS tissue staining with Evans blue consequently to intraarterial infusion of hypertonic arabinose. Since dye Evans blue, binds to albumin, is unexpected to permeate through the BBB, this observation suggested a BBB alteration by hypertonic arabinose [7]. Thus the combination of arabinose along with the Evans blue led to the alteration within the cell (hypertonicity). Thus the hypertonicity increases the diffusion of the actives through [8].

2.3. Overcoming Active Efflux at the BBB Xenobiotics are the foreign materials that are considered to be pharmacologically harmful for regular functioning. Active efflux transporter in BBB acts as a purification system inhibiting the entry of xenobiotic. Drug delivered to the brain acts as a xenobiotic inhibiting its entry into the brain. So, this active efflux plays a major barrier for the ITDT. Structural modification of the drug to reduce the efflux transporter affinity, co-administration of transport inhibitors, and many others are few of the approaches implemented for the delivery of drug across BBB. The CNS protective effect of BBB acts as a barrier for the treatment of brain malignancies or brain metastases, whereas the peripheral diseases are well controlled [9].

CHAPTER 13

Intrathecal Targeted Drug Delivery System

215

Strategies for intrathecal drug delivery

Ultrasound

Osmotic mechanism

Overcoming active efflux

Passive diffusion

FIG. 1 Strategies for intrathecal drug delivery system.

Ultrasound

Barrier for drug penetration

Ultrasound waves helps in the penetration of the drug through the porous formation

FIG. 2 Blood-brain barrier (BBB) disruption by ultrasound.

The CNS barrier can be partially overcome in the case of efflux transporter substrates by modulating the transporter proteins. ATP-binding cassette transporters are multispecific efflux transporters that bind to several of the efflux in the body. A slight modification in structure or the surface charge of this efflux helps in the penetration of the drug acting as xenobiotics for the BBB [10].

2.4. Passive Diffusion of Drugs Normal endothelial cell in BBB

Shrinkage of the endothelial cell in the presences of osmotic shock

FIG. 3 Blood-brain barrier (BBB) disruption by osmotic

mechanism.

Dispersion of drug across the membrane without energy source is known as passive diffusion. Most of the small molecular diffuse passively across the BBB and their flow is accelerated by partially association-dissociation between anions and cations for the formation of neutral

216

Modeling and Control of Drug Delivery Systems

ion-pair species in solution [11]. The passive diffusion of drug, with slight modification in the structure, molecular weight, lipophilicity, polar surface area, and others will help in the penetration of drug across BBB (Figs. 4 and 5). Polar surface area, charge, and hydrogen bonding: An alternative measure for BBB permeability is the permeability surface area products (PS) which is traditionally expressed as log PS although log P is a reliable indicator of permeability. In case of the liposomes, the change in the surface charge (cationic liposomes) helps in the binding of that nanocarriers toward specific receptor. The polar surface area is the sum of the polar atoms in the molecule. Thus the change in the sum of polar atom leads to surface modification and increases the affinity. Generally, molecules with a large polar surface area remain undiffused through the BBB, with the upper limit estimated between 60 and 90 Å [12].

FIG. 4 Passive diffusion of drug through blood-brain barrier

Passive diffusion of drugs

(BBB).

Polar surface area Hydrogen bonding Molecular charge Molecular flexibility

FIG. 5 Parameters for the passive diffusion of the drug

across blood-brain barrier (BBB).

3. EMERGING TRENDS IN INTRATHECAL DRUG DELIVERY 3.1. Nanoparticulate Drug Carrier System Nanocarriers are used in various drug delivery systems such as dendrimes, magnetic nanoparticles (NPs), lipid nanocarriers, and gold particles as shown in Fig. 6. Dendrimers (DDs) are hyperbranched molecule with nanosized-scale dimensions, composed of three layers of polymers: a central core, branched layer, and terminal functional group. The branched groups are attached to the central core and they usually describe the generation of the DDs. For example, the number of the layers of branched groups is three, it is indicated as a G3 DD and the terminal groups are responsible for the charges created over the surface. The surface groups may have positive, negative, or neutral charges, which act as an important parameter for the penetration across BBB. Poly(amidoamine) (PAMAM) is considered as one of the smallest and precious components for dendrimer synthesis. Because of nanosize and highly compact structure of PAMAM dendrimer, it is considered as ideal form for brain delivery (Table 2). Because of open structure of DDs, small molecules get easily encapsulated and covalently attached to it. DDs are mainly nanosized, and they provide better stability making them ideal for the delivery across BBB [27]. Table 2 indicates the improvement in the bioavailability and stability with the use of DDs as nanocarrier. Lipid NPs: Lipid-based nanocarriers are composed of physiological lipids, making them well-tolerated, nontoxic, and are degraded to a nontoxic residue. Lipid NPs are divided into two groups: liposomes and other lipid NPs, such as solid lipid nanoparticles (SLNs) and nanostructured lipid carriers (NLCs). Liposomes are bilayer lipid particles mainly composed of phospholipid and cholesterol and size ranging from 100 to 1000 nm. Liposomes are considered as safe and selective therapeutic tool for the delivery of the drug across BBB. The advantages of liposomes for intrathecal drug delivery are as follows: (1) the lipid composition of the liposomes helps in the easy penetration across BBB. (2) Intrathecal delivery of the liposomes allows diffuse distribution. (3) Encapsulation of drug in liposomes changes the pharmacokinetic of the free drug providing control release of the drug. (4) Drug entrapped in the liposomes will avoid the side-effect associated with continuous infusion. The surface-charged liposomes also act as ideal delivery of drugs and genetic material, as they provide facilitate interaction with the cell membrane and improves uptake [28].

217

Lip

osom

es

Gold nanoparti cles

Dendrimers

Lipid rti nanopa cles

Intrathecal Targeted Drug Delivery System

c eri ym ti Pol opar nan cles

M nan agnet opa ic les rtic

CHAPTER 13

FIG. 6 Nanocarrier drug carrier system for intrathecal drug delivery.

TABLE 2

Nanocarrier Drug Carrier System for Intrathecal Drug Delivery. S. No. Nanocarrier

Formulation

Material and Methods

Observation

1.

In vivo delivery of siRNA to the brain by carbosilane dendrimer

Carbosilane, cysteamine hydrochloride, 2,20 dimethoxy-2-phenylacetophenone, fluorescein isothiocyanate isomer I, iodomethane using thiolene reactions

Dendrimers helped to [13] transfect siRNA to HIVinfected human primary astrocytes and achieved gene silencing without causing cytotoxicity

Efficient gene delivery targeted to the brain using a transferrinconjugated poly(ethyleneglycol)modified poly(amidoamine) dendrimer

Plasmid pEGFP-N2, poly(amidoamine) (PAMAM), human holo-transferrin, Bolton and Hunter reagent, and 2-iminothiolane hydrochloride, 5,5-Dithiobis, NHSPEG-MAL

Transfection efficiency [14] of PAMAM-PEG-Tf/ DNA dendimers complex was much higher than PAMAM/ DNA and PAMAMPEG/DNA complexes in Brain capillary endothelial cells

Doxorubicin-PAMAM dendrimer complex attached to liposomes: cytotoxic

Hexadecylphosphocholine, Egg yolk phosphatidylcholine, PAMAM, (N-tris(hydroxymethyl)methyl-2aminoethanesulfonic acid,

Encapsulation [15] efficiency was found to be 97% and doxorubicin–PAMAM

Dendrimers

References?

Continued

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Modeling and Control of Drug Delivery Systems

TABLE 2

Nanocarrier Drug Carrier System for Intrathecal Drug Delivery—cont’d S. No. Nanocarrier

2.

3.

Lipid nanocarrier

Formulation

Material and Methods

Observation

References?

studies against human cancer cell lines

stearylamine, sulphorodamine B, trichoroacetic acid. Thin film hydration method

complex 3:1 M ratio was found to be most stable. Liposomal formulations composed of lipids and of a drug– dendrimer complex act better for release

Brain delivery of Baclofen as a hydrophilic drug by nanolipid carriers: two characteristics and pharmacokinetics evaluation

Baclofen, glyceryl monostearate, glyceryl distearate, glyceryl trioleate, Tween 80. Double emulsion-solvent evaporation technique

Particle size and zeta [16] potential of baclofen NLC was found to be 127  5 nm and  14 nV, respectively. Encapsulation efficiency was found to be 38.5%–42%.

Intrathecal Dimyristoyl-sn-glycero-3administration of phosphocholine, neostigmine, liposomal chloroform neostigmine prolongs analgesia in mice

For the neostigime[17] liposomal formulation, the median analgesic duration was 1.0, 1.5, and 6.0 h for the 20, 40, and 80 mg doses, respectively

Bupivacaine, Cetyl Palmitate, Optimized NLC: a Dhaykol 6040, Capryol 90, Precirol nanotechnological approach to improve ATO5, Pluronic F68 (P68) the anesthetic effect of bupivacaine

In vivo analgesic effect [18] elicited by NLCBupivacaine was found to be twice that of free bupivacaine

Methoxypolyethylene glycol amine, N-hydroxysulfosuccinimide, 1-ethyl-3-carbodiimide. Multipleparticle tracking method

Particle size of PNPs [19] was found to be 85 nm and larger nanoparticles is expected to allow more uniform, longerlasting, and effective delivery of drugs within the brain

Foxp3 plasmidencapsulated PLGA nanoparticles attenuate pain behavior in rats with spinal nerve ligation

PLGA, dichloromethane, poly(vinyl alcohol), plasmid

Foxp3 NPs attenuated [20] pain behavior induced by spinal nerve ligation in rats for 7 days by suppressing microglial activity

ChABC-Loaded PLGA nanoparticles: a comprehensive study on biocompatibility,

PLGA, poly(vinyl alcohol), bovine serum albumin, ChABC enzyme, Luxol fast blue, cresyl violet, Span 60, Tween 80, dichloromethane (DCM)

[21] The scarglial degradation in animals studies by the ChABCloaded particles was

Polymeric A dense poly(ethylene nanoparticles glycol) coating (PNPs) improves penetration of large PNPs within brain tissue

CHAPTER 13

Intrathecal Targeted Drug Delivery System

219

TABLE 2

Nanocarrier Drug Carrier System for Intrathecal Drug Delivery—cont’d S. No. Nanocarrier

Formulation

Material and Methods

functional recovery, and axonal regeneration in animal model of spinal cord injury 4.

Magnetic Magnetic field nanoparticles distribution modulation of intrathecal delivered ketorolac iron-oxide nanoparticle conjugates produce excellent analgesia for chronic inflammatory pain

5.

Gold L-DOPA nanoparticles functionalized, multibranched gold nanoparticles as brain-targeted nanovehicles To deliver gold nanoparticles to the brain by conjugation with a peptide that recognizes the transferrin receptor

References?

proved using immunohistochemistry

Ketorolac, iron oxide, carbodiimide, tetramethylazanium hydroxide

Superparamagnetic Rhodamine, maghemite, poly(ethylene glycol) liposomes for MRI monitoring and external magnetic field-induced selective targeting of malignant brain tumors Magnetic brain tumor targeting and biodistribution of long-circulating PEGmodified, crosslinked starch-coated iron oxide nanoparticles

Observation

Starch-coated, magnetite, N-hydroxysuccinimidyl, methoxyl poly(ethylene glycol), dimethylsulfoxide, sodium phosphate, sodium phosphate

Gold chloride, methyldopa, mPEG thiol, sodium citrate, 4-ethylcatechol. Seed-mediated method

Ninhydrin, 2-mercaptoethanol, gold (III) chloride hydrate, sodium tribasic dehydrate, 5(6)-carboxyfluorescein, collagen type IV, fibronectin. Citrate reduction of HAuCl4

Intrathecal ketorolac- [22] conjugated iron oxide particle administration demonstrated a magnetic fielddependent analgesia effect in CFA pain model with a significant reduction in COX expression The [23] magnetoliposomes are longer then liposomes, as they retained over a 24-h period

Selective magnetic [24] brain tumor targeting (t¼ 1 h) of PEG-MNPs was confirmed in 9 Lglioma tumors, with up to 1.0% injected dose/g tissue nanoparticle delivery The amount of

[25]

L-DOPA-AuNFs

transported across the BBB (1.22% ID/h) is higher in vitro BBB models (0.1%–1% ID/h) [26] Peptide sequence interacts with the transferrin receptor in the microvascular endothelial cells of blood–brain barrier, causing an increase in the permeability of the conjugate in brain

220

Modeling and Control of Drug Delivery Systems

SLNs are nanocarriers composed of phospholipid, dispersed in water or in surfactant solution (50–1000 nm). The absence of the hydrophilic domain in SLNs enables it to carry lipophilic compounds but allows nanoparticle to cross the BBB with ease [29]. NLCs are modified SLNs composed of solid and liquid lipid. The combination of lipid in NLCs helps to improve drug loading capability, biocompatible, and stability [30]. Polymeric nanoparticles (PNPs): PNPs are colloidal particles comprising of active ingredient entrapped or adsorbed on solid macromolecule and the size varying from 1 to 1000 nm. PNPs in case of the intrathecal drug delivery provide advantage of greater stability, high encapsulation efficiency, controlled release kinetics, and charge modification. Systemic delivery for the CNS includes polymer such as poly(butyl cyanoacrylate), poly(lactic acid), poly(glycolic acid) (PLGA), chitosan, and poly(ethylenimines). These polymers help in the entrapment of drug and peptide molecules. This polymers biocompatible, biodegradable, helps in the better entrapment of drug, helps to cross BBB and prolong the drug lifetime in the body. Local delivery for the CNS includes polymers such as PLGA, poly(methylidene malonate), poly(epsilon-caprolactone), and chitosan [31]. Camptothecin-loaded PLGA NPs were delivered locally and demonstrated to be effective for treatment of an intracranial tumor model [32]. Magnetic NPs (MNPs) are nanomolecules consisting of iron-oxide as a metal core often coated with the organic fatty acid, polysaccharides, and polymers. The

metal core has an unpaired electron that contributes to its magnetic property and the polymers help in the stabilization and prevent the separation into particles and carrier medium. Iron as core material provides low toxicity, easy elimination and is used in oxide form due to its stability. MNPs are considered as one of the impending system for the brain delivery, due to easy permeation through BBB and low toxicity. Fig. 7 shows the method for delivery of the drug to the affected site of the spinal cord. The unpaired electron charge of the iron gets attracted toward the magnet and the movement of the magnet helps in the delivery of the drug [33]. Colloidal gold nanoparticles (AuNP) appear to be popular carrier for drug, due to their inert and nonimmunogenic properties, better bioavailability, and ease of preparation. AuNP is composed of gold which act as metallic core and the polymers are coated covalently over this particle with an organic layer. The covalent bond between gold and polymer facilitated to improve biocompatibility, biophysical properties, and targeting specially in case of proteins, peptide, and gene therapy [34]. The tunable, nanoconjugate, easy penetration through BBB and low toxicity properties of the gold with polymer, AuNPs can be considered as ideal drug delivery for the intrathecal targeted drug delivery, as shown in Table 2.

3.2. Hydrogels-Mediated Drug Delivery Hydrogels are 3D nanostructured cross-linked networks polymer consisting of large amount of water molecule. Highly superior property such as native extracellular

Spinal cord

Drug delivery

Magnet Intrathecal delivery

FIG. 7 Magnetic nanocarriers for intrathecal drug delivery system.

CHAPTER 13 matrix (ECM) allows them easy encapsulation of the hydrophilic drugs and excellent bioavailability. Due to their high tuneable mechanical property (0.5 kPa to 5 MPa) hydrogels are used as excellent carrier for proteins and enzymes to avoid their degradation. Highly matrix structure of hydrogel allows control and sustained release for both drug and nucleotide-based drugs such as plasmid DNA [35] (Table 3). In case of the spinal damages, nanosized and microsized particles injected for the control release of the actives. Unlike hydrogel, many of the nanoparticles are unable to fill the helix-loop-helix transcription. The ideal system for the spinal injuries should provide local and sustained release, be biodegradable, be single doe administration and noninflammatory in the CNS. Due to in situ gelation, hydrogels can be exactly injected to fill up the spinal damages by exact geometrical reshaping to avoid more surgeries. As shown in the Fig. 8, in situ hydrogel are thermolabile, that is, change in the temperature leads to gelation leading to control release of the drug. Thus hydrogels are promising carriers for controlled drug release, single dose administration and biocompatibility for clinical application [39].

3.3. Microbubble-Assisted Ultrasound-Based Drug Delivery Microbubbles (MBs) is a noninvasive method used for the increasing the permeability through BBB. MBs are

Intrathecal Targeted Drug Delivery System

221

used in the analysis and medical treatment of diseases, due to its tunable and transient effect on vasculature. Because of their high compressibility and tendency to cavitate, microbubbles can transform the kinetic energy from the traveling acoustic wave to the local microenvironment. This acoustic wave helps in the penetration across the BBB creating a gas bubble. While cavitating, microbubble increases the fluid streaming within the diameter’s range away their surface. By the period, the fluid stream reach the surface of the tissue, this creates the bubble. The bubble helps in the penetration of the actives through the cell, and thus increasing the bioavailability of the drug. As the main component of ultrasound contrast agent, MBs with diameter less than 10 μm can pass through pulmonary circulation and enhance the contrast of ultrasound imaging in diagnosis [40]. MBs therapy used along with IV method helps in widening the interendothelial clefts and tight-junctions of the BBB. This method is also known as sonoporation. The method relies on the mechanical action of the gas MBs in ultrasonic pressure waves. The MBs are about 1–10 μm in diameter containing a lipid or protein shell loaded with heavy gasses, which can be excreted by exhalation and make MBs more stable. In MBs, molecules are attached to the layer of the BBB and through the caustic wave’s drug penetrate to the stem cell and also viral vectors [41].

TABLE 3

Hydrogel-Mediated Intrathecal Drug Delivery System. S. No.

Formulation

Material and Methods

Observations

References?

1.

Intrathecal delivery of a polymeric nanocomposite hydrogel after spinal cord injury

Sodium hyaluronate, methylcellulose, poly(lactic glycolic acid), poly(vinyl alcohol). Water-oil-water double emulsion method

Composite hydrogel was well tolerated in intrathecal space of spinal cord injured rats and no significant effect on locomotor function up to 28 days

[36]

2.

Click cross-linked injectable hyaluronic acid hydrogel is safe and biocompatible in the intrathecal space for ultimate use in regenerative strategies of the injured spinal cord

Furan-modified hyaluronic acid, 2-(N-Morpholino)ethanesulfonic acid, 4-(4, 6-dimethoxy-1, 3, 5-triazin-2-yl)-4methylmorpholiniumchloride Water/oil/water (W/O/W) double emulsion procedure

Encapsulation efficiency of hyaluronic acid hydrogel was found to be 47.2% and a loading of 34.6 ng BDNF/ mg nanoparticles was obtained. Delayed release upto 76  9% was obtained

[37]

3.

Hydrogel-assisted antisense LNA gapmer delivery for in situ gene silencing in spinal cord injury

20 O-methyl RNA-DNA AON gapmer, acetonitrile, bisdiphenylacetyl disulfide, ammonium hydroxide

75% downregulation was obtained within 5 days after hydrogel-assisted antisense LNA gapmer injection

[38]

222

Modeling and Control of Drug Delivery Systems

Hydrogel formulation

Release of the drug from gel during change in temperature

FIG. 8 Mechanism action of hydrogel by intrathecal route.

3.4. Intranasal Drug Delivery Intranasal drug delivery is the administration of the drug through nose by insufflated method. Brain-targeted intranasal delivery helps to transport drug across the olfactory epithelium to brain. The distal areas of the CNS, olfactory bulb or in the brainstem are the main part for the distribution of the drug through intranasal administration. Fig. 9 shows the direct delivery of the formulation to the brain with the help of the nasal route. In this administration, extracellular route is the main pathway for the delivery. In extracellular route the drug is present between the supporting cells, where the drug is passing through paracellular cleft, the lamina propria, perineural space, and reaches at the subarachnoid space. Many of the recent studies have proved the direct uptake of drug by brain and cerebrospinal fluid by adjusting the molecular weight and lipophilicity of the formulation. Table 4 shows the increase in the bioavailability and permeability of lipidformulation given by the nasal route. A new approach in intranasal delivery includes use of in situ gelation method. In in situ gelation method, the sol gets converted to the gel formation helps in the adhesion to the mucous and leads to control release of the drug [45].

3.5. Receptor-Mediated Opening Receptor-mediated opening is the process of getting nutrients and materials into the cell. A specific receptor

FIG. 9 Intranasal drug delivery.

is present on the cell surface, which bind tightly to the extracellular macromolecule, thus forming a pathway for the entry of the outside materials. Adenosine is purine nucleoside involved in a myriad host functions and adenosine receptor (AR) substrates modulate BBB

CHAPTER 13

Intrathecal Targeted Drug Delivery System

223

TABLE 4

Intranasal Drug Delivery System. S. No.

Formulation

Material and Methods

Observation

References?

1.

Intranasal drug delivery of frovatriptan succinate–loaded PNPs for brain targeting

Frovatriptan, poly(vinyl alcohol), poly(lactic-coglycolic acid) Double emulsion (w/o/w) method

Frovatriptan-loaded nanoparticle size was found to be 264.4  0.04 nm, zeta potential 35.17  0.07 mV, and 65.2  0.06% entrapment efficiency was obtained. Sustained release up to 72 h was achieved

[42]

2.

An enhanced charge-driven intranasal delivery of nicardipine attenuates brain injury after intracerebral hemorrhage

Chitosan, tripolyphosphate, poly(lactic acid), poly(methyl methacrylate), nicardipine, ionic crosslinking

Nicardipine particle size was found to be 439.6  11.9 nm with a PDI of 0.307, and the zeta potential was +21.05  0.48 mV

[43]

3.

To deliver ziconotide to cerebrospinal fluid by intranasal pathway for the treatment of chronic pain

Ziconotide acetate, chitosan, Krebs-ringer bicarbonate, acetone, hydrochloric acid

The elimination rate was found to be constant of ziconotide in cerebrospinal fluid intranasal and intravenous administration of ziconotide solution was found to be 0.54  0.08 h1 and 0.42  0.10 h1, respectively

[44]

permeability. AR modulation leads signaling to brain endothelial cells causing opening of the BBB and facilitate permeability of the drug across the cells into the CNS. AR is used for improved drug delivery into the brain by activating A2A receptors or blocking the entry of neurotoxic agents or inflammatory immune cells into the brain [46].

3.6. Carbon Nanotubes Carbon nanotubes (CNTs) are novel cylindrical nanomaterial consisting of hexagonal arrangement of hybridized carbon with attractive physical, chemical, and electronic properties made from graphene. CNTs have the unique property of translocating across plasma membrane, which makes it novel alternative for the delivery of the drug across BBB. The nanomeshworks of single-walled CNTs proved to show support to the neuron growth. Kafa et al. studied the penetration ability of amino-functionalized multiwalled carbon nanotubes (f-MWNT) to cross the BBB in vitro using a co-culture BBB model. Co-culture model proved the maximum transportation of 13.0  1.1% after 72 h. After intravenous injection in mice, f-MWNT exhibited substantial brain uptake (1.1  0.3% injected dose/g) at 5 min [47].

4. CONCLUSION Intrathecal drug delivery system has significant approach to treat patients with the chronic pain. The new strategies for the ITDT show the effect of the delivery to increase the bioavailability and permeability through the BBB. BBB is considered to be the main barrier for the delivery of actives through the brain cell. The new strategies help in the disruption of the BBB indirectly helping in the penetration of the drug through the thick cell wall. The upcoming nanotechnology has proved its effect for the delivery of the brain through noninvasive method. The nanotechnology helps to overcome the barrier through techniques such as lipid solubility, polymer coating, small molecular weight polymer, and use of magnetic drug delivery system increasing the bioavailability of the drug. The nasal and in situ hydrogel drug delivery has proved its effect for the brain delivery. Computational fluid dynamics (CFD) is the novel forthcoming approach used to study the pharmacokinetics for ITDT. CFD is paradynamic studies used for equation solving of fluid motion to produce the data of fluid flow phenomena. Some of the recent advantages in the intrathecal pump such as Prometra pump, Synchromed II pump, and Flowonix pump proved its advantage over the traditional pump

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Modeling and Control of Drug Delivery Systems

by reducing the pain and ease of delivery associated with ITDT. However, as modernization in intrathecal drug delivery limits the area of research due to cost of therapy, requirement of trained personal for administration of medicine and pain associated with route. Although nanotechnology has helped to overcome such barrier, the safety and efficacy regarding this technology will be resolved in the near future.

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CHAPTER 14

Biopolymer-Based Hydrogel Wound Dressing

MONA ALIBOLANDIa,b • ELNAZ BAGHERIa,b • MARZIEH MOHAMMADIc • ELHAM SAMEIYANb • MOHAMMAD RAMEZANIa,b a

Pharmaceutical Research Center, Pharmaceutical Technology Institute, Mashhad University of Medical Sciences, Mashhad, Iran, bDepartment of Pharmaceutical Biotechnology, School of Pharmacy, Mashhad University of Medical Sciences, Mashhad, Iran, cDepartment of Pharmaceutics, School of Pharmacy, Mashhad University of Medical Sciences, Mashhad, Iran

1. INTRODUCTION Chronic ulcer is a critical healthcare problem worldwide. A retrospective study in 2018 demonstrated that 8.2 million people experienced wounds whose treatment cost ranging from 28.1 to 96.8 billion USD [1]. In such wounds, the healing process is impaired due to several diseases and conditions such as peripheral vascular insufficiency, age, poor nutrition, diabetes, and local pressure effects which disrupt organized cellular and molecular processes required for wound healing [2]. Nonhealing wounds impose a significant cost on the patient and the medical system [3]. Their prevalence rate is reported to be 2% of the population and it was demonstrated that 6.5 million people were affected by chronic wounds in 2009 only in the United States [4]. Injuries because of a disease or accidental/intentional trauma lead to disruption of the integrity of the tissue (mucosa, skin, or any organ) activating a cascade of cellular pathways to restore the tissue integrity. In general the healing process involves hemostasis, inflammation, proliferation, and tissue remodeling. At the time of the damage, the arterial vessels are constricted to prevent exsanguination leading to hypoxia and acidosis and consequent release of nitric oxide and vasodilation. The enhanced release of histamine improves vascular permeability and infiltration of inflammatory cells into the wound environment. In the inflammatory phase, chemotaxis and the entrance of neutrophils, macrophages, and lymphocytes remove bacteria and also the damaged host tissue called debridement. Besides, they stimulate angiogenesis and release of growth factors regulating the inflammatory response. As long as the bacteria and the debris exist, the inflammatory phase persists that leads to tissue damage, delayed proliferation leading to nonhealing wound. After achieving the

hemostasis and the inflammatory balance, proliferation phase begins through angiogenesis, collagen deposition, and re-epithelialization [5]. Keratinocytes are the major cells which migrate from the wound edge to the wound to restore the epidermis. Conventional wound dressings were made of cotton gauze which acts as a physical barrier to pathogens. However, their capacity to absorb the wound exudate is low. Besides, such dressings adhere to the damaged area and hurt the fragile neotissue [6]. Among different wound dressings, hydrogel-based dressings are of great interest. Hydrogel is a 3D cross-linked network of hydrophilic polymer chains which can absorb high amounts of water between the chains [7]. In addition to high water content, flexibility, good mechanical stability, and the availability of gaseous exchange, hydrogels can incorporate bioactive materials, regenerative cells, and drugs thus making them ideal wound dressings. This feature is due to the similarity between the hydrogels and extracellular matrix (ECM) that both are made of polymeric networks in aqueous environment [8]. Xiao et al. developed collagen-chitosan hydrogels loaded with QHREDGS peptide to protect keratinocytes against oxidative stress in wounds in diabetic mouse model to enhance re-epithelialization [2]. Hydrogels are produced from the biopolymers, the synthetic polymers, or the composites. Biopolymers are polymers made of natural sources. Natural polymers are made of animals [gelatin, chitosan, silk, collagen, and hyaluronic acid (HA)], plants (starch and cellulose), or algae (alginate). Their main advantage is biocompatibility and biodegradability [9] mimicking the ECM of the injured tissue. Liu et al. fabricated hydrogels composed of chondroitin 6-sulfate (CS) and heparin which were used to deliver basic fibroblast growth factor (bFGF) to improve healing process of full thickness

Modeling and Control of Drug Delivery Systems. https://doi.org/10.1016/B978-0-12-821185-4.00019-1 © 2021 Elsevier Inc. All rights reserved.

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wounds in diabetic mice [10]. They proved that the clinical potential of proteoglycans as the hydrogel and the growth factor in accelerating the healing process. In this chapter, various biopolymeric hydrogels used in wound dressings are categorized, and recent studies describing the characteristics and their exceptional features are discussed. Additionally the ideal properties of wound dressings to accelerate the healing process and to diminish the scar formation are explained. Moreover, their clinical application and FDA-approved products are briefly described.

2. WOUND DRESSING AND ITS IDEAL PROPERTIES In the past, strips soaked in oil or grease, honey, oil, and wine were used as treatment agent for wound healing. Besides, wool boiled in water or wine was applied as wound bandage. Wound dressing is placed directly on the wound and then fixed in place using bandage. During the 19th century, the antiseptic technique was discovered to manage the infection [11]. Wound dressing appeared in 20th century [12], and they were developed into more sophisticated tools in the late 20th century. Wound dressings were fabricated with increasing reepithelialization, angiogenesis, and collagen synthesis properties by providing acidic pH and hypoxia to reduce the chance of wound infection [13]. During the mid1990s, synthetic wound dressings with the features of providing moisture and absorbing wound exudates were considered [14]. As an ideal wound dressing, it should have properties such as preserving moist environment, improving epidermal migration, allowing gas exchange between environment and wounded tissue, increasing angiogenesis and connective tissue synthesis, defending against bacterial infection, improving leucocytes migration, having long shelf life, being free of toxic contaminants, and protecting wound from further trauma [15, 16]

3. WOUND DRESSING BASED ON BIOPOLYMERS Polyphenols, peptides, polyesters, polysaccharides, and other natural polymers with repeating units/monomers originate from the living organisms such as fungi (chitin), animal (chitosan, HA, and collagen), bacteria [bacterial cellulose (BC) and exopolysaccharides], plant (natural rubber, starch, and cellulose), and algae (alginate). These polymers with biodegradability, renewability, biocompatibility, and lower antigenicity properties are preferred over the synthetic materials [17–20]. In addition, they play a key role in wound-healing process

with antiinflammatory, targeted actions to specific cells, antibacterial, and proliferative features. Biopolymers have been blended with other polymers to promote biomimetic and mechanical strength features for the development as skin substituent [21]. Considering the inherent physicochemical and biological properties of ECM, proteins, and polysaccharides possess high level of biomimicry. In the construction of hydrogel, collagen, gelatin, and HA can be considered as ECM support, while RGD and LDV sequences in silk fibroin and keratin are used as cell-recognition domains and biomolecule-binding sites. In addition, chitosan and alginate have antibacterial and antiinflammatory properties [22].

3.1. Dextran Dextran is a type of polysaccharide composed of anhydroglucose rings [23]. Dextran with alpha-1,6 linked D-glucopyanose residues, possesses biodegradability, biocompatibility, nontoxicity, and hydrophilicity properties [24–26]. D-Glucopyranose residues can be easily oxidized by NaIO4 to generate aldehyde groups providing then functional group to be conjugated with amine side groups of gelatin, chitosan, and other polymers to form various hydrogels as antibacterial materials, smart sustained drug delivery vehicles, and hemostatic agents. In addition, dextran-based hydrogel is used as a wound dressing due to its ability to stimulate wound-healing process [27–29]. Excessive inflammation accompanied by the immediate damage of blood flow at the injury site of burn inducing pain which are the major differences between incisional and burn wounds. Owing to the ineffectiveness of using single therapeutic agent in promoting burn wound healing, a combination therapy is usually used for the treatment. In a study, vascular endothelial growth factor (VEGF) was applied as an angiogenesis agent which is working by accumulation of inflammatory cells in the damage site and thus stimulating the proliferation and migration of the endothelial cell [30]. Although naked DNA of VEGF is unstable and displayed low transfection efficiency, gene therapy approach using polyethylenimine (PEI) combined with pDNA encoding VEGF provided versatile gene transfection ability. Moreover, resveratrol with immunomodulatory, antioxidant, and chemopreventive properties is a natural polyphenolic compound found in grape skin which could be implemented as a wound-healing accelerator [31]. Resveratrol can upregulate the expression of VEGF in human skin cells [32]. However, it has poor water solubility. To overcome this problem, a hydrogel scaffold consisted of dextran (Dex), HA, and β-cyclodextrin (β-CD) was produced

CHAPTER 14 under ultraviolet (UV) irradiation through reductive chemical reaction. In the aforementioned system, HA exhibited cellular proliferation induction and migration roles while dextran improved skin regeneration and facilitated neovascularization [33, 34]. At the final stage, PEI/ pDNA-VEGF complex and resveratrol were co-loaded in the prepared hydrogel. The obtained results verified the excellent characteristics of the prepared platform as a wound dressing [35]. Gelatin and dextran having biodegradability and biocompatibility properties were used as self-healing materials. Chlorhexidine acetate (CHA) acted as an antimicrobial agent. Beside, bFGF can promote angiogenesis, cell proliferation, migration, and tissue repair [36]. One of the barriers to using bFGF is its short half-life. In this regard, a sequential drug delivery system of CHA and bFGF was used to overcome this obstacle [37]. Hydrogel is a suitable dressing material with sequential release properties for wound healing. The hydrogel network was prepared by the dynamic reaction under physiological conditions (pH 7.4) implementing oxidized dextran (OD), aminated gelatin, and adipic acid dihydrazide. Encapsulation of bFGF in PLGA (poly(lactic-coglycolic acid)) microspheres can control the release of bFGF from the hydrogels. After the application of bFGF@PLGA/CHA/hydrogel into the wound site, sequential delivery of bFGF and CHA and thus acceleration of wound-healing process occurred [38]. In another attempt, Du et al. applied a hydrogel dressing with self-healing, antibacterial, and hemostasis activity. Chitosan and OD were used to produce the composite hydrogel. In this regard, positively charged chitosan interacted with negatively charged bacterial membrane leading to the leakage of cellular proteins and other cellular components. Besides, hydrophobically modified chitosan (hmCS) has hemostatic and antibacterial activity. Aldehyde groups of OD interacted with the hmCS to obtain hydrogels with great potential as a wound dressing [25]. Zheng et al. designed a hydrogel based on poly(vinyl alcohol) and dextran-aldehyde which were cross-linked using freeze-thaw and freeze-drying method. The porous structure of hydrogel was uniform with 5–10-μm pore size. The prepared hydrogel accelerated the time frame of skin regeneration. Silver nanoparticles (AgNPs) are effective antimicrobials agents. On AgNPs aggregation, the release of silver ions would decrease, thereby reducing the antimicrobial activity of AgNPs. To overcome this problem, the antifouling hybrid hydrogel consisted of cationic-thiolated chitosan and anionic maleic acid-grafted dextran was used to incorporate AgNPs and provide release of silver ions.

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As a result, the expression of CD68+ and CD3+ increased leading to improved wound-healing process. The prepared AgNPs-loaded hydrogel introduces an excellent candidate for the treatment of diabetic ulcers [36]. In another study, poly(vinyl alcohol) (PVA)/dextran/ chitosan hydrogel was applied for the preparation of a novel wound dressing. In this hydrogel, dextran reacted with PVA using glutaraldehyde (GA) as the cross-linker to improve angiogenic responses and regeneration of skin during burn wound-healing process. In this regard, chitosan, dextran, and PVA have also positive impact on antimicrobial, angiogenesis, and cell proliferation properties of this hydrogel [39]. Li and coworkers applied the block copolymer, DA95B5, consisted of dextran-block-poly((3–4 acrylamidopropyl) trimethylammonium chloride-co-butyl methacrylate), to remove biofilms of various multidrugresistant Gram-positive bacteria. DA95B5 was selfassembled to form core-shell nanoparticles with cationic core and dextran shell. Dextran shell increased solubility of the bacteria-nanoparticle complex and thus decreasing the biofilm formation. The removal of biofilm occurred by weakening the bacterial attachment. This hydrogel was used to combat against multidrug-resistant Gram-positive bacteria producing biofilms [40]. In another study, a hydrogel composed of dextranpoly(ethylene glycol) (PEG) was prepared. In the structure of this hydrogel, a thiol group of each antibiotic including polymyxin B and vancomycin (Vanco) was conjugated to the hydrogel network to provide a versatile antibacterial wound dressing against Gram-positive and Gram-negative bacteria [41].

3.2. Collagen Collagen is a bioactive polymer extracted from skin, tendons, bones, ligaments animal cells, and forms 30% of the total protein in the body [42]. Due to its well-known structure, biocompatibility, hemostasis ability, bioresorbability, and reduced manufacturing cost properties, collagen is abundantly used as drug release support [43–49]. However, resistance to in vivo enzymatic degradation and its mechanical properties limited the application of collagen. To overcome the aforementioned obstacles, various strategies for its conjugation with other biopolymers or various cross-linkers were used [50, 51]. In this regard, matrices, hydrogels, fibers, and membranes are among various forms of collagen [52–56]. The porous form (spongious matrices) of collagen is capable of absorbing large amounts of wounds exudate and preserving a moist environment for wound healing. As a result, spongious matrices of collagen loaded with antiinflammatory drug, can

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be used for the treatment of tissue injury and burn damage [57]. Flufenamic acid as an NSAID is a therapeutic which is used in soft tissue injuries and different rheumatic disorders. In a study, spongious matrices of dextran and collagen which were cross-linked using lyophilization method were applied for the flufenamic acid delivery. Drug release from this hydrogel can be controlled by the degree of sponge swelling and drug diffusion through the swollen polymeric matrix. This hydrogel accelerated the regeneration of the affected epithelial tissues thereby improving the wound-healing process [58]. In another attempt, three natural sources of raw materials consisting of carboxymethyl chitosan/collagen peptide (COP)/oxidized konjac composite (KGM) were cross-linked to form a new hydrogel. To prepare this hydrogel, in the first step, KGM was oxidized by NaIO4 producing OKGM. Then OKGM was acted as cross-linker agent and facilitated the formation of uniform porous composite structure of CMCS/COP/ OKCM. By increasing the OKGM, gelation time also increased while its pore diameter and water evaporation rate decreased and swelling performance improved. In addition, in vitro study indicated that this hydrogel stimulated the proliferation of NS-FB cells. Besides, this composite showed potential application as wound dressing [59]. Mousavi et al. compared the properties of chitosan/ gelatin and chitosan/collagen hydrogels. Blending of Col and CS resulted in homogenous and porous structure of hydrogel while mechanical strength and hydrophilicity properties were improved. A significant reduction in degradation and swelling rate for the CS/ Col was observed in comparison with CS/Gel hydrogel. Besides, in vitro study indicated that the CS/Col hydrogel had better cell survival and adhesion to their surfaces in comparison with CS/Gel hydrogel [60]. In another study the microbial transglutaminase was implemented as an effective catalyst for the conjugation of amino group of chitosan to the COP molecules. Hydrogel was prepared by coupling of amino groups of the CS-COP and active aldehyde of OKGM via the Schiff-base reaction. Water retention capacity, equilibrium swelling, and blood compatibility properties were all improved by using this hydrogel. In addition the results indicated that the prepared hydrogel accelerated the wound-healing process by absorption of the wound exudate and providing moist wound environment [61].

3.3. Chitosan Chitosan is derived from chitin by alkaline deacetylation process and most commonly found in cell walls of fungi and the shells of some arthropods [62].

Chitosan as a linear and semicrystalline natural polymer [63] has important properties including biodegradability [64, 65], biocompatibility [66, 67], mucoadhesion [68], antiinfection, low toxicity, antimicrobial, and other bio-functional properties [69] which can help to speed up the wound-healing process [70–74] and facilitate homeostatic activity [75, 76]. Moreover, the positive charge of chitosan [77], its reactive amino groups [78], solubility in acidic medium [79], and low cost [80] has made it a good candidate for drug delivery and wound-healing purposes [81]. Chitosan-based hydrogels are usually used for the treatment of skin lesions in diabetes. There are various studies indicating that chitosan hydrogels have been effective in improving of these lesions. For example, in a study using chitosan hydrogel containing flavonoids extracted from Passiflora edulis leaves, antioxidant, and wound-healing properties were observed in rat skin lesions [82]. Flavonoids have been shown to have high biological potentials including antioxidant and antiinflammatory properties, activating enzymes and interfering with some signaling pathways [83]. In this study, results of the in vitro tests performed to evaluate the biocompatibility and release of flavonoids from chitosan hydrogels, showed rapid and continuous release of flavonoids. Furthermore, in in vivo studies, histopathological analysis and macroscopic evaluation of male Wistar diabetic rats confirmed the wound-healing properties of the prepared formulation. In another study performed on a wound model of diabetic mouse, fluorinated methacrylamide chitosan (MACF)-based hydrogel was used to enhance oxygenation to the wound site [84] which can improve wound healing [85, 86]. The results showed that wound dressing with MACF-oxygenating hydrogel sheet increased collagen synthesis and neovascularization augmentation, which was shown by higher collagen content and greater number of blood vessels and capillaries, respectively. In some cases, other polymers cross-linked to chitosan improved its elasticity and strength properties [87]. PVA as a synthetic polymer has remarkable properties including strong water retention ability, noncarcinogenicity, biocompatibility, and biodegradability features [88]. Therefore in combination with chitosan, it can improve the mechanical and physicochemical properties of hydrogel. To this end, a hydrogel composite made of chitosan lactate and PVA was developed by Mahato et al., which can be an effective candidate for wound dressing [89]. They prepared composite drugloaded hydrogels by blending of PVA and chitosan lactate followed by GA cross-linking. The in vitro results

CHAPTER 14 showed that fabricated hydrogels are compatible with cells and facilitate cells adhesion. In addition, sustained release of ciprofloxacin from the designed drugcontaining hydrogel prevented the growth of Escherichia coli, thereby assisting antimicrobial activity under physiological conditions. Xu and colleagues have developed a PVA/chitosan hydrogel composite that could accelerate wound healing and tissue regeneration by maintaining moisture and sustained release of the stromal cell-derived factor 1, as a chemokine and significantly recruit bone marrow mesenchymal stem cells both in vitro and in vivo [90]. In addition the hydrogel was found to be greatly biocompatible with skin tissue and no side effect or irritation was observed. Although PVA/chitosan composite hydrogels have high biocompatibility and good antibacterial ability, their poor mechanical strength limits their application in wound dressing. Moreover, PVA/chitosan hydrogels, as an environmental conditioner cannot ideally meet wound dressing requirements to accelerate wound healing. Therefore novel hydrogels are made up of this composite such as lignin/PVA/chitosan which can more effectively accelerate wound healing [91]. It has been shown that adding lignin to the PVA/chitosan hydrogel could greatly enhance the mechanical strength, wound environmental regulation ability, and protein adsorption capacity of the hydrogel. Thus application of this composite hydrogel significantly accelerated wound healing in murine wound model. AgNPs are extensively applied in biomedical research due to their unique properties such as effective antibacterial, antifungal, antiviral, antioxidant, and antiinflammatory activities. These features can be significantly useful in treating many types of infections in diabetic and chronic and nonhealing wounds [92]. Therefore mixing hydrogel structures with AgNPs could be an effective approach for wound healing. In this regard, Jiang and colleagues designed a chitosan hydrogel that cross-linked with konjac glucomannan as a biocompatible polysaccharide with great therapeutic potential [93]. On the other hand, the AgNPs were incorporated into the hydrogel to improve its antimicrobial activity. Because of its swelling ability, this nanocomposite hydrogel could absorb wound exudates and exhibit self-healing properties making the structure stable. Furthermore the hydrogel, as a carrier is able to modulate the release of silver ions, consequently reducing AgNPs cytotoxicity. Rat models with infected skin defects were used to evaluate the wound healing. As a result, the AgNPs hydrogel dressing was able to stimulate wound healing and reduce inflammatory response and show valid clinical application potentials.

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In another study, the development of a chitosan-PEG hydrogel impregnated with AgNP was reported to enhance wound-healing process [94]. This formulation could provide a sustained and slow release of AgNP for the treatment of chronic diabetic wounds. The results indicated a higher degree of swelling, higher water vapor transition rate and higher porosity for AgNP-impregnated hydrogel in comparison with chitosan-PEG hydrogel alone. This hydrogel exhibited antioxidant and antimicrobial properties in vitro and improved wound-healing ability in vivo in diabetes-induced rabbits. The hydrogel displayed a sustained release of AgNPs for at least 7 days demonstrating the slow biodegradation of designed hydrogel. Xie et al. have developed a chitosan hydrogel network which AgNPs were incorporated into it with a view to reinforce the mechanical features and improve the antibacterial properties of chitosan hydrogel [95]. In this work, they also evaluated the swelling characteristics of the chitosan hydrogel network and the efficacy of the wound healing on Sprague-Dawley rats. The results displayed that the novel hydrogel owing to intermolecular and intramolecular interactions, has a porous three-dimensional network with extremely high mechanical properties. In addition, hydrogel showed higher antibacterial activity compared with the control group and significantly increased the rate of epithelial remodeling and collagen deposition, which accelerated the woundhealing process. In recent years, various types of hydrogel-based biological materials have been developed to respond to environmental stimuli such as temperature, pH, biomolecules, or a combination of these factors [96]. The pH of the cutaneous wounds has been shown to be dynamic and associated with the wound-healing process [97–99]. Through taking advantage of the pH difference between healthy skin (4.0–6.3) and chronic wound (7.15–8.93), one can develop stage-specific wound treatments to respond to these environmental symptoms using pH-sensitive hydrogels. Meanwhile, due to the cationic nature of chitosan, it can be pH-responsive and able to indirectly regulate tissue enzymatic activity that is raised at the alkaline pH of chronic wounds [100]. In this context, Zhu et al. designed a pH-sensitive chitosan hydrogel that was functionalized with methacrylic anhydride to achieve physically and chemically adjustable properties [101]. The pH-sensitive methacrylated chitosan (MAC) hydrogels exhibited tunable mechanical properties, pH sensitivities and swelling ratios without affecting its degradation behavior in vitro cell responses.

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The results show that the amino groups of MAC hydrogel (pKa 6.5) [102] become protonated at low pH and their interaction with water molecules increased. Finally, these hydrogels presented good ability to undergo volume changes during pH range transition in wound-healing environment which may indicate their suitability as wound dressing. Another pH-responsive chitosan hydrogel is the hydrogel in which genipin is used as a cross-linker [103]. The chitosan-genipin hydrogel is capable of absorbing large volumes of fluids at different pH ranges for chronic wound dressing applications. Characterization studies showed that the constructed chitosan-genipin hydrogel was able to neutralize an environmental pH, whereas showing an average of 230% absorption of aqueous solution, suggesting its use as a potent wound dressing. Studies on bacterial activity demonstrate the ability of hydrogel to inhibit E. coli growth by 70%, whereas retaining its biocompatibility toward keratinocyte and fibroblast cells in vitro. In addition, chitosan-genipin hydrogel increased the cell proliferation and immune response in induced pressure wounds in mice. Temperature is another important environmental factor on which smart hydrogels can be designed. The hydroxybutyl-chitosan-dopamine (HCS-DOPA) composite hydrogel is an example of these temperatureresponsive gels [104]. The ability of this composite as a temperature-responsive hydrogel at 4°C and 37°C was investigated by in vitro phase transformation method. The HCS-DOPA composite hydrogel produced below 5.0% hemolysis rate and had no cytotoxicity in mouse fibroblast cells. The in vitro antibacterial studies of this composite revealed >8 h activity against Staphylococcus aureus growth. The in vitro whole blood test confirmed that blood clotting time treated with HCS-DOPA-2 composite hydrogel was reduced to 95.6 seconds. In vitro evaluation of homeostasis indicated that HCS-DOPA-2 composite hydrogel could be used as a favorable wound dressing. In another study, a PVA/chitosan/honey/clay nanocomposite hydrogel was used as a wound dressing which was responsive to both pH and temperature [105]. In this nanocomposite, nanoparticles of clays help to regenerate and form the gels, while honey plays the role of a plasticizer (softener) and diminishes the possibility of PVA crystals formation under a freezethawing cycle through disrupting the order of the PVA chains. Swelling studies were carried out at 20°C and 37°C at different pH. The results displayed that swelling was enhanced as a result of temperature increase and maximum swelling happened at a pH of 2. Furthermore, faster honey release rate occurred at higher pH values.

The ability of the proposed system to smart release of drug against pH and temperature changes confirms that this system can release the drug according to the need of treatment of skin injury. MTT assay results demonstrated no cytotoxicity in nanocomposite hydrogel system. The designed nanocomposite indicated more than 99% antibacterial activity and in vivo studies also confirmed the ability of the introduced system in effective wound healing.

3.4. Cellulose Cellulose with natural fibril structures and low antigenicity is a renewable resource derived from plant cell walls [106, 107]. Cellulose fibers organized in a complex three-dimensional (3D) with great flexibility and high tensile strength are produced by bacteria (e.g., Acetobacter and Agrobacterium). BC-based wound dressings accelerate wound healing by providing an adequate moist and confirmed liquid and gas permeability. The 3D network of cellulose chains interacted with secondary active compounds, for example, antimicrobial and bactericidal agents. However, relatively short drying-out time of BC is a considerable drawback. As a result, a second layer was used for covering BC-based dressings. By control of microfibrillar organization, material porosity and chemical modifications of BC surface, the BC water holding capacity (WHC) could be concisely tuned. Various biopolymers such as alginate [108, 109] and chitosan [110] promoted water absorption capacity of BC. Table 1 represented the cellulose-based wound dressing hydrogels. In this regard, Sulaeva et al. applied BC hydrogel as wound dressing platform. In this study, to improve WHC and wound sticking behavior of BC hydrogel, alginate, as a secondary hydrophilic component, was used. In addition, poly(hexamethylene biguanide) hydrochloride as cationic antimicrobial agent was impregnated into the composite hydrogel for the treatment of highly contaminated wounds [117]. Tobramycin (TB) is an aminoglycoside antibiotic which has a broad activity against a variety of Grampositive and Gram-negative organisms involved in skin infection [118–120]. Carbonyl group of cellulose can react with TB to form imine bond without using a cross-linking agent [121, 122]. As a result, imine bonds in the structure of the hydrogel provided a slow rate of hydrogel degradation under mild acidic condition. However, the delivery efficiency of TB cannot be easily controlled because of low water solubility and its easy sublimation. Cyclodextrin (CD) was applied to overcome the aforementioned problem. CD by capturing volatile hydrophobic molecules acted as a molecular container, thereby increasing the solubility of

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Biopolymer-Based Hydrogel Wound Dressing

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TABLE 1

Cellulose-based composite hydrogels as wound dressing. Composite of hydrogel

Incorporated materials

Method of fabrication

Function

2+

References

2,2,6,6Tetramethylpiperidine-1-oxyl oxidized bacterial cellulose (TOBC)

Calcium alginate

Zn

Bacterial nanocellulose (BNC)grafted poly(acrylic acid) (AA)-graphene oxide (BNC/P (AA)/GO)

Graphene oxide

Electronbeam irradiation

Improve biocompatibility of hydrogel

[112]

Poly(vinyl alcohol)/ carboxymethyl cellulose/ polyethylene glycol (PVA/CMC/PEG)

PVA and PEG

Freezethaw method

Inhibit bacterial penetration and control the moisture loss at the wound site

[113]

ZnO/PVA/carboxymethyl cellulose

ZnO and poly(vinyl alcohol)

Freezethaw method

Promote water vapor transmission rate (WVTR), proper degree of swelling ratio (DSR) and good ability to heal and protect wounds

[114]

Hyaluronic acid (HA) and hydroxyethyl cellulose (HEC)

HA

–

Improved the absorbent capacity and maximum swelling rate

[115]

Bacterial nanocellulose/acrylic acid (BNC/AA)

AA

Electron beam irradiation

Both as a cell carrier and accelerator of wound healing

[116]

Carboxymethyl chitosan (CMC) and rigid rod-like dialdehydemodified cellulose nanocrystal (DACNC)

CMC

–

Pain killer and stop scar forming

[112]

hydrophobic drug molecule and its bioavailability. In this study, dialdehyde carboxymethyl cellulose (DCMC) is applied as the main base of a pH-responsive hydrogel. The aminoglycoside antibiotic drug, tobramycin as a cross-linking agent, reacted with DCMC through imine bonds to form a hydrogel. Borneol/mono-6-(2-hydroxy-3-(trimethylammonio)propyl)-β-cyclodextrin (BN/EPTAC-β-CD) by electrostatic force was reacted with carboxyl groups of DCMC and dispersed inside the hydrogel. Continuous release of the moisture and drugs occurred to promote wound healing [123]. In another attempt, carboxymethyl chitosan and oxidized CMC were coupled to each other through Schiffbase reaction to form a hydrogel/microgel (Gel/MG) composite. Mechanical performance and stable properties of microgel were improved after addition of microgel to the hydrogel. Besides, drug release occurred in both acid and alkali environment. In this hydrogel, silver sulfadiazine acted as an antibacterial agent. The

[111]

release of the drug at pH 9.5 was more than that of pH 7.4 and pH 5.5. Gel/MGs hydrogel was used as drug delivery and wound dressing platform [124]. In another study, after incorporation of bioactive molecules and chemical modifications of bacterial nanocellulose (BNC), it was used as suitable candidate for wound dressing applications. Vancomycin and ciprofloxacin, as antimicrobial agents, were incorporated into BNC. Besides, the modifications of BNC was performed by grafting of glycidylmethacrylate and crosslinking trough stable CdC bonds with ethylene glycol dimethacrylate that produced cellulose carbon-centered radicals scavenged by methacrylate structures. BCN-CD was applied as drug career for wound dressing and tissue-engineering applications [125]. Curcumin (CUR), as natural antibiotic, is used as an antibiotic in wound management and regenerative medicine [126]. Though, low hydrophobicity limited its applications for wound management. In a study, cellulose produced by Gluconacetobacter xylinus (ATCC

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23770) provided a 3D network structures with unique properties. For loading of hydrophobic CUR in ATCC 23770, CUR was encapsulated in CD to decrease its hydrophobicity. As a result, water soluble CUR:hydroxypropyl-β-cyclodextrin was loaded into ATCC 23770 hydrogel to facilitate wound healing [127]. Henna, as an agent against fungal infection with hepatoprotective, antidiabetic, and antiinflammatory properties extracted from the Lawsonia inermis has been used for wound-healing purposes [128, 129]. Hydroxyethyl cellulose (HEC) is one of the water soluble derivative of cellulose, considered as hydrogel-like material that has both solid- and liquid-like properties [130]. Because of its physiochemical stability, the shape and the structural integrity were maintained after the elastic strength. Poly(N-vinylpyrrolidone) (PVP) as main component of temporary wound dressings with good hydrophilicity is a synthetic polymer. In this regard, Raafat et al. applied HEC as a natural polymer blended with PVP through γ-irradiation method to improve gel strength, and moisture retention providing a versatile wound dressing. They loaded henna into the HEC/ PVP hydrogel and observed good antibacterial and skin regeneration properties [131]. Graphene oxide (GO) possesses antibacterial activity. Besides, sodium CMC (NaCMC) is one of the cellulose derivative with good biodegradability, high drug loading capacity, and high water content properties [132, 133]. In a study, after reduction of GO using sodium hydroxide (NaOH), reduced GO (rGO) was incorporated into NaCMC polymer forming a composite hydrogel. This hydrogel indicated antibiofilm activity [134]. Bacterial strains of the genra Acetobacter produce BC for hydrogels fabrication, as the potential scaffolds for tissue engineering and wound-healing applications [135, 136]. Mohamad and coworkers, co-loaded human dermal fibroblasts and human epidermal keratinocytes into the BC/acrylic acid (AA) wound dressing hydrogel. The prepared hydrogel facilitated re-epithelialization and accelerated burn wound-healing process. Cell transport and wound protection were the important applications of the BC/AA hydrogel [137]. For the treatment of infected chronic wounds, broadspectrum antibiotics should be applied, due to the presence of the multiresistant bacteria strains. Pinho et al. developed a composite wound dressings by joining CD/cellulose-based hydrogels and cotton textile substrate. It was demonstrated that the CD/cellulose enhanced the applicability of textile substrates as wound dressing. Antiinflammatory agent, garlic acid

and phenolic acid as the antimicrobial agent, were co-loaded into the composite hydrogel. CD improved the release of phenolic acid and enhanced the antimicrobial and antiinflammatory properties of the garlic acid. This composite offered an optimum approach for wound dressing [138]. In another study, a composite hydrogel based on cellulose nanocrystals, gelatin, and HA was constructed through amide and hydrogen bond formations followed by freeze-drying. The prepared hydrogel possessed spongy properties, with the pore diameter about 80–120 μm. Gelatin- and HA-simulated ECM and could improve penetration of cells into the hydrogel. This threedimensional porous composite hydrogel stimulate the attachment, growth, and proliferation of fibroblasts and accelerated skin tissue regeneration [139]. The combination of CMC and human hair keratin with citric acid as a cross-linker was fabricated as a composite hydrogel. The prepared composite hydrogel as an ECM-mimetic material promoted cellular homing and their proliferation. Clindamycin, an efficient antibiotic, was also loaded into the composite hydrogel for healing soft tissue and serious skin infections at the burn wound sites [140]. In this study, keratin had impact on sustained release of clindamycin, water uptake and water vapor transmission rate. The in vitro release study showed that the release rate of clindamycin decreased when CMC was attached to keratin and the growth of S. aureus colonies was inhibited after 24 h [141]. Silver ions possessed a growth-inhibitory capacity against a broad spectrum of microorganisms. Usually, AgNPs are aggregated and loss their antibacterial property. In a study, silver ions were incorporated into the polymer matrix of hydrogel based on PVP, polyethylene glycol (PEG), agar, and CMC. CMC was used in this hydrogel featuring high swelling capacity, high water solubility, and low cost. In vivo study indicated that silver ions-loaded hydrogel improved wound healing by neovascularization and proliferation of fibroblasts in the damaged tissue [142]. Chuah and coworkers developed poly(acrylic acid) and BC nanofibers to manufacture hydrogel hybrid biomaterial composites used for wound dressing. BC component provided a microporous sponge-like structure and enhanced the hydrogel strength. Amoxicillin (AM) as antibiotic was loaded into this hydrogel to fight against wound infection. This hydrogel was pH sensitive since at an acidic pH of medium, the swelling of the composite decreased while the swelling ratio increased at an alkaline pH medium. Absorption of more exudates occurred from the chronic wounds and the release of AM was improved at higher pH. Hence, this hybrid

CHAPTER 14 hydrogel could be used as a wound dressing especially for chronic wounds [143].

3.5. Alginic Acid Alginate, as a linear acidic polysaccharide consisting of manoronic and gluronic acids in adjustable ratios has favorable properties including low cost and high biocompatibility. This biopolymer has been able to make an important contribution to the industry and medicine in a variety of applications, for example, scaffold and wound dressing [144, 145]. Specifically, this natural polymer is capable of forming stable aqueous hydrogels through physically cross-linking with divalent cations such as calcium or barium salts [146–148]. Hydrogen sulfide (H2S) is known to be an important gasotransmitter which can accelerate wound healing by improving angiogenesis. Zhao and colleagues designed an alginate hydrogel which played the role of a sponge loaded with H2S [149]. This sponge consisted of a functional sodium alginate (SA) incorporating JK-1 molecule (SA/JK-1) which acts as a pH-dependent H2S donor. It was able to release H2S constantly under acidic pH condition and supply a protective and moist healing environment by absorbing exudate at the wound interface. SA/JK-1 displayed good cyto-compatibility and further improved fibroblast proliferation and migration in vitro. Moreover, evaluations of in vivo full thickness dermal defect models demonstrated that SA/JK-1 can considerably accelerate wound-healing process with improved granulation tissue formation, collagen deposition, re-epithelialization, and angiogenesis. Exosomes (EXOs) are known as membrane lipid vesicles secretory products in most cells with diameters of 30–150 nm [150], which play an important role in intercellular communications [151]. Stem cells-derived EXOs can be very effective in tissue repair and wound healing because of their significant benefits such as high stability, homing effect, nonimmune rejection, and easy control of its concentration [152, 153]. In addition, EXOs are able to enhance the process of proliferation, migration, and angiogenesis in the wound area by modulating the secretory activity of dermal fibroblasts and enhancing the collagen/elastin synthesis and secretion, which eventually re-epithelialization [152, 154]. In a study, Shafei et al. applied EXOs derived from adipose-derived stem cells (ADSCs) supernatant and loaded them onto the alginate-based hydrogel which acts as a bioactive scaffold to maintain the EXOs in the wound area [155]. This bioactive wound dressing displayed good biocompatibility and biodegradability as well as significant potential for wound closure, vessel formation, and collagen synthesis at the wound site.

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There are various studies on the positive effect of platelet concentrate on soft tissue regeneration in terms of wound healing, pain reduction, and consequently improvement of patient’s quality of life [156, 157] with reduced hospital sleep-over and associated costs. A freeze-dried sponge consisting of a combination of alginate (Alg), silk sericin (SS) with platelet lysate (PL) has been proposed to produce a bioactive wound dressing for skin lesions care [158]. In this study, the highest level of growth factor release was reported within 48 h as the optimal time to undergo the healing process in vivo. It was also observed that when PL was incorporated into the biomembrane, its ability to protect cells against oxidative stress and proliferation induction was enhanced. Furthermore, results of in vivo studies in a mouse skin lesion model demonstrated that the biomembranes with PL accelerated the healing process, resulting in faster burst of inflammation, new collagen deposition and formation of granulation tissue thereby causing a rapid skin regeneration. The emergence of antibiotic-resistant pathogens has created serious problems in the treatment of infected burn wounds. Nowadays, phage therapy offers promising treatment options for combating against antibioticresistant pathogens [159–163]. Hence, Kaur et al. introduced a wound-dressing system based on PVA-SA hydrogel for topical delivery of antibiotics and bacteriophages with the aim to treat the infected burn injuries [164]. In this system, hydrogel membrane creates wound-healing environment while the adsorbed antibiotic/bacteriophages protect the wound against the infection. Different concentrations of SA and PVA were blended together to achieve an ideal ratio for wound dressing. The final blend displayed great protein adsorption, gel fraction, swelling index, hemocompatibility, and ideal mechanical properties. Moreover, they obtained self-adherent, antibacterial activity and biocompatible membrane properties through the in vitro antibacterial and cell cytotoxicity assays. The in vivo potential was assessed by a murine burn wound model of MRSA-infected, which exhibited notable bacterial reduction, diminished inflammation and wound contraction in the membrane-treated groups compared with the control group. The double-coated hydrogel membrane providing both MR10 phage and minocycline has been used as a better therapeutic approach for treating resistant burn wound infection than antibiotic and phage alone. In another system, alginate hydrogels containing AgNPs were used to determine the effect of honey on nosocomial wound infections of multiresistant clinical strains of Acinetobacter baumannii and Pseudomonas

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aeruginosa [165]. In this system, optimized colloidal solution consisting of 50% of honey and 8 nm AgNPs was applied to create nanocomposite Ag/alginate hydrogels. These nanocomposites were produced in three different forms including microbeads, microfibers and discs which retained all AgNPs and a large fraction of honey components. Antibacterial activity assays revealed that at 9 μg/mL concentration of total released silver, the hydrogel demonstrated stronger antibacterial activity against standard and most of the examined multiresistant hospital strains. However, this study recommended the need for further in-depth studies on the bacterial resistance mechanisms and the potential of the novel Ag/alginate hydrogels with honey combination to accelerate healing and combat wound infection as an antibacterial, nonsticky, and bioactive dressing. CUR has significant properties in wound healing, including antiinflammatory, antioxidant, antimicrobial and antifungal activities which has led to its introduction as a model drug [166–168]. Acevedo et al. used an alginate hydrogel and polycaprolactone nanoparticles loaded with CUR for potential applications in wound healing [169]. This alginate membrane revealed various functional features required to act as a substitute for synthetic skin, such as high swelling capacity and adhesion to the skin, the ability to have pores to regulate the loss of transepidermal water, drug-controlled release by the incorporation of nanoparticles and transparency for monitoring the wound. The composition of the nanoparticles supported the drug permeation into different layers of the skin and thus can solve the problems of CUR solubility. The clinical application of this system protected the wide areas of mixed first- and seconddegree wounds without the need for any removal, thereby reducing patient discomfort and the risk of new epithelial malformations. Another type of alginate hydrogel was made from the sodium alginate solution by Kazi et al. which was used as a surgical sealant material for wound healing [170]. They made a surgical incision with bleeding on mouse dorsal surface and then used alginate solution at the wound site. It was then observed that the alginate sample successfully sealed the bleeding wound after 1 and 2 weeks and also promoted tissue regeneration without the help of other surgical/dressing tools. This suture-free wound closure may be of great value for those wounds on which sutures are difficult to be placed or when there are concerns about the aesthetic appearance. The ability of alginate to form a gel in a mild environment has made it a suitable choice for injectable wound dressing materials. By taking advantage of this alginate potential, an injectable nanocomposite hydrogel

comprising nanosized calcium fluoride particles was produced through the in situ precipitation process [171]. In this study, it was found that the amount of fluorine ions released from nanocomposite hydrogels increased with increasing CaF2 content inside the composite hydrogel, and these ions also in turn induced the proliferation and migration of fibroblast cells in vitro. The antibacterial activity of the composite hydrogel against S. aureus and E. coli was carried out by using colony formation test, in which the number of bacterial colonies was significantly reduced. The results of in vivo studies based on a full-thickness wound model exhibited that the nanocomposite hydrogel efficiently increased the ECM deposition and thereby accelerating the wound-healing process. Since hydrogels are able to maintain a moist environment, they can be a suitable candidate for burn wounds, as a wound dressing. In particular, wound dressing based on the combination of gelatin and alginate biopolymers showed great potential [172]. At first, both polymers were modified by introducing photo-crosslinkable functionalities then forming methacrylated gelatin and methacrylated alginate (gel-MA/alg-MA). The gel-MA films were incubated in alg-MA solutions and cross-linked subsequently into double networks. In vitro experiments indicated an improved cell adhesion for lower alginate films and superior mechanical features. Moreover, good biocompatibility with compatible cell attachment characteristics was demonstrated for hydrogel dressings. In general, a good ratio of gelatin-alginate can allow the use of materials as wound dressing for some days without tissue ingrowth. Impaired wound healing is one of the important challenges of diabetes that results in incomplete healing of chronic diabetic wounds. Persistent and elevated reactive oxygen species (ROS) have been identified in vivo and have been found to be related to impaired wound treatment in chronic nonhealing wounds [173]. Thus ROS can cause inflammation in the diabetic chronic wounds. Therefore scavenging the ROS in wound may be an impressive treatment for chronic ulcers. Fan et al. used edaravone, as a potent effective radical scavenger and loaded it into nanocomposite hydrogels composed of alginate and positively charged Eudragit nanoparticles [174]. On the one hand, alginate hydrogels augmented protection and sustained release of edaravone, while Eudragit nanoparticles increased the solubility of edaravone. In this study, it was demonstrated that nanocomposite hydrogel was able to improve wound healing in a dose-dependent manner. In vivo studies also showed that low doses of edaravone-loaded nanocomposite hydrogel could

CHAPTER 14 improve wound healing in diabetic mice, whereas high doses of edaravone might postpone the treatment. These results suggest that ROS plays a dual role in the treatment of chronic wounds. These evaluations also revealed that dose factor could be an important limiting key in the translational application of antioxidants in wound healing. Summa and colleagues designed another system based on sodium alginate (NaAlg) that, in combination with povidone iodine (PVPI), could improve wound healing [175]. In this system, alginate, with its wound-healing properties and PVPI with antibacterial and antifungal properties, were able to provide a controlled antiseptic release. They also showed that NaAlg/PVPI films could efficiently reduce the inflammatory response in both rodents and in human foreskin fibroblasts after lipopolysaccharide stimulus wound induction. In addition, animals treated with NaAlg/ PVPI films showed remarkably higher wound closure than untreated animals at all-time points. It is noteworthy to mention that complete wound closure was attained in 12 days in film-treated group only, demonstrating that full-thickness ulcers healed more rapidly in these animals. Taken together, the results confirmed the efficacy of NaAlg/PVPI films in accelerating wound healing. In another study using composite scaffolds consisting of chitosan and sodium alginate, different concentrations of sodium alginate were used to enhance the physicomechanical properties [176]. The system also used lentil seed extract (LSE) loaded on the composite bioscaffolds because of its antioxidant/antibacterial activities. The results showed that with increasing concentration of sodium alginate, the mechanical properties of the scaffold improved dramatically, confirming that sodium alginate enhanced the flexibility of the scaffold. Furthermore, it was shown that LSE-loaded scaffolds had higher antibacterial activity and greater free radical scavenging ability than blank scaffolds, resulting in faster wound healing. In conclusion, these studies show that the use of alginate hydrogels with the ability to combine with various materials can be used for effective wound healing. Therefore, because of the high potential of this biopolymer to create an ideal environment for accelerating wound healing, it can be used as a bioactive wound dressing.

3.6. Starch Starch, as a natural biodegradable and hydrophilic polysaccharide which can be modified chemically or physically, is a more effective wound-healing material when

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compared with synthetic polymers [17]. Due to its hydrophilic nature and poor mechanical properties, starch cannot form a stable hydrogel. Thus starch widely combined with other polymers to prepare versatile hydrogel structure [177]. Hassan et al. blended PVA and starch to synthesize a hydrogel membrane. PVA is a suitable polymer for hydrogel preparation due to its compatibility, nontoxicity, solubility, and biodegradability features [178]. Biopolymers when compared with synthetic polymers are more effective wound-healing materials [17]. In a study, cross-linking of PVA with starch formed a hydrogel membrane. In addition, turmeric, as an antibacterial agent, was added to the hydrogel membrane to create stronger hydrogen bond interactions. The prepared hydrogel demonstrated antibacterial activity and wound-healing capability [179]. In another study, PVA, starch (St), and chitosan (CS) were blended to prepare a composite hydrogel membrane. At neutral pH, chitosan lost its positive charge and do not have antibacterial activity. Due to the aforementioned reason, zinc oxide (ZnO) that possesses antibacterial activity was added to the prepared composite hydrogel membrane. In vitro and in vivo studies confirmed wound-healing potency for this hydrogel-based wound dressing [180]. Transparency is a pivotal property for a wound dressing due to the easy screening and observation of wound site without removing the hydrogel. In this regard, starch by heterogeneous dispersion of nanoparticles could increase light transparency [181].

3.7. Gelatin Gelatin is a mixture of several peptides and proteins which can be extracted from the alkaline and acid hydrolysis of collagen [182]. Due to the amine and carboxylic groups of gelatin, it can be reacted with other polysaccharides to provide cellular adhesives composite for tissue regeneration [183]. RGD motifs in the gelatin structure led to the cell adhesion property of the gelatin-based scaffolds [184]. Epichlorohydrin and GA are used as cross-linker to overcome low stability of gelatin [185–187]. Due to the high toxicity of epichlorohydrin and GA, their biological application is limited and thus, nontoxic and green cross-linkers are recommended [188]. Table 2 represents the gelatin-based composite hydrogels with wound dressing potency. Attack of pathogen toxins or pore-forming toxins (PFTs) the cellular membrane, the cell permeability was altered in term of bioactivity [193, 194]. Numerous approaches such as small-molecule inhibitors, antisera, and monoclonal

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TABLE 2

Gelatin-based composite hydrogels as wound dressing. Incorporated materials

Method of fabrication

Materials function

References

Sodium 2-acrylamido-2-methylpropanesulfonate (Na-AMPS)/gelatin

Na-AMPS

UV photoinitiation

Wound healing without any harm

[189]

A poly(vinyl alcohol) (PVA)/gelatin hydrogel

PVA

Esterification followed by solution casting

Inhibitory effect against two main skin pathogens, the Gram-positive Staphylococcus aureus and the Gram-negative Pseudomonas aeruginosa

[190]

Gelatin, and glycosaminoglycans (HA and chondroitin sulfate) incorporated with asiatic acid (a triterpenoid) and nanoparticles (zinc oxide and copper oxide)

ZnO, CuO, GAG, asiatic acid

–

Applied as a wound healing in second degree burns

[191]

Gelatin paste containing collagen type I/ chondroitin 6-sulfate (coll/chondroitin) sponge and gelatin

Chondroitin

In this hydrogel, collagen bundles resembling normal skin and wide formation of blood vessels around the wound of dermis

[192]

Composite hydrogel

antibodies [184] and recently, functional nanoparticle incorporated into hydrogel were applied for the removal of PFT by injecting the hydrogel into the target site and subsequent detoxification by interaction with diffused toxins. In a study, nanoparticle-functionalized macroporous hydrogel was fabricated as a supporting matrix in the macroporous detoxification program. Polydiacetylene (PDA) nanoparticles were encapsulated in gelatin cryogel and formed a hydrogel to improve the therapeutic efficacy by capturing a broad spectrum of PFTs. PDA nanoparticles are fabricated through a self-assembly method and then blended with the gelatin monomer solution for polymerization. The absorption of the PFTs and a long retention time were the features of this hydrogel for local cure against bacterial infections [195]. Due to both spray-filming and self-healing abilities of hydrogels, they are ideal barrier against wound infection. For preparation of spray-filming form of hydrogel, the gelation time must be reduced. Borax is unsuitable for reduction of gelation time because of its carcinogenicity [196–198]. By forming Schiff base bond between monoaldehyde-modified sodium alginate (SA-mCHO) and adipic acid dihydrazide-modified gelatin (ADHGel), the spray-filming hydrogel was prepared with selfhealing properties. The gelation time was reduced to 2–21 s. The growth of S. aureus and Candida albicans were suppressed after 12 h exposure to this hydrogel [199].

Aloe vera (AV) extract, as a therapeutic agent, possesses suitable properties on cell proliferation and wound-healing process. In addition, to obtain a sustained release of therapeutic agents, it must be incorporated into particle-based carrier. In a study, a sustained released niosomal formulation of AV was incorporated into the AG hydrogel. AV-loaded niosomes incorporated into the AG hybrid hydrogel indicated an extended sustained release and could be proposed as a promising candidate for wound healing [200]. In another study, poly(sodium 2-acrylamido-2-methylpropane sulfonate) (NaAMPS) and gelatin (GE) were used to prepare antibacterial hydrogel. In addition, chlorhexidine gluconate (CHG) is an effective antiinflammatory and antiseptic drug against a wide range of Gram-negative and Grampositive bacteria [201]. CHG was loaded into Na-AMPS/GE hydrogel which was fabricated through UV-initiated polymerization. This hydrogel had biocompatibility, high exudate absorption capacity, and excellent bacteria barrier activity [202]. In another attempt, due to the biocompatibility, transparency, and ECM mimicking properties of gelatin, microporous structure of the injectable gelatins (IGs) was used as a promising network to heal wounds without any growth factors. IGs accelerated neovascularization by recruiting immune cells. To modify temperature sensitivity of gelatin, it was conjugated to

CHAPTER 14 copolymers that can spontaneously self-assemble in response to temperature. Nondegradable tri-block copolymer, poly(ε-caprolactone-co-lactide)-b-poly(ethylene glycol)-b-poly(ε-caprolactone-co-lactide) (PCLAb-PEG-b-PCLA, called in short PCLA), with gelatin formed temperature-sensitive in situ hydrogel. GelPCLA was used as a wound dressing for protection against external contamination [203]. Chuysinuan and coworkers blended essential oil extracted from Eupatorium adenophorum Spreng plant (Crofton weed) as the emulsion agent with a 10 wt.% gelatin solution to provide a hydrogel. The release outlines and the antimicrobial activity of the hydrogels were investigated on open wound. The results indicated that by increasing initial amount of emulsion in the gelatin hydrogel, the antimicrobial activity also increased. Thus this hydrogel show potential for use as antibacterial wound dressing [204]. Due to the alternative therapeutic properties of ADSCs for recovering damaged skin, gelatin-based hydrogelencapsulated ADSCs were applied as the wound-healing agent in mouse and porcine wound models. The obtained results indicated that, ADSCs isolated from porcine, considerably increased cell growth and differentiation in comparison with mouse isolated one. Furthermore the in vivo study proved the clinical applicability of this platform for skin regeneration [205]. Pectin is a main polysaccharide of plant cell walls. Biocompatibility, low cost and availability are some properties of pectin-based hydrogels. In a study, oxidized pectin (OP) was cross-linked to gelatin. It was used as drug delivery vehicle for wound-healing applications. In addition, swelling ratios, sustained release rate, and stability of the hydrogel increased at 32°C [206]. Triiodothyronine (T3) regulates growth and metabolism of the body and has potency in wound healing by stimulation of growth factor secretion. Alginate, gelatin, and PVA-based hydrogel as wound dressing were used for T3 loading to eliminate exudates and enhance the rate of wound healing. In vivo study on Wistar rat model indicated enhanced skin healing process with increased angiogenesis and collagen deposition [207]. Citric acid (CA) with three carboxylate groups is a nontoxic cross-linking agent. Citric acid could be used as the cross-linker of gelatin matrices to provide reservoir for drug delivery. In a study, carboxyl groups of CA were reacted with amine groups on gelatin during annealing process and then the composite combined with AgNPs. In addition, cefixime was loaded into the prepared composite hydrogel as an antibiotic agent. The results showed that drug release and swelling were dependent on pH, and increased at pH 7.4 compared

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with pH 1.2. In addition, AgNPs incorporated into hydrogel had an antibacterial effect against S. aureus and E. coli. Due to pH sensitivity, this hydrogel can be considered for oral drug delivery and wound-healing applications [208].

3.8. Hyaluronan HA is a linear polysaccharide consisting of repeating disaccharide units of N-acetyl-D-glucosamine and D-glucuronic acid linked by beta-1–4-glycosidic bonds [209]. Due to biological and physicochemical features, this polysaccharide has biomedical applications in drug delivery, tissue engineering, and wound dressing. Properties of HA could be improved by using a cross-linker by covalent connection or chemical modifications [210]. Covalently cross-linked HA-based hydrogels with improved characteristics were applied as a versatile biomaterial for drug delivery and tissue engineering [211]. In addition, they decrease bacterial adhesion in wound site when they are implemented as wound dressing [212]. Rao and coworkers synthesized HA-zinc oxide (ZnO) nanocomposite hydrogel (NCHs) by one-pot synthesis method. In this method, HA hydrogel with 1,4butanediol diglycidyl ether (BDDE) cross-linker as a network structure was pursued by the formation of ZnO nanobelt, which were homogeneously dispersed. This NCHs improved cell proliferation and adhesion, hemostasis, and exhibited antibacterial properties for wound dressing uses. It should be noted that antiadhesive feature of HA hydrogel decreased after incorporation of ZnO [213].

3.9. Keratin Due to the presence of Leu-Asp-Val (LDV) and Arg-Gly-Asp (RGD) motifs in the keratin molecule, it is a suitable polymer for the preparation of the nanofibrous matrix. These motifs improved cellular adhesion, migration, and proliferation of the keratin-based hydrogels [214, 215]. Besides, it has metabolic impacts, biocompatibility, and desirable biological and physical features in wound-healing applications [216]. On the other hand, to overcome poor spinnability properties of pure keratin, different additives were used [217]. Ren et al. designed tannic acid (TA), keratin biocomposite hydrogel coupled to photoluminescent GO quantum dots using citric acid as cross-linker for wound care application. Wound healing occurred by faster keratinocytes proliferation within a compact period [218]. In this hydrogel, TA acted as pH-responsive biomaterial owing to the presence of weakly acidic polyphenolic part [219]. On the other hand, blending TA with keratin

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biomaterial improved the cell adhesion, avoided from secondary infections and preserved wet environment. In addition, GO-quantum dots have excellent properties comprising good electrical conductivity, high carrier mobility, and large surface area [220]. The role of citric acid as cross-linker in this hydrogel was to provide resistance against dissolution in water, improving stability, and swelling capacity [221]. In another study, different composite scaffolds were fabricated by simultaneous electrospinning and electrospraying techniques. The scaffolds were consisted of thermos-sensitive hydrogel particles and fibrous mats. Electrospinning was applied for the fabrication of keratin-based fibrous composite scaffolds. Besides, electrospraying hybridized scaffolds with thermosresponsive hydrogel particles which were consisted of tri-block copolymers conjugated with tragacanth gum. In addition, when BC was added, the diameter of keratin/PEO fibers decreased. Conjugation of hydrogel with fibrous scaffolds did not change its fiber diameter and porous structure [222]. Villanueva and coworkers hybridized a keratin hydrogel with zinc oxide nanoplates. This keratin-based hydrogel demonstrated pH-dependent behavior while incorporated zinc oxide showed antibacterial activity. The gel matrix was increasingly expanded by swelling at basic pH, and collapsed at acidic pH [223]. Since the pH of chronic bacterial wound infection is basic, the keratin hydrogel by increasing the pore size can be swelled leading to the faster release of zinc oxide [98, 100]. The optimum concentration of 5% for zinc oxide was also determined in wound-healing application [223]. Composite hydrogels consisted of microparticle systems, and hydrogels were used as a dressing material in wound healing. In this regard, a gelatin-CUR hybrid was interleaved within an acrylate hydrogel. The role of hydrogel in this system was to prevent from ROS toxicity. Besides, a quercetin-loaded lipidized keratin-based microparticle system was incorporated to the hydrogel to improve cell proliferation and control the release of the antimicrobial therapeutic agent [224]. In a study, a protein-polysaccharide matrix consisted of human hair proteins (KER), konjac glucomannan (KGM) which was loaded with ethanolic extract of Avena sativa (OAT) was used as diabetic wounds dressing. In this system, KGM, biocompatible gelling agent, acted as immunomodulator and stimulator of fibroblast production. KER provided abundant cell adhesion sites, as well as keratinocyte migration, collagen expression, and fibroblast attachment to improve wound

healing. The prepared hydrogel scaffold was inexpensive and safe wound dressing which can effectively improve treatment of diabetic wounds [225]. In another attempt, Ponrasu et al. produced morin (MOR)-loaded hydrogel scaffolds consisted of psyllium seed husk polysaccharide blended with human hairderived keratin using freeze-drying method. The prepared scaffold had good porosity, suitable swelling behavior, notable antibacterial and antioxidant properties to enhance diabetic wound healing. The role of psyllium as gelling agent in this hydrogel scaffolds was wound exudates absorption. On the other hand, keratin as a natural protein could enhance vascularization and collagen synthesis [226].

3.10. Silk SS is a natural hydrophilic protein extracted from silkworm cocoon [227–230]. Due to the cytoprotective and mitogenic activity of serein on keratinocytes and fibroblasts, it is widely applied with other synthetic or natural polymers for the fabrication of hydrogels and scaffolds for skin and tissue repair [231, 232]. Since sericin has poor mechanical strength, natural, or synthetic polymers could interact with hydroxyl, carboxyl, and amine groups of sericin to provide favorable mechanical properties [233–235]. In a study, silver ions were reduced to AgNPs in situ by tyrosine residues of sericin. In this regard, PVA was combined with AgNO3 and sericin to produce AgNPssericin/PVA (AgNPs-SS/PVA) with excellent mechanical performance, biocompatibility and antimicrobial activity properties for wound-dressing application [236, 237]. In another attempt, hydrogel based on PVA, fibroin and sericin was synthesized by high-pressure carbon dioxide (CO2) method. The structure of hydrogel consisted of two layers, a fibroin-based hydrogel mixed with PVA was the upper layer and the lower layer was composed of fibroin. Interaction between PVA and fibroin reduced the gelation time of fibroin hydrogel. In addition, the linkage between PVA and sericin improved flexibility, length and elongation properties of the composite hydrogel [238]. In another study, a superabsorbent and photoluminescence sericin/PVA hydrogel with swelling and hydrophilicity behavior was designed through repetitive freeze thawing. Gentamicin with bactericidal activity was loaded into SS/PVA hydrogel to prevent bacterial growth. The SS/PVA hydrogel with excellent cytocompatibility on mammalian cells had capacity to load and release small molecule drugs [239].

CHAPTER 14 Since, SS could not self-assemble strongly, it is combined with other polymers to provide favorable characteristics of a hydrogel-based wound dressing. In a study, the short-chain poly(ethylene glycol)-diacrylate (PEGDA) was reacted with peptide chains of SS by radical polymerization to improve water content, compressive strength, biodegradability, sol fraction, and gelation time. The optimum mass ratio of SS/PEGDA to form excellent biocompatible hydrogel was 15/85, respectively [240]. As we know, active matrix metalloproteinases (MMP) and the poly-microbial infections, delay wound healing. Thus Sonamuthu et al. hybridized metal chelating dipeptide, L-carnosine, and CUR with the biocompatible silk protein hydrogel (L-car@cur/SF) to overcome polymicrobial infections problem and inactivation of MMP. Silk fibroin (SF) is a natural biopolymer with favorable biodegradability and biocompatibility which can be used for the fabrication of wound dressings. Besides, carnosine (β-alanyl-L-histidine) is a natural dipeptide found in muscles and tissues acted as ionchelating agent to regulate and protect enzyme and protein activity. The chelating effects of Zn2+ ions from the MMP-9 active center inactivate the matrix metalloproteinase-9 (MMP-9). This reaction demonstrated the therapeutic rational of cur/SF matrix using L-carnosine in diabetic wound ulcer [241]. One example of hydrogel scaffold synthesized with 3D printing technique is transparent modified gelatin/ SS hydrogel. The hydrogel scaffold was formed by a rapid prototyping process with synchronous UV radiation using free radical polymerization of GelMA. The properties of this hydrogel scaffold were high transparency, well-distributed macroporous structure, high swelling ratio, and excellent biocompatibility for wound healing and tissue engineering purposes [228]. Another hydrogel as a wound care material is consisted of HA containing corn silk extract (CSE) and AgNPs, which was completely nontoxic and biocompatible. This hydrogel was synthesized by a microwaveassisted green technique. The CSE in the hydrogel structure acted as a biostabilizing and bioreducing/capping agent for the biosynthesis of AgNPs [242]. In another attempt, FGF1-loaded silk fibroin hydrogel (human acidic fibroblast growth factor 1) was prepared. The prepared hydrogel swelled to a 17.3-fold maximum swelling behavior over 12 h [243]. Karahalilo
glu and coworkers produced CUR-loaded silk fibroin e-gel by the electrogelation method to

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stimulate the cell proliferation. The multipotential therapeutic effect of this hydrogel had antifungal and antibacterial properties for the healing of burns and wounds [244]. Silk-elastin has been able to stimulate the migration of macrophages and fibroblasts. Studies indicated that when the concentration of water-soluble silk-elastin was 4% (w/v), hydrogel was formed at body temperature. In an aqueous condition, silk-elastin-based hydrogel covering wound and preserving moist condition without inflammation led to epithelialization of full-thickness skin defects [245].

4. CLINICAL APPLICATION Regardless of the type of the wound, the wound repair process is similar in acute and chronic wounds. Acute wounds such as surgical incisions or traumatic injuries pass through the critical wound-healing processes (hemostasis, inflammation, proliferation, and remodeling) relatively fast. However, the healing process is delayed for more than 12 weeks in chronic wounds as the edges of the wound cannot be approximated [5]. In nonhealing wounds, the cascades responsible for healing process are impaired as there are underlying conditions such as age, diabetes, or nutrition insufficiency [4]. One of the prerequisites for wound dressings is that they should facilitate re-epithelialization and enhance cellular migration to the damaged tissue while providing a moist environment to decrease scar formation. Wound dressings are also used to protect the wound from pathogens while allowing water vapor and air through [3]. Hydrogels are moist occlusive dressings which provide a hydrated environment to promote the own body’s wound-healing process and inhibit the fluid loss. Also they absorb the exudate and keep the wound clean due to improvement of autolytic debridement [6]. Hydrogels are also used for dry and necrotic wound as they possess high water content. One of the main differences of hydrogels with traditional dressings is that hydrogels showed low adherence to the tissue so that the neo-tissue would not be affected and the associated pain during changing the dressing is reduced [8]. Another important issue is the mechanical stability (under tension and pressure) of the dressing during application and its elasticity to allow the dressing to adapt to the applied location. Table 3 represents the biopolymer-based hydrogels available on the market. Collagen-based products provide a favorite

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TABLE 3

Biopolymer-based wound dressings on the market [7]. Product

Main constituent

Company

Algisite M

Alginate

Smith & Nephew

Kaltostat

Alginate

Convatec

Biostep

Collagen/ alginate/ carboxymethyl cellulose/gelatin

Smith & Nephew

Puraply

Collagen

Organogenesis

Promogran Matrix

Collagen/ oxidized cellulose

Johnson & Johnson Wound Management

Fibracol

Collagen/alginate

Johnson & Johnson Wound Management

Chitoderm plus

Chitosan

Trusetal Verbandstoffwerk GmbH

Tegaderm alginate Ag

Alginate, Ag

3M

SilverCel

Alginate, Ag

Acelity

Algidex Ag

Alginate, Ag

DeRoyal

Kendall Curasorb

Alginate

Covidien

Regenecare HA

Hyaluronic acid

MPM Medical Inc.

Hyalomatrix

Hyaluronic acid

Anika Therapeutics

Apligraft

Bovine collagen gel, fibroblasts, keratinocytes

Organogenesis

Alloderm Regenerative Tissue Matrix

Collagen and elastin

BioHorizons

TissueMend

Collagen

Stryker Orthopaedics

environment attracting different cell types needed for the skin regeneration and deactivate free radicals or matrix metalloproteinases, which negatively affect the healing process though enzymatic degradation of the neotissues [246]. Another important issue in developing wound dressings is the antimicrobial properties. Chitoderm plus is a nonadherent chitosan cryogel with bacteriostatic

properties. In addition to its antimicrobial activity, positively charged chitosan adsorbed growth factors and enhance angiogenesis and proliferation [8]. Campani et al. studied the effects of Chitoderm pressure ulcers in patients who needed bed rest. They demonstrated that the treatment was effective in 90% of the patients with the reduction of the area of the lesion [247]. To improve the antimicrobial efficacy, silver ions are incorporated into the alginate-based dressings. Calcium alginates used as natural hemostats with gel-forming capability help to change the dressing without trauma. However, alginatebased gels absorb fluids up to 20 times of their weight and that might be the reason of their adherence to the wound base if were not replaced for a week [248]. Silverlon is a calcium alginate-based dressing with a silver contact layer. The dressing absorbs exudate and keeps the environment moist while the release of silver ions provides an antimicrobial action. Algidex Ag hydrogel composed of alginate, maltodextrin, and silver ions is effective against broad spectrum of microorganisms such as Methicillin-resistant S. aureus and fungi. Studies have shown that hyaluronan possesses antimicrobial properties. Also hyaluronan-based dressing indicated angiogenic and blood clotting properties. Hyalomatrix and Hyalograft are matrices made of HA and autologous fibroblasts. The use of autologous fibroblasts provides secretion of new ECM into the wound with no immune response. Such dressings are called as skin substitutes [6]. Skin substitutes are made of biologically derived materials combined with platforms to allow its placement on the wound. The presence of biologically active materials on wounds leads to continuous release of growth factors while the presence of live cells helps maintaining the structure and function of dermis. Apligraft is a composite allograft made of bovine collagen gel and neonatal fibroblasts and keratinocytes. It is FDA-approved for diabetic foot ulcers and healing of venous leg ulcers that showed no response to other treatments for 1 month [249]. Such dressings are highly expensive and their cost-effectiveness should be evaluated. However, studies showed that the use of skin substitutes leading to even 1 day less hospitalization is an enormous cost saving [3]. The physicochemical properties of hydrogels made them tunable materials with versatile properties. In a recent study, injectable hydrogels made of chitosan and konjac glucomannan were designed with long lasting self-healing behavior. Inherent antibacterial activity of chitosan combined with pain relief and easy application of the dressing significantly shortened healing time and accelerated epithelialization [250].

CHAPTER 14 Over the last decades, tremendous efforts have been put into designing functional dressings. There have been major improvements in developing antibacterial/regenerative platforms leading to fabrication of antibacterial dressings, moisture-retentive dressings, and tissueengineered grafts. Despite all the significant progress, it is necessary to setup multifunctional dressing with combinatory performances to overcome the nonhealing wounds in patients with underlying diseases.

5. FUTURE PERSPECTIVE Despite the amazing advancements in developing new materials and the production technologies, the management of nonhealing wounds is still a big challenge in medicine. During the last 50 years, hydrogels have gained full attention as they showed fascinating properties as wound dressings. There are many studies focusing on the design of biocompatible hydrogel-based dressings with the aim of promoting the healing process. Biopolymers are great candidates as ingredients to prepare hydrogels mainly because of their biocompatibility which enable them to show minimal damage to the surrounding tissue and reduce the possible adverse reactions. However, the performance of such dressings should be enhanced to resolve the increasing need. An efficient wound dressing should promote tissue remodeling and regeneration while reducing the formation of scars. To this aim, multifunctional dressings capable of facilitating angiogenesis, cell migration and differentiation, and protecting against infections are essential. Besides, it would be ideal if specifications such as the age of the patient or disease related differences are considered for developing the wound dressings. Hydrogels containing bioactive agents and regenerating cells might meet the stated need. The promising approach could be producing multifunctional hydrogel-based dressings using novel manufacturing technologies such as 3D printing to cover the defected area. Due to the current improvement in technology, this approach is not far from reality. However, reproducibility of the dressing and the reliability of the wound models are important factors which should be considered. Therefore we believe that the remarkable advancement in fabrication technologies and the materials science offer outstanding opportunities enabling us to develop multifunctional wound dressings with the aforementioned capabilities.

ACKNOWLEDGMENTS The authors are grateful to the Mashhad University of Medical Sciences.

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DECLARATION OF INTEREST The authors declare that they have no conflicts of interests.

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CHAPTER 15

Novel Controlled Release Pulmonary Drug Delivery Systems: Current updates and Challenges

DALJEET S. DHANJAL a,* • MEENU MEHTAb,i,* • CHIRAG CHOPRAa • REENA SINGHa • PARVARISH SHARMAb • DINESH K. CHELLAPPANc • MURTAZA M. TAMBUWALAd • HAMID A. BAKSHId • ALAA A.A. ALJABALIe • GAURAV GUPTAf • SRINIVAS NAMMIg • PARTEEK PRASHERh • KAMAL DUAi,j • SAURABH SATIJAb,i a

School of Bioengineering and Biosciences, Lovely Professional University, Phagwara, Punjab, India, School of Pharmaceutical Sciences, Lovely Professional University, Phagwara, Punjab, India, c Department of Life Sciences, School of Pharmacy, International Medical University, Bukit Jalil, Kuala Lumpur, Malaysia, dSchool of Pharmacy and Pharmaceutical Sciences, Ulster University, Coleraine, Northern Ireland, United Kingdom, eFaculty of Pharmacy, Department of Pharmaceutics and Pharmaceutical Technology, Yarmouk University, Irbid, Jordan, fSchool of Phamacy, Suresh Gyan Vihar University, Jaipur, India, gSchool of Science and Health, Western Sydney University, Penrith, NSW, Australia, hDepartment of Chemistry, University of Petroleum and Energy Studies, Dehradun, India, i Discipline of Pharmacy, Graduate School of Health, University of Technology Sydney, Ultimo, NSW, Australia, jPriority Research Centre for Healthy Lungs, Hunter Medical Research Institute (HMRI) & School of Biomedical Sciences and Pharmacy, University of Newcastle, Callaghan, NSW, Australia b

1. INTRODUCTION The lung is a highly susceptible internal organ to different infections and injuries as it continuously gets exposed to chemicals, particles, and infectious organisms present in the air [1, 2]. Worldwide, two billion people are prone to toxic gases emitted by ineffective burning of fuel in fireplaces or poorly ventilated indoors [3, 4]. Approximately, 1 billion people inhale the polluted air and tobacco smoke. These factors are leading to respiratory impairment responsible for causing disability and mortality globally in all social classes [5]. Generally, deprived living conditions are the main causes for increasing vulnerability among the group of people suffering from respiratory disorders. Four of the diseases such as asthma, chronic obstructive pulmonary disease (COPD), tuberculosis (TB), and lung cancer are the common causes of mortality and severe illness worldwide [6–9]. These wide arrays of the disease are termed as “chronic respiratory diseases (CRDs).” All around the world, millions of people of varied ages are affected by these CRDs. A large proportion of people *Authors having first authorship equal contribution.

(i.e., approximately 50%) are from developing countries and its prevalence is increasing exponentially, especially in children and elder people [10]. The burden of prevailing CRDs is affecting the life of the affected individuals. The World Health Organization (WHO) has stated that in 2005 approximately 4.6 million people died earlier because of these CRD, and this number is eventually going to grow in the nearby future [11]. However, the preventive measure can aid in managing the burden of CRD in both developing and developed countries, but cost-effectiveness of the interventions restricts it usage ([12]). Many risk factors such as allergens, indoor pollution, occupational agents, tobacco smoking, and in few cases diseases like sickle-cell or schistosomiasis have been identified and their prevention can play a significant role in regulating mortality and morbidity rate worldwide [12]. Furthermore, insufficient attention is given to preventable CRD and their associated risk factors by community, families, government officials, healthcare, patients, and media. Hence, people suffering from CRD remain undiagnosed and untreated. Although there are many pharmacotherapies available for treating these CRD and manage the suffering patient under

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medication for their lifetime [13], but limitation and side-effects associated with drugs used for controlling and managing CRD have encouraged us to search for novel treatment [14]. In recent years, advancement in the field of nanotechnology and their role in medicine has led us to develop multifunctional nanoparticles, which can act as nanocarriers for loading different drugs ( [15]). These nanocarriers have emerged as a valuable approach for drug delivery with features like prevention of drug degradation, allows the controlled release of the drug, and targeted delivery of the drug at the target site [16]. Therefore this chapter intends to feature the efficacy and physiological aspects of targeted drug delivery, mechanism of pulmonary drug administration. Moreover, it will also emphasize different nanocarriers for drug delivery. Additionally, this will help in gathering deep insight into the challenges and complications associated with developing the pulmonary disease, which will open new avenues for pharmacists with the purpose of minimizing the medical and technical gaps. This chapter is primarily divided into five sections including the introduction as Section 1. Section 2 describes the background and rationale of the review work envisaged highlighting the global scenario of COPD and anatomy and physiology of lungs to give clear understanding of the underlying mechanism involved in the drug delivery. Section 3 describes the methods and mechanism of drug administration. Section 4 describes about the nanocarrier drug delivery, whereas Section 5 describes about the various drug delivery approaches in controlled pulmonary drug delivery systems along with the clinical studies and challenges associated with controlled drug delivery. Section 6 discusses about the future directions of controlled pulmonary drug delivery systems, and Section 7 provides the concluding remarks.

2. BACKGROUND 2.1. Global Scenario Millions of people worldwide of all ages are suffering from CRDs such as asthma, COPD, TB and lung cancer. Approximately, 500 million people suffering from these CRDs are from developing nations [17]. In general, CRDs are the diseases associated with airways and different parts of lung. Globally, millions of people are suffering from preventable CRD such as asthma and COPD, out of which 50% of them are from developing countries. And prevalence of these CRD is exponentially increasing especially in children and elder people [18]. The burden induced by these CRDs has adversely affected the lives of many people. In past, the WHO

estimated the death 4.6 million people due to CRDs and also stated that this number will considerably increase in coming future. The major risk factors that have been identified to be responsible for these CRDs includes allergens, air pollution, tobacco smoking, and some disease such as sickle cell anemia and schistosomiasis [19]. Asthma, a prevailing inflammatory disease of airways, is especially linked with hyperresponsiveness of airway organs which causes hindrance in airflow [20]. The most important risk factor of asthma is allergen sensitization. It is also linked with inflammation in nasal mucosa and rhinitis. Both children and adults are highly prone to this disease [21]. Around 300 million people of all cultural backgrounds and ages are suffering from asthma. The two studies, that is, European Community Respiratory Health Survey and the International Study of Asthma and Allergies in Childhood have comprehended the asthma prevalence worldwide [22]. This survey revealed that the global prevalence of asthma has parallelly increased in all countries in relation to allergy. Urbanization and the modern lifestyle are also the factors associated with increasing asthmatic cases. It has been estimated that asthma accord for 250,000 deaths every year globally. Even, prevalence is high in those countries where access to drugs is limited. The disability-adjusted life years have ranked 22nd position to asthma worldwide. The countries such as Australia, Brazil, India, Northern Europe, North America, and some parts of Latin America [23]. COPD, a chronic respiratory disease of airways, which obstructs the airflow during breathing due to inflammatory response to noxious agents such as tobacco smoke, biomass fuel, and industrial contaminants [24]. This disease is highly prevalent in elder people. Globally, COPD has affected approximately 200 million people, out of which 65 million were already suffering from severe respiratory disease. The leading cause of this disease is cigarette smoking which causes damage to lung tissues and obstruct the airway via inflammation and elevated mucus production in bronchi [25]. The clinical symptoms involve breathlessness and cough. As per the prediction by WHO, COPD will attain the 5th position for disability and 3rd position for mortality in 2020 [26]. About 12 Asian countries have been comprehended to be suffering from COPD. Currently the burden of obstructive lung disease is caring out the survey studies in developing countries. The countries such as Australia, Brazil, Denmark, England, Iceland, India, Spain, and the United States are having the high number of COPD patients [27].

CHAPTER 15 Novel Controlled Release Pulmonary Drug Delivery Systems TB, an infectious disease commenced by the infection of bacterium Mycobacterium tuberculosis, which affects the lungs and other pulmonary organs. This disease is known to spread from sick person through the expulsion of bacteria in the air because of cough or sneeze [28]. The person infected by M. tuberculosis shows symptoms such as cough with sputum and blood (in few cases), chest pains, fatigue, fever, weight loss, sneezing, and night sweats. The people who do cigarette smoking and have compromised immune systems, diabetes mellitus, HIV are at higher risk to develop TB [29]. In 2018, TB has been accorded for 10 million people globally of all age groups. The development of multidrug resistant (MDR) from rifampicin has also become the global concern, as the cases of MDR-TB is increasing exponentially [30]. The high burden of TB is accounted by countries such as Bangladesh, China, India, Indonesia, Nigeria, Pakistan, Philippines, and South Africa [31]. Lung cancer, a malignant lung tumor disease caused by smoking and is responsible for the growth of unregulated of genetically altered cells going repeated division [32–35]. In 2012, 8.2 million deaths have been accorded for this disease as estimated by GLOBOCAN. It generally occurs by smoking which causes mutation in protective genes and damages the DNA. This alteration and damages accumulate with time, which is the reason that people who start or quit smoking also suffer from this deadly disease [36]. Other than smoking, there are other risk factors associated with lung cancer such as asbestos, passive inhaling of biomass fuel, diesel exhaust, passive inhalation of tobacco, and other environmental factors [37, 38]. The clinical symptom of this disease involves chest pain, fatigue, hoarseness, persistent cough, loss of appetite, excessive weight loss, and shortness of breath [39]. As per the consensus of WHO, lung cancer holds the 2nd position for mortality globally and moreover, the 70% of death because of lung cancer occurs in low- and middle-income countries such as Australia, Belgium, Denmark, France, Hungary, India, Ireland, Netherlands, New Zealand, Norway, and the United States [40, 41].

2.2. Anatomy and Physiology of the Lungs Lungs are a pair of air-filled spongy organs that performs the function of gaseous exchange and delivers the oxygen to every cell. It comprises a total of five lobes, where three are in the right lung and two are in the left lung [42]. The lung interior contains smaller air-passages, alveoli, bronchi, blood vessels, and lymph tissues. The bronchi portion of the lung is further subdivided into primary (1 degree) and secondary (2 degree) bronchi as well as bronchioles and at last the alveoli. Moreover,

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the lung contains more than 0.3 billion alveoli [43]. Additionally, all the alveolus is covered with respiratory capillaries, which form a complex network containing around 0.28 trillion capillaries, facilitating the large surface area of about 70 m2 [44]. Primarily the alveolar gasexchange takes place at the lining comprising alveolar epithelium, endothelium, and interstitial cell layers. Also, among the alveolar epithelium and capillaries, single endothelial layer is also present. The distance among the capillaries and alveoli is very less approximately 0.5 μm, this acute thinness owes to the blood-gas interface, where exchange of gas take place by diffusion [45]. These alveoli are covered with alveolar fluid as well as mucus, which are primarily made up of surface proteins and phospholipids. The phospholipid surfactant layer of alveoli allows the proper functioning of gaseous exchange and reduces the surface tension in the alveoli [45]. Moreover, these distant respiratory passages are maintained intact by connective tissue, which are bounded by several types of cells such as fibroblasts, lymph vessels, macrophages, and nerves cells. This aid as an ideal site for drug administration as it has access to both lymphatic and pulmonary system [46].

3. METHODS 3.1. Mechanism of Drug Administration Over the past few decades, the systemic absorption of different types of drugs after pulmonary applications have been tested on animals and humans [47]. Through the pulmonary pathway, the therapeutic agent can be administered in our body through two ways, that is, intranasal and inhalative administration. The intranasal administration has anatomical constrain like narrow lumen of airway [48]. Whereas oral inhalative administration has been found to better than intranasal administration, as it allows the administration of small particles and reduces the loss of therapeutic drug before reaching the target site [49]. Moreover, oral inhalative administration is further categorized as intratracheal inhalation and intratracheal instillation. Most common method is “intratracheal instillation” in which minute amount of therapeutic solution is delivered by special syringe into the lungs. As it allows the quantitative and fast delivery of drug to the lungs [50]. Additionally the confined drug deposition is attained that also on relatively limited absorption area. Hence, this makes this process easy and nonexpensive but drug distribution in this case is nonuniform [51]. During preclinical assessment on animals, intratracheal instillation method was used to determine the absorption capacity

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and bioavailability of therapeutic drugs in pulmonary system, especially because of the effectiveness and precision associated with this approach [52]. But results obtained from these studies are hard to transfer this approach as an aerosol form in human. Whereas inhalative administration method uses the aerosol technique and facilitates the uniform distribution of drug with effective penetration [53]. Despite this the approach is expensive and ineffective in determining the exact dosage in pulmonary system. There are three main mechanisms such as diffusion, gravitational sedimentation, and inertial impact that are involved in drug deposition through aerosol administration method. In case, where size of drug is larger, then deposition of drug take place by gravitational sedimentation and inertial impaction [53]. Whereas when the drug is smaller in that diffusion mechanism comes into play and follows the rule of Brownian motion. Other than the morphology aspect of the pulmonary system, parameters like the size of drug droplets or particles and geometry are very important. Additionally the size of the drug droplet or particles in relation with diameter, surface charge and shape are also crucial, as they determine the influence of deposition of drug via pulmonary route [54]. On analyzing the mechanism of drug administration in the pulmonary system, although the conventional techniques are effective in relieving the patient suffering from respiratory disease but during the ADME process of drug, the concentration of the drug at target site is lower than the desired dosage. Hence, nanocarrier drug delivery system has gained the interest of pharmaceutical industry to invest in the development of these nanocarrier system [55].

4. NANOCARRIER DRUG DELIVERY SYSTEMS 4.1. Advantages of Nanocarrier Drug Delivery System In recent years, the use of nanotechnology in the field of medicine has gained significant attention [56–59]. Particularly, the nanocarrier drug delivery system is the emerging field due to its associated advantages such as improved circulation time of drug, high concentration of drug at targeted site, reduced degradation and loss of drug, and most important is its easy administration procedure [60–62]. With the application of nanoparticle, it is easy to regulate the release of drug at the targeted site. Generally the nanocarriers are the colloidal particles within the size range of 10–200 nm [63, 64]. The selection of nanocarrier depends on the nature of drug which is to be entrapped inside these nanocarriers by

different techniques that are retained on the basis of their interaction between nanocarrier and drug. As the interaction of drug is not compatible with all the types of nanocarriers [65]. To date, various different types of nanocarrier systems have been developed named carbon nanotubes, dendrimers, liposomes, mesoporous silica, micelles, polymeric nanoparticles, protein nanoassemblies, and many more. On the basis of the material and variation on the surface, these nanocarrier systems possess different drug release characteristics and properties [16]. These nanoparticles share a high resemblance with biological entities such as viruses and proteins, which allows them to interact with cell surface and with the cells. Due to which nanocarrier system has gained a lot of attention and various new nanocarriers are being developed with different properties, making it difficult to underline all new findings [66]. Still, some of the properties which are important for developing an effective drug delivery nanocarrier system is illustrated in Fig. 1 and comprehended below:

4.1.1. Easy surface amendment Mostly surface amendment of nanocarrier system is done to improve the biodistribution and circulation time. For example, hydrophobicity often aids in binding of these nanocarriers with blood components. Whereas nonamended surface hydrophobic nanocarriers are easily removed by our own system [67]. Therefore to enhance the circulation time of these nanocarrier in blood, these nanocarriers are coated with hydrophilic polymers or surfactants like chitosan and polyethylene glycol (PEG). Interestingly, PEGylated nanocarrier have shown improved mucous penetration, whereas chitosan-amended nanocarriers have shown improved circulation time [68]. These benefits of nanocarrier system highlight their importance in treating CRDs, as it prolongs the availability of therapeutic agent at the target site. Other than modifying these nanocarrier system with chemical agents, they can also be amended by biological fluid which forms corona properties on the surface, which changes the properties of nanocarrier [69]. Additionally, the modification of these nanocarrier with phospholipids has drastically improved the cellular uptake and the toxicity of it [70].

4.1.2. Targeted delivery Nanocarrier drug delivery system offers the additional benefit of specific cell or tissue targeting, which has greatly improved therapeutic effect of the drug and reduces the chances of drug toxicity [71]. This targeting is achieved by both the active and passive drug delivery method. Passive accumulation of these nanocarriers has

CHAPTER 15 Novel Controlled Release Pulmonary Drug Delivery Systems

257

FIG. 1 Diagrammatic illustration of the properties important for developing nanocarrier system for drug

delivery.

been found to improve permeability and retention time of the drug. Moreover, charge on the nanocarrier system also aid passive targeting [72]. For example, paclitaxel loaded in cationic liposomes were compatible to target the endothelial cells in tumors of cancer patient. Moreover, passive uptake of these nanocarriers by macrophages aid in targeted drug delivery, as in case of TB the microbe M. tuberculosis resides in the macrophages [73, 74]. On the other hand, active targeting involves the interaction between the targeted ligand with nanocarrier surface which has an affinity for specific molecule in the diseased cell. For example, liposomes loaded with doxorubicin have been developed which targets integrin molecule on colon cancer cells and inhibits their progression, which was an advantage of this modified liposome over the traditional liposomes. Hence, conjugation of the cell-specific entity on the nanoparticle aid in targeting the specific cell and increases the specificity of the therapeutic agent [75].

4.1.3. Regulated release of drug For therapeutic success, regulated release of therapeutic agent from nanocarrier is important. Control release of drug from these nanocarriers could be sustained or stimuli response [76]. In case of liposomes and polymeric nanocarriers, there is a sustain release of the drug either by the process of diffusion or gradual degradation these nanocarriers over the period of time.

Whereas stimuli-responsive release of the therapeutic agent is more targeted approach and can be achieved by altering the biological environment by changing the pH or disease-targeted enzyme [77]. For example, the nanocarrier synthesized with pH sensitive linkers attached on the surface, allows the easy removal of outer coat made up of polymer on their uptake [78]. In another example, PEG-peptide-lipid conjugate nanocarriers were used for the responsive release of the therapeutic agent, as the PEG molecules get removed from the nanocarrier surface by the action of matrix metalloproteinases, as they are overexpressed in the tumor cells [79, 80].

5. DRUG DELIVERY APPROACHES FOR PULMONARY RESPIRATORY DISEASE There are numerous nanocarriers drug delivery system that have been developed for treating these CRDs as depicted in Fig. 2. Additionally, these nanocarrier drug delivery systems have been briefly discussed later in the following subsections:

5.1. Liposomes Liposomes are spherical-shaped nanocarriers of lipid bilayers made up of cholesterol and nontoxic phospholipids. Generally, it comprises an aqueous core in center which is surrounded by lipid bilayer resembling the

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FIG. 2 Systematic illustration of the nanocarrier drug delivery system for treating chronic respiratory diseases.

shape of membrane, separating the inner core from the external agents [81]. The natural or synthetic lipids can be used for synthesizing liposomes, moreover, they are not exclusively dependent on lipids, even polymers are used for synthesizing liposomes (also known as polymersomes) [82]. The biocompatibility, and biodegradable properties of these nanocarrier, makes it suitable for medical research [83]. The major advantage of liposomes is its ability to form compartments and solubilize both hydrophobic and hydrophilic material inside it. It can be used to load variety of biomolecules such as drug molecules, proteins, plasmids, nucleotides, and viruses [35]. Moreover, specialized liposomes can be prepared

by surface modification by using antibodies or targeted ligands which recognize and bind with targeted molecule. The modification improves drug absorption ability, prolongs biological half-life, reduce metabolism and toxicity [84]. These properties of liposomes make it the suitable nanocarrier for targeted drug delivery. Recently, budesonide-containing liposomes have been developed and evaluated for managing asthma [85]. Various nebulizing liposome systems have been formulated and reached clinical trials like Arikace and Pulmaquin are the one which have reached the advanced stages of clinical trials [86]. Additionally, usnic acid encapsulated liposomes have also been

CHAPTER 15 Novel Controlled Release Pulmonary Drug Delivery Systems evaluated their interaction with anti-TB agents effective against clinical MDR-TB isolate [87]. Lately, stealth liposomes have gained significant attention in lung cancer treatment as they circulate in blood system for longer time and increase the chances of targeted delivery of therapeutic agent at the affected site and regulate the spread of lung cancer [60].

5.2. Niosomes Niosomes are multilamellar vesicular nanocarrier structure similar to liposomes and are composed of cholesterol and nonionic surfactant. These multilamellar vesicular serve as an alternative to liposomes [88]. These nanocarriers comprise two components such as additives and nonionic surfactant. The nonionic surfactant includes alkyl amides, alkyl ethers, alkyl esters, amino acid compounds, and fatty acid. These nonionic surfactants are used to form the vesicular layer of nanocarrier whereas additives like charged molecules and cholesterol are used for preparing niosomes [89]. These additives aid in improving the fluidity, permeability, and rigidity of bilayer. These nanocarriers are used to deliver antigens, bioactive agents, and hormones as it protects them from premature degradation or inactivation because of unwanted immunological effect. Moreover, these nanocarriers are also used to overcome challenges such as instability, insolubility, and rapid degradation of therapeutic drugs [90]. Niosomes have emerged as a promising therapeutic carrier system. Recently, salbutamol sulfate containing niosomes have been developed and evaluated for stability, sterility, and pharmacological studies for therapeutic potential against Asthma [91]. Even, polysorbate 20 has been encapsulated in niosomes for targeted delivery in COPD patients [92]. The comprehended literature reported about synthesis of ofloxacin and rifampicinencapsulated niosomes, which has been further evaluated against drug-resistant M. tuberculosis. The result obtained from this study showed significant inhibition and controlled growth of drug resistant M. tuberculosis [93]. Lately, group of researchers have developed Adriamycin containing niosomes and evaluated in lung cancer-bearing mice. The result revealed that growth of tumor delayed for longer time and suggested therapeutic administration of adriamycin can be enhanced by using it as niosomes [94].

5.3. Nanoparticles 5.3.1. Magnetic nanoparticles These are those nanocarrier system which gets influenced under the magnetic force and get to the target site within the body either by active or passive approach,

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due to the presence of ligands present on its surface [95]. Usually, these nanocarriers are composed of super paramagnetic entities as core, which is of the size less than 25 nm. This core contains the material such as nickel, gold, iron, and cobalt and is covered by surface coat which obstructs the interaction of core with the other particles [96]. Moreover, these magnetic nanoparticles show thermic effect when magnetic field is applied from external source, as it initiates the apoptosis signal when 42°C temperature is attained and triggers direct killing on reaching the temperature to 45°C. Whereas in case of nonbiodegradable magnetic nanocarrier, they are coated with layer which leaches the magnetic core from them and the core gets excreted through kidneys [97]. Due to biocompatibility, good loading capacity and definite shape of this nanocarrier make it the effective system for both radiotherapy and diagnosis purpose [98]. The advancement in the field nanotechnology has enabled to developed polyethylene glycol-coated magnetic nanocarriers for noninvasive magnetic resonance imaging (MRI) and specific targeting in asthmatic patient [56, 59, 99]. Another study reported about the synthesizing of antibody-associated magnetic nanoparticles for specifically targeting the alveolar macrophage and MRI in animal model with COPD [100]. Moreover, iron oxide magnetic nanoparticles have emerged as bacterial detection and therapy approach because of its magnetic properties. And, now this approach is also used for imaging and therapy for treating infection caused by M. tuberculosis [101]. One more study reported about the development of Fe3O4surface-coated magnetic nanoparticle conjugated with poly(lactic-co-glycolic acid) for controlled drug delivery of quercetin at targeted site in lung cancer cells [102].

5.4. Polymeric Nanoparticles Polymeric nanoparticle is an effective strategy for targeted delivery of therapeutic agents as they are easy amendable and morphology can be easily altered as per the need [103]. Different types of polymers like alginic acid, gelatin, polylactic acid, chitosan, polylactideco-glycolide, and polycaprolactone are generally used for constructing these nanocarriers. Further, supplementing of sulfide bond to these polymeric nanoparticles provides it the ability to control the release of drug [104]. Usually, cationic polymers are found to be toxic and they also get accumulated at the targeted site because of their poor biocompatibility and nondegradable nature. This provokes the need of consistent monitoring of toxicity, as interaction of these cationic polymers with biosurfactant elicits the complication

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like breathlessness in immune-compromised patient [105]. Recent developments in the field of nanotechnology have improved the biocompatibility and biodegradability of polymeric nanoparticle making it less toxic. Moreover, these modifications have largely improved the recognition ability and activity during targeted drug delivery [57, 106, 107]. Even though inhaling steroids is the first choice of therapeutic agents to treat asthmatic patient. But, inhaling therapeutic agent decreases the bioavailability of therapeutic agent at the targeted site. Hence, to improve the efficacy and to reduce the side effects of the drug, betamethasone phosphate containing polymeric nanoparticles have been developed. Additionally, its effects are assessed in asthmatic murine model [108]. Moreover, RNA interference, that is, RNAi-based treatment approach is believed to be an endogenous approach for regulating the gene expression. Furthermore, microRNAs (miRNAs) can be used as traceable targets for COPD, for which miR146a was conjugated along with poly(glycerol adipate-co-ω-pentadecalactone) and PGA-co-PDL polymeric nanoparticles to suppress the gene expression of target IRAK1. The result obtained from this study revealed that poly(glycerol adipate-coω-pentadecalactone) and PGA-co-PDL polymeric nanoparticles is a suitable candidate for targeted drug delivery system for treating COPD [109]. Recently, alveolar macrophage targeting pyrazinamide containing polymeric nanoparticles have been developed to regulate the dosing frequency and profile of pyrazinamide and treat pulmonary TB [110, 111]. Another study reported by the development of Crizotinib-encapsulated polylactide tocopheryl polyethylene glycol 1000 succinate polymeric nanoparticle which regulated the release of the therapeutic agent at target site and caused noteworthy toxicity in NCIH3122 cells of lung cancer. Additionally, another group of researchers developed polycaprolactone/poly(ethylene glycol)/polycaprolactone (PCEC) polymeric nanoparticle loaded with paclitaxel and is used in combination with chrono-modulated chemotherapy procedure in lung cancer [112].

5.5. Solid Lipid Nanoparticles Solid lipid nanoparticles (SLNs) are the improved alternative of conventional delivery system. Electron microscopic techniques like scanning electron microscopy (SEM) and transmission electron microscopy (TEM) have aided in determining the range between 50 and 1000 nm and spherical shape of these nanoparticles [113]. The biocompatibility and safety profile of these nanoparticles in pulmonary system has made it recommended drug delivery system. These SLNs comprises

solid lipid in the 0.2%–30% (w/w) range, which allows it to get dissolved in the aqueous solution. Furthermore, to improve the stability of these nanoparticles, 0.5%– 5% surfactants are used [6, 114]. These SLNs come in the category of nanoparticles comprising lipids which maintain their solid structure even at room temperature. Moreover, they have additional advantages like they are easily amendable, biocompatible with lipophilic drug, less toxic, efficient to deliver therapeutic agents at targeted site in contrast to other carriers. The large surface of these nanoparticles allows the loading of high amount of therapeutic agent and protect it from environmental factors, as a resultant bioavailability of drug increases [115]. The main reason for this nanoparticle to be ideal targeted drug delivery system is that its traits are the amalgam of emulsion, liposomes and polymeric nanoparticles. Bee wax, cholesterol butyrate, Dynasan, and emulsifying wax are the common lipids material used for developing these nanoparticles [35]. The previously published literature has highlighted the therapeutic potential of curcumin. Therefore the group of researchers have developed curcumin-loaded SLNs to improve the efficacy of curcumin in rat model of asthma, which showed this approach will be promising in treating asthma [116]. The long-acting β agonist is an effective treatment approach for COPD patients. Hence, Salmeterol Xinafoate containing SLNs have been developed, characterized and assessed on bronchial epithelial cells in the model organism and found to be effective in treating COPD [117]. Additionally, inhalable formulation in powder form has been developed of Ethambutol-containing SLNs for treating TB [118]. Lately, paclitaxel-containing SLNs and coated with chitosan as well as folate-poly(ethylene glycol) has been formulated and reported about the reduction in IC50 value in M109HiFR cell line of lung cancer [119].

5.6. Dendrimers Dendrimers are the branched synthetic polymers that are of size ranging from 10 to 100 nm. These nanocarriers have well-defined size and structure which makes them unique among the other nanocarriers. Chiefly, they consist of globular structure with core, dendrons and a surface-active molecule which makes it the perfect candidate for the controlled drug delivery system [120]. Dendrimers are composed of from a core element (with two identical functional groups), dendrons (monomers linked to the core, forms a layer around it) that are further attached to the bifunctional surface molecule. The presence of bifunctional group on the surface of nanocarrier makes biocompatible and is responsible for physiochemical properties of the dendrimers [121]. These

CHAPTER 15 Novel Controlled Release Pulmonary Drug Delivery Systems nanocarriers are chemically synthesized by controlled polymeric reaction embracing both hydrophobic and electrostatic interactions. Moreover the easy amendment of surface, biocompatibility, high loading capacity, multiple point of conjugation and proportioned shape of this nanocarrier makes it the effective nanocarrier drug delivery system [122]. Recently, beclometasone dipropionate (BDP) containing polyamidoamine (PAMAM) dendrimers have been developed and evaluated for aerosolization, solubility and drug release properties. The result of the study revealed that these BDP-dendrimers could be potential nanocarrier system for inhalation using both air-jet and vibrating-mesh nebulizers in patient suffering from pulmonary disease like asthma [123]. With the formulation of autophagy-inducing compounds, now researchers are expanding their horizon by developing dendrimer formulation for inhalation for stable, longer half-life and targeted delivery of therapeutic agent [124]. The published literature has reported about the development of mannosylated G5 EDA-PPI dendrimer containing ethylene diamine in its core for selective delivery of rifampicin in patients suffering from TB [125]. Another study reported about the formulation of polyplex for targeted delivery of RNAi genes as a treatment for cancer. This polyplex was modified form of PAMAM dendrimers with the help of bromodecanoic acid and PEG. The in vitro evaluation on A549 cell line of lung cancer of this polyplex revealed about the efficacy of this nanocarrier in knocking down the Bcl-xL expression and inducing apoptosis in cells [126].

5.7. Micelles For targeted and regulated release of hydrophobic antineoplastic drug, micelles nanocarrier systems are usually used. They serve as the candidate of interest as they are composed of hydrophobic core of co-polymers. This co-polymer hydrophobic core allows the loading of hydrophobic drugs whereas hydrophilic shell allows the attachment of hydrophilic drugs [127]. Moreover, hydrophilic shell of this nanocarrier system also enhances is stability in the biological system. This nanocarrier system is found in different sizes ranging from 20 to 100 nm, which makes it effective for loading high amount therapeutic agent [128]. As micelles provide the benefit like enhanced drug permeability, improved circulation time and uniform distribution has made it suitable candidate for passive-targeted therapy. Surface modification of these micelles substantially improves it targeting efficiency and biodegradable nature of this nanocarrier system makes it effective tool for drug delivery at targeted site [129].

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As angiogenesis is one of the characteristic features of asthma and to regulate this, anti-angiogenesis therapeutic agents have gained significant attention. The study was conducted in which αvβ3-mixed micelles was loaded with docetaxel-prodrug and evaluated for controlling the hyper-response to mite dust triggered asthma in rat. The result obtained revealed about effectiveness of the nanocarrier system in ameliorating the inflammatory response in rat model [130]. Another study reported about self-assembling micelle formulation containing chafuroside A, which was evaluated for anti-inflammatory effect in COPD rat model. The result obtained from this study showed improved dissolution of chafuroside A and target delivery of the therapeutic agent suggesting the benefit of this approach [131]. Moreover, the group researchers have developed rifampicin-loaded HPMA-PLA polymeric micelles, which showed improved result against both resistant and sensitive M. tuberculosis [132]. Another study reported about the development of matrix metalloproteinase 2/9 (MMP2/9)-triggered-release micelles, which were evaluated for both in vivo and in vitro studies against lung cancer. The approach showed effective result in reducing the toxicity induced by chemotherapy on healthy lungs in nude mice model [133].

5.8. Micro-emulsions Microemulsions are the monophasic, optically isotropic, thermodynamically stable, and transparent colloidal dispersion made up of co-surfactant, surfactant, water and oil having size within the range of 100 nm. These are also stated as monodispersed spherical droplets of oil in water (o/w) or water in oil (w/o), which depends on the surfactant used for its preparation [134]. These nanocarriers are found to be thermodynamically stable and have the ability to enclose both hydrophobic and hydrophilic drugs. Additionally, this nanocarrier system provides more bioavailability as well as solubility to drug and has high permeable power than other nanocarrier system [135]. These properties are key features of these nanocarrier systems which makes it ideal candidate for targeted drug delivery. With the advancement in the nanomedicine, the group of researchers formulated micro-emulsion nanocarrier system for improving the bioavailability and solubility of fenofibrate for asthma (L. [136]). The comprehended previous literature reported about the optimization of simvastatin-loaded microemulsion and their evaluation in in vivo system was also done. The result showed the better efficacy of this system in targeted delivery of insoluble drug via oral route for COPD [137]. Another study reported about the

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encapsulation of Isoniazid, Pyrazinamide and Rifampicin in varied combinations in microemulsions and evaluated for its efficacy in in vitro release for treating MDR-TB [138]. Moreover, Ganoderma lucidum-derived polysaccharides (GLP) coated oil-based microemulsion has been developed and evaluated for stability of this nanocarrier in A549-bearing xenograft mice for targeted lung cancer therapy [139].

5.9. Carbon Nanotubes Carbon nanotubes (CNs), an another nanocarrier system having tubular morphology mainly composed of carbon with diameter ranging from 4 to 100 nm. Moreover, their shape and size can be amended by changing the graphene molecules arrangement [140]. Basically, these CNs are not readily soluble in any aqueous or organic solvent. But major challenge associated with this nanocarrier system is its toxicity which demands for the solution. Progression in the technology has allowed us to chemically modify so that their biocompatibility can be enhanced, toxicity could be decreased and make it water-soluble nanocarrier system [141]. The large surface area of the CNs allows it load high amount of therapeutic agent. Additionally the unique electron emission, mechanical and optical properties makes it an effective nanocarrier system. Furthermore, this nanocarrier system has penetration power as it resembles with fine-needle and conjugation of functional groups on surface adds additional benefit of targeting the disease cell [142]. There have been studies that have used multiwalled CNs which augments the pulmonary eosinophilic inflammation and triggers the other responses by synthesizing cysteinyl leukotriene. These finding prompted to explore the pharmacological agents which can cease the leukotriene synthesis in asthmatic patients [143]. Still, the expedition is going among the researchers to load the drug for targeted delivery in COPD patients. Due to the blood circulation, the lower availability of anti-TB drug at targeted site has raised the concern for targeted delivery of drug. To resolve this issue, the researchers have developed isoniazid containing carbon nanotubes for effective delivery of therapeutic agent in bone TB [144]. Another study reported about paclitaxel-loaded single-walled CNs and result obtained from the study highlighted the potential of the nanocarrier system for treating lung cancer [145].

5.10. Quantum Dots Recently, development in the field of nanotechnology has enabled us to fabricate the colloidal nanoparticles having the properties similar to atom and is known

by the name “quantum dots (QDs).” These nanoparticles are unique as their surface modification improves both solubility and biocompatibility of these nanocarrier system [146]. It is considered to be effective fluorescent probe in contrast to other fluorophores (especially organic). High photo bleaching, wide absorption spectrum range and photo stability are few unique features of these QDs [147]. Mostly, QDs includes the elements from group II–IV (such as zinc sulfide, cadmiumselenide, and cadmium-telluride) to group III–V (such as indium arsenide, gallium arsenide and gallium nitride). QDs are composed of core and cap/shell-like structure which are further coated with polymer layer [148]. The cap/shell of the QDs serves as protective shield for core, as it contains metal complexes. These QDs are extensively are used for bioimaging, labeling, and targeting of biological molecules. In addition to these applications, now more avenues like drug delivery to target site are being explored for therapeutic purpose [149]. Still, the researchers have not considered quantum dots to be suitable nanocarrier system for targeted delivery in chronic respiratory disease [150]. The major concern for slow exploration is the heavy metal toxicity. But certain changes have been made and evaluated on the murine model. It is expected that in near future researchers will unveil the real potential of this nanocarrier and make them suitable for treating these chronic respiratory diseases [151]. Clinical studies of drug delivery system. Currently, various nanocarrier based drug delivery formulations are under pre-clinical and clinical trials for their approval by government agencies like European Medicines Agency (EMA), Europe and Food and Drug Administration (FDA), the United States [152, 153]. Moreover, different nanocarrier based formulations have already got the approval for treating lung cancer after their clinical trials. Doxil encapsulated in liposomes, known by the name “Abraxane” approved by FDA is commercially available in the market for treating cancer [154]. List of the nanocarrier based drug formulation in their clinical trials against these chronic respiratory diseases have been summarized in Table 1. Largely, most of the nanocarrier-based drug formulations are targeting the respiratory diseases such as COPD and lung cancer because of their high prevalence worldwide. But this also highlights that need for the development of nanocarrier based formulations for treating other respiratory diseases like asthma, TB and others [165]. As these formulations are effective and have improved the therapeutic potential of drugs, the

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TABLE 1

Enlist of nanocarrier-based drug formulation in their clinical trials against these chronic respiratory diseases. Clinical Phase

Drugs

Nanocarrier

Disease

References

Phase-I

AZD4635 combination with Durvalumab

Oral administered nanosuspension

Lung cancer

[155]

Phase-II

Albumin bound paclitaxel

Nanoparticle

Lung cancer

[156]

Phase-II

BIND-014 (Docetaxel)

Nano-suspension for injection

Lung cancer

[157]

Phase-II

DOTAP:Chol-TUSC2, Erlotinib

Nanoparticle

Lung cancer

[158]

Phase-II

Doxorbucin combined with Ifosfamide

Liposomal

Lung Cancer

[159]

Phase-II

Genexol-PM and Gemcitabine

Polymeric micelle

Lung cancer

[160]

Phase-II

Paclitaxel, Paclitaxel loaded polymeric micelles (Genexol-PM®)

Polymeric micelles

Lung cancer

[161]

Phase-II

SN-38

Liposomal

Lung Cancer

[162]

Phase-III

Irinotecan, Topotecan

Liposomal

Lung cancer

[163]

Phase-III

Paclitaxel combination with Cisplastin

Micelle for injection

Lung cancer

[161]

Phase-I

PF 06260414 solid dose formulation

Nanoparticles

COPD

[164]

substantial efforts in this direction will enable us to improve the clinical approaches for treating these CRDs [166]. Challenges associated with controlled drug delivery. Even though nanotechnology has brought the paradigm shift in drug delivery and has become successful, as evident from the commercially available products, whereas few did not meet the same achievement. To overcome the challenges, new nanomaterials are being continuously synthesized [167]. However, few challenges still exist and are demanding for the modification in synthesizing process. For this, certain modification are being done on the nanomaterials to improves physiochemical characteristics and properties such as enlarged functional surface area, prolonged circulation time in bloodstream, ability to cross the biological barrier, the safety of therapeutic agents from degradation and site-specific targeting [168]. Further, there is another challenge associated with nanomaterial is it large-scale production, as being the outmost need

after the development of laboratory and pilot technology for their commercialization. Moreover, production, and material cost, prevent the development of various nanocarrier drug delivery system up to commercial level [169, 170]. The challenges associated with scaling-up involves agglomeration, chemistry process, and very less concentration of nanomaterial, as these challenges are easy to amend at lab scale but process become complex when applied at large scale [62, 171]. Additionally, the sustainment of composition and the size of nanomaterial at commercial scale is also one of the barriers. In spite of the large number of patents for nanocarrier drug delivery techniques have been granted, but commercialization of these nanocarrier are still in its infancy stage [172, 173]. One of the reasons is that most of the research on nanocarrier drug delivery is conducted in academia but marketing of this product requires sophisticated handling and extensive research, which demands for the collaboration of pharmaceutical industries. Regrettably, various pharmaceutical companies feel risk to use nanotechnology as priorities approach as proper

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guidelines are unavailable and scaling-up is also prevailing challenge [174]. Though, it has been envisioned that with passing time, expiring of patents and market loss, the pharmaceutical companies will take shift towards the manufacturing of nanocarrier drug to compete constructively. These advancements will further demand the improvement in regulation guidelines according for pharmacokinetic and physicochemical properties of nanocarrier drug as they differ from other traditional drugs [175]. For this, EMA and FDA have taken the initiative to recognize few potential scientific and regulatory challenges. Additionally, International Organization for Standardization has established technical committee, those who will deal in the field of nanotechnologies to extend standards related to their terminology, nomenclature, characterization, measurement, health benefits and safety in comparison with other standards. And, these standards are still under progress [176]. Furthermore, progress in R&D related to nanocarrier drug delivery system has raised the concern about the safety of this nanomaterial in our human body. As few of the nanomaterial synthesized are biodegradable whereas few are not, in some cases, the by-product formed also becomes the reason of concern (D. [177]). For example, material used on macroscale was found to be safe but that same material was found to be toxic, as the physicochemical characteristic of that material gets altered at the nanoscale. The safety parameter for these nanomaterials should not be only limited to their effect on patient population but should also involve complete manufacturing as well as disposal processes [178]. In addition to the above context, traditional safety measures used in pharmaceutical industries are not valid for developing and manufacturing nanomaterial. Instead, they demand for extra precautions in order to safeguard the environment from the negative impact of these nanomaterials [179]. Though the progress in the nanotechnology has reduced the development cost but only few materials are still being produced in bulk, which are also doubtful as if they will be commercial successful or expensive technology [180].

6. FUTURE DIRECTIONS In the past decades, nanomedicine has developed into fascinating area for research. In the last two decades, around 1500 patent from research in this field has been filed and out of which few of them are on edge of completion of clinical trials. As summarized in the above table, lung cancer appears to be suitable example, where

this controlled drug delivery approach has served the both diagnostic and therapeutic purposes. With the aid of different type of nanocarriers system have been developed to deliver the specific amount of therapeutic agent to the affected cell, without affecting the normal cell, show the potential of nanocarrier-drug delivery system. As it is becoming the trend and will remain in the nearby future in the arena of research. Hence the expedition concerning the consistency, drug loading capacity, releasing capacity and uniformity would serve for the future prospect in this field of research. Significant developments of different nanocarrier system for drug delivery have been comprehended in this chapter. Despite being a deep understanding of the future outlook of nano-drug delivery system and nanomedicine, its real application in the healthcare sector, even in cancer diagnosis/therapy is very restricted. It has been only two decades since the introduction of this field of science and which has brought the paradigm shift in this research area but still many central attributes remain unknown. One of the chief features of this research is to explore fundamental markers related to chronic disease, so that targeted treatment could be done without amending the normal biological processes. Eventually, this field of nanomedicine will improve our knowledge about chronic diseases up to the molecular level and identification of biomarkers will unveil new opportunities for developing new diagnosis or therapeutic methods. Therefore, profound knowledge about the molecular signature of chronic disease improves the application of nanomedicine in the nearby future. However, the different types of nanocarrier systems for controlled drug delivery have been comprehended, whereas further exploration in this field will open new avenues in nanomedicine. The idea of controlled release of particular drug at targeted site, the method for analyzing these events, effect of drug at cellular/tissue level, and mathematical model for predicting are still in their development stage. Various nanomedicine studies are just focusing on the formulation and type of biomaterial, which appears to be the foundation stone of nanomedicine application. Multidisciplinary research and animal trials can play a significant role in collecting valuable data for diagnosis and therapeutic potential of nanocarriers for delivering the therapeutic drug, but it demands for both resources and time. On seeing the global trend and development of more advance diagnosis method and personalized medicine, the scope of targeted nanomedicine and nanodrug delivery method looks unparallel. The extensive research is being conducted to develop nanorobots that could be used for diagnosis purposes

CHAPTER 15 Novel Controlled Release Pulmonary Drug Delivery Systems and repair mechanism with full regulation from outside the body. This is just futuristic view yet not the reality but this could be achieved in the coming time. Along with the benefits, there are potential risks associated with these nanomedicines for both environment and humans, which requires extensive research. Therefore it is foremost important to assess the acute and chronic effects of these nanocarriers on both the environment and humans. With gaining popularity of these nanomedicines, exploration for affordable nanomaterial will be a new area of research. At last, the controlled drug delivery by different nanocarriers as discussed in above section will continuously evolve along with advancement and development in nanomedicine.

7. CONCLUSION Recently, a controlled pulmonary drug delivery system has gained significant attention because of its multiple advantages. This approach has numerous advantages over the conventional drug delivery system in the respiratory system. The large surface area of nanoparticles allows the rapid absorption of the therapeutic agent in the lungs. Additionally the small size of these nanocarriers certainly serves as an improvement over conventional dosage forms. Even the nonsoluble therapeutic agents can be incorporated in different nanocarrier systems. Liposomes and SLNs have an additional advantage because of their composition. But, biodegradable nanocarriers impose challenge of exhibiting constant drug release and this mechanism is still being investigated. Various studies are being conducted to evaluate the safety of using nanocarrier drug delivery system, many of which are in their preclinical stage of drug development and some of them are now available in the market. Alongside, while overcoming the challenges associated with this approach, the development of an effective nebulization device is also important. Furthermore, development of aerosol devices is also of utmost importance in accordance with incorporated active pharmaceutical ingredient and particle type. Additionally, the study on cell line is also important to know the range of safe and toxic concentration, as it is challenging to correlate the result of in vitro studies with in vivo studies. This prompts to infer the results with great precision so that both results of in vitro and in vivo analysis can be compared. Other than the formulation of these nanocarriers, the uptake and clearance mechanism of these nanocarrier in respiratory is highly important. Still the uncovering of the mechanisms involved in the transportation of drug across the epithelium of pulmonary system is under evaluation. Fluorescent or

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radioisotope labeling are the two option which can be used to trace the mechanism of drug uptake in the cell, but further challenge of correlating the realistic disease with model is still a mystery. In conclusion, various opportunities are available for exploration in the pulmonary system for the triumph of nanocarrier system on human trials.

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CHAPTER 16

Nanoparticle Formulations and Delivery Strategies for Sustained Drug Release in the Lungs

´ S BRITO DEVOTOa • MARI´A A. TOSCANINIb • MARI´A L. CUESTASa • TOMA a ´ N A. ISLA ´ Nc • GUILLERMO R. CASTROc,d MARI´A J. LIMERES • GERMA a

University of Buenos Aires, CONICET, Institute for Research in Microbiology and Medical Parasitology (IMPaM), Buenos Aires, Argentina, bUniversity Buenos Aires, Faculty of Pharmacy and Biochemistry, Institute of Nanobiotechnology (NANOBIOTEC), Buenos Aires, Argentina, cLaboratory of Nanobiomaterials, CINDEFI, Department of Chemistry, Faculty of Exact Sciences, National University of La Plata-CONICET (CCT La Plata), La Plata, Argentina, dMax Planck Laboratory for Structural Biology, Chemistry and Molecular Biophysics of Rosario (MPLbioR, UNR-MPIbpC), Partner Laboratory of the Max Planck Institute for Biophysical Chemistry (MPIbpC, MPG), Center for Interdisciplinary Studies (CEI), National University of Rosario, Rosario, Santa Fe, Argentina

1. INTRODUCTION Inhalation therapy is one of the ancient medical approaches to treat pulmonary diseases. The inhalation of therapeutic vapors and aerosols and the smoking of herbal preparations for medicinal purposes have been used for thousands of years. It is estimated that the origins of inhalation therapy for asthma and other lung diseases dated back more than 2000 years to Ayurvedic medicine in India, although the introduction of the first pressurized metered-dose inhaler (pMDI) in 1956 opened the era of the modern pharmaceutical aerosol industry [1]. One of the first and most notable drug delivered by the respiratory tract was opium that was used for therapeutic and recreational purposes [2]. Remarkable advances in the technology of devices, aerosol formulations, and nanotechnology for pulmonary drug delivery (PDD) have occurred since then (Fig. 1). The unique anatomical and physiological features of the lungs (i.e., large surface area of the pulmonary epithelium of 100–140 m2, an extremely thin absorptive mucosal membrane of