Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine [1 ed.] 0323953514, 9780323953511

Table of contents :
Contents
Contributors
1. Microbial, animal, and biotechnologically originatedpolysaccharides • Moumita Das Kirtania, Nancy Kahali, Anwesha Barik, Sanjay Dey and Ranjan Kirtania
2. Glycogen-based hydrogels • Bijaya Ghosh and Tapan Kumar Giri
3. Hydrogel based on hyaluronic acid • Roberta Cassano, Federica Curcio, Roberta Sole and Sonia Trombino
4. Hydrogels based on chitosan • Sujit Kumar Debnath, Monalisha Debnath, Rohit Srivastava and Abdelwahab Omri
5. Hydrogels based on heparinand its conjugates • Hemant Ramachandra Badwaik, Kalyani Sakure and Tapan Kumar Giri
6. Xanthan gum and its compositebasedhydrogels • Kaushik Mukherjee, Pallobi Dutta, Hemant Ramachandra Badwaik and Tapan Kumar Giri
7. Gellan gumebased hydrogels • Kaushik Mukherjee, Pallobi Dutta, Hemant Ramachandra Badwaik and Tapan Kumar Giri
8. Hydrogels based on dextran • Pankaj V. Dangre, Vishal C. Gurumukhi, Satish S. Meshram and Sankalp M. Zade
9. Hydrogels based on scleroglucan • Tapan Kumar Giri
10. Pullulan-based hydrogels • Anca Giorgiana Grigoras
11. Hydrogels based on levan • Álvaro González-Garcinuño, Antonio Tabernero and Eva M. Martín del Valle
12. Hydrogels based on schizophyllan • Yachen Hou and Jingan Li
13. Curdlan based hydrogels • Natasha Aquinas, Ramananda Bhat M and Subbalaxmi Selvaraj
14. Chitosan nanogel for drug delivery and regenerative medicine • Neslihan Kayra and Ali Özhan Aytekin
15. Heparin-based nanocomposite hydrogels • Amrita Thakur, Vinay Sagar Verma, Jyoti Ahirwar, Sandeep Kumar Sonkar and Hemant Ramachandra Badwaik
16. Hydrogels based on chondroitin sulfate nanocomposites • Leena Kumari, Kalyani Sakure and Hemant Ramachandra Badwaik
17. In situ gel based on gellan gum • Jieyu Zhu, Yijun Pan, Haizhou Peng, Jinzhang Fang, Guoxin Du, Akshaya Tatke and Bo Liang
18. Preclinical and clinical study of polysaccharide-based hydrogels • Bijaya Ghosh, Moumita Das Kirtania and Ranjan Kirtania
Index

Citation preview

Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine

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Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine Edited by Tapan Kumar Giri Department of Pharmaceutical Technology, Jadavpur University, Kolkata, West Bengal, India

Bijaya Ghosh NSHM College of Pharmaceutical Technology, NSHM Knowledge Campus, Kolkata Group of Institutions, Kolkata, West Bengal, India

Hemant Badwaik Shri Shankaracharya Institute of Pharmaceutical Sciences and Research, Junwani, Bhilai, Chhattisgarh, India

Elsevier Radarweg 29, PO Box 211, 1000 AE Amsterdam, Netherlands The Boulevard, Langford Lane, Kidlington, Oxford OX5 1GB, United Kingdom 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States Copyright © 2024 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. ISBN: 978-0-323-95351-1 For information on all Elsevier publications visit our website at https://www.elsevier.com/books-and-journals Publisher: Stacy Masucci Acquisitions Editor: Patricia M. Osborn Editorial Project Manager: Matthew Mapes Production Project Manager: Jayadivya Saiprasad Cover Designer: Vicky Pearson Esser Typeset by TNQ Technologies

Contents Contributors

1.

xi

Tissue engineering applications 2.5.1 Bone tissue regeneration 2.5.2 Wound healing and skin tissue regeneration 2.5.3 Self-healing hydrogel in tissue engineering 2.6 Conclusions References

Microbial, animal, and biotechnologically originated polysaccharides Moumita Das Kirtania, Nancy Kahali, Anwesha Barik, Sanjay Dey and Ranjan Kirtania 1.1

Introduction 1.1.1 Hydrogel-forming ability of polysaccharides 1.2 Natural polysaccharides: Their sources and applications 1.2.1 Polysaccharides derived from animal sources 1.2.2 Microbial and biotechnologically derived polysaccharides 1.3 Preparation of polysaccharides through biotechnological approach 1.4 Future prospect of the polysaccharide hydrogels in drug delivery and regenerative medicine 1.5 Conclusion Acknowledgments References

2.

2.5

3. 1

2.3 2.4

Introduction Glycogen 2.2.1 Source 2.2.2 Chemical structure and composition 2.2.3 Physicochemical properties 2.2.4 Hydrogel-forming ability Drug delivery applications Tumor targeting

Introduction Dermatological applications Ophthalmic applications Hyaluronic acid injectable hydrogels 3.5 Inhalable hyaluronic acid hydrogels 3.6 Hyaluronic acid hydrogels and their applications in tissue engineering 3.7 Cartilage and bone regeneration 3.8 Wounds treatment 3.9 Hyaluronic acid hydrogel and neuroregeneration 3.10 Hyaluronic acid hydrogel and stem cells 3.11 Conclusions References

2 2

7 14

16 16 16 17

4. 21 21 21 22 22 22 23 24

27 30 31

Hydrogel based on hyaluronic acid

3.1 3.2 3.3 3.4

Glycogen-based hydrogels 2.1 2.2

25

Roberta Cassano, Federica Curcio, Roberta Sole and Sonia Trombino

2

Bijaya Ghosh and Tapan Kumar Giri

24 24

35 36 37 38 39 40 41 42 42 43 44 44

Hydrogels based on chitosan Sujit Kumar Debnath, Monalisha Debnath, Rohit Srivastava and Abdelwahab Omri 4.1 4.2

4.3

Introduction Classification of hydrogel 4.2.1 Crosslinking 4.2.2 Electric charge 4.2.3 Crosslinked junctions Hydrogel preparation using chitosan 4.3.1 Chemical crosslinking

47 47 48 48 48 49 49 v

vi

Contents

4.3.2 Physical crosslinking 4.3.3 Smart hydrogel 4.4 Application of chitosan-based hydrogel 4.4.1 Drug delivery 4.4.2 Tissue engineering 4.4.3 Commercially available chitosan-based hydrogel 4.5 Conclusion and prospects References

5.

6.2.1 6.2.2

Source 89 Chemical structure and composition 90 6.2.3 Production 91 6.2.4 Physicochemical and rheological properties 92 6.3 Xanthan gum-based hydrogel in drug delivery applications 92 6.3.1 Oral controlled release drug delivery 92 6.3.2 Ophthalmic drug delivery 95 6.3.3 Vaginal drug delivery 95 6.4 Xanthan gum-based hydrogel in regenerative medicine 96 6.4.1 Xanthan gum hydrogel for bone tissue regeneration 97 6.4.2 Xanthan gum hydrogel for cartilage tissue regeneration 99 6.4.3 Xanthan gum hydrogel for skin tissue regeneration 102 6.4.4 Xanthan gum hydrogel for neuron tissue regeneration 104 6.5 Conclusion 105 References 105

51 54 59 59 62 63 64 65

Hydrogels based on heparin and its conjugates Hemant Ramachandra Badwaik, Kalyani Sakure and Tapan Kumar Giri Introduction 69 HEP-based HGLs 69 5.2.1 Chemically crosslinked HEP HGLs 69 5.2.2 HEP-based HGLs formed via Michael-type addition reactions 71 5.2.3 Amide coupling for crosslinking of HGLs 72 5.2.4 Enzyme-mediated crosslinked HGLs 74 5.2.5 Other covalent bonding approaches 74 5.3 Physically crosslinked HEP HGLs 76 5.3.1 Electrostatic interaction HGLs 76 5.3.2 Hosteguest interaction HGLs 77 5.3.3 Hydrogen bonding (growth factor-crosslinked HGLs) 78 5.3.4 Hydrophobic interaction HGLs 78 5.3.5 Other physical interactions 78 5.4 Combined interaction (duel crosslinked) HGLs 80 5.5 Smart HGLs or stimuli-responsive HEP HGLs 81 5.5.1 Enzymatically responsive HGLs 83 5.5.2 Glutathione-responsive HGLs 83 5.6 Conclusion and prospective 84 References 84 5.1 5.2

6.

Xanthan gum and its compositebased hydrogels Kaushik Mukherjee, Pallobi Dutta, Hemant Ramachandra Badwaik and Tapan Kumar Giri 6.1 6.2

Introduction Xanthan gum

89 89

7.

Gellan gumebased hydrogels Kaushik Mukherjee, Pallobi Dutta, Hemant Ramachandra Badwaik and Tapan Kumar Giri 7.1

Introduction 7.1.1 Source 7.1.2 Chemical structure 7.1.3 Types 7.1.4 Properties 7.1.5 Gelling properties 7.2 Gellan gumebased hydrogel in drug delivery 7.2.1 Oral drug delivery 7.2.2 Ocular drug delivery 7.2.3 Nasal drug delivery 7.3 Gellan gumebased hydrogel in regenerative medicine 7.3.1 Gellan gum hydrogel for bone tissue regeneration 7.3.2 Gellan gum hydrogel for skin tissue regeneration 7.3.3 Gellan gum hydrogel for cartilage tissue regeneration 7.3.4 Gellan gum hydrogel for neuron tissue regeneration 7.4 Conclusion References

109 109 110 110 110 111 111 111 114 114 115 116 119 120 122 124 124

Contents

8.

Hydrogels based on dextran

10.2.6 10.2.7

Drugs for diabetic disease 161 Drugs for neurodegenerative diseases and other neurological conditions 162 10.2.8 Vaccine formulations 163 10.2.9 Therapeutic proteins 164 10.3 Regenerative medicine 165 10.3.1 Hydrogels for burn injuries 165 10.3.2 Hydrogels with antiadhesion and tissue regeneration properties 166 10.3.3 Hydrogels for cartilage tissue engineering 167 10.3.4 Hydrogels for dentin tissue engineering 167 10.4 Controlled and sustained release of drugs in regenerative medicine 168 10.5 Conclusions and perspective remarks 172 Acknowledgments 172 References 172

Pankaj V. Dangre, Vishal C. Gurumukhi, Satish S. Meshram and Sankalp M. Zade 8.1 8.2

Introduction to hydrogel Dextran 8.2.1 Introduction 8.2.2 Structure of DEX 8.2.3 Source of DEX 8.2.4 Properties of DEX 8.3 Gelation 8.4 Methods of preparations 8.4.1 Physical methods 8.4.2 Chemical methods 8.5 Biomedical application 8.5.1 Tissue engineering applications 8.5.2 Drug delivery applications 8.6 Conclusion References

9.

129 129 129 130 130 131 131 131 131 132 133 133 134 135 135

Hydrogels based on scleroglucan

11.

Tapan Kumar Giri 9.1 9.2

Introduction Scleroglucan 9.2.1 Source 9.2.2 Composition and chemical structure 9.2.3 Physicochemical properties 9.3 Drug delivery applications of scleroglucan 9.3.1 Oral sustained drug delivery 9.3.2 Colon-targeted drug delivery 9.3.3 Topical drug delivery 9.4 Tissue engineering applications of scleroglucan hydrogel 9.5 Conclusion References

10.

139 139 139

Introduction 175 11.1.1 Levan structure and mechanism of formation 175 11.1.2 Levan properties and associated potential applications 176 11.2 Production mechanisms 177 11.2.1 Microbial production 177 11.2.2 Cell-free synthesis 178 11.2.3 Levan purification 178 11.3 Levan-based hydrogels 179 11.3.1 Rheology of levan solutions 179 11.3.2 Type of hydrogels: A brief summary of classification and preparation techniques 180 11.3.3 Levan-based hydrogels 181 11.4 Conclusions and future perspectives 181 References 184 11.1

140 140 140 140 145 145 147 147 147

Anca Giorgiana Grigoras Introduction Drug delivery systems 10.2.1 Antiinflammatory drugs 10.2.2 Antimicrobial drugs 10.2.3 Drugs for imagistic diagnostic 10.2.4 Drugs for wound healing 10.2.5 Drugs for cancer therapy

151 152 152 153 158 158 159

Hydrogels based on levan A´lvaro Gonza´lez-Garcinun˜o, Antonio Tabernero and Eva M. Martı´n del Valle

Pullulan-based hydrogels 10.1 10.2

vii

12.

Hydrogels based on schizophyllan Yachen Hou and Jingan Li 12.1

Introduction 12.1.1 Preparation and structure

187 187

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Contents

12.1.2 Gelation behavior Biological function 12.2.1 Antineoplastic activity 12.2.2 Gene delivery 12.2.3 Immunoadjuvants 12.2.4 Drug delivery 12.2.5 Wound dressing 12.2.6 Other function of SPG 12.3 Conclusion Acknowledgments References Further reading 12.2

13.

14.2.4

Chitosan nanogels with other polymers 221 14.2.5 Stimuli-responsive chitosan nanogels 221 14.2.6 Chitosan nanogels with different delivery approaches 225 14.3 Chitosan nanogels in regenerative medicine 226 14.3.1 Chitosan nanogels applications in wound healing 226 14.3.2 Chitosan nanogels applications in dentistry 227 14.4 Characterization methods of chitosan nanogels in drug delivery 228 14.5 Conclusion 229 References 229

188 189 189 191 195 195 196 197 197 197 197 202

Curdlan based hydrogels Natasha Aquinas, Ramananda Bhat M and Subbalaxmi Selvaraj 13.1 13.2 13.3

Overview Biosynthesis of curdlan Polymer properties with structural features related to hydrogel formation 13.4 Importance of curdlan hydrogel in drug delivery 13.5 Method of preparation of curdlan hydrogel 13.5.1 Carboxymethylation 13.5.2 Phosphorylation 13.5.3 Sulfation 13.6 Evaluation parameter in animal model for curdlan hydrogel 13.7 Application of curdlan-based hydrogel 13.7.1 Drug delivery application 13.7.2 Tissue engineering application 13.8 Conclusion References

14.

203 203

205 205 206 206 207 207 208 209 209 211 211 212

Chitosan nanogel for drug delivery and regenerative medicine Neslihan Kayra and Ali O¨zhan Aytekin 14.1

14.2

Introduction 14.1.1 Chitosan 14.1.2 Chitosan nanogels 14.1.3 Chitosan nanogel preparation methods Chitosan nanogels in drug delivery 14.2.1 Chitosan-based nanogels 14.2.2 Chitosan derivative-based nanogels 14.2.3 Functionalized chitosanbased nanogels

215 215 216 216 217 218 218 220

15.

Heparin-based nanocomposite hydrogels Amrita Thakur, Vinay Sagar Verma, Jyoti Ahirwar, Sandeep Kumar Sonkar and Hemant Ramachandra Badwaik Introduction 233 Advantages of nanocomposite hydrogel over hydrogel 234 15.2.1 Mechanical 235 15.2.2 Electrical 235 15.2.3 Biological 235 15.2.4 Magnetic 235 15.3 Types of nanocomposite hydrogels 235 15.4 Heparin 235 15.4.1 Heparin’s role in biological systems 236 15.5 Preparation of heparin-based nanocomposites 237 15.5.1 Electrospinning 237 15.5.2 Layer-by-layer assembly 238 15.5.3 Covalent functionalization 238 15.5.4 Cosynthesis method 238 15.5.5 Physical self-assembly method 238 15.5.6 Spontaneous emulsion solvent diffusion method 239 15.5.7 Coprecipitation and solvothermal method 239 15.5.8 Other methods 239 15.6 Application of heparin-based nanocomposites 239 15.6.1 Drug delivery 239 15.6.2 Tissue engineering 240 15.7 Conclusion and prospective 245 References 245 15.1 15.2

Contents

16.

Hydrogels based on chondroitin sulfate nanocomposites

Ophthalmic in situ gelling drug delivery 17.4.2 Intranasal in situ gelling drug delivery 17.4.3 Oral in situ gelling drug delivery 17.4.4 Injectable in situ gelling drug delivery 17.4.5 Other in situ gelling drug delivery routes 17.5 Conclusion References

17.4.1

Leena Kumari, Kalyani Sakure and Hemant Ramachandra Badwaik 16.1 16.2

Introduction Structure and bioactivities of chondroitin sulfate 16.3 Method of preparation of nanocomposite hydrogel systems 16.3.1 Crosslinking method 16.3.2 In situ method 16.3.3 Inwards diffusion method 16.3.4 Codependency and crosslinking method 16.4 Drug delivery, biomedical, and other applications of chondroitin sulfate nanocomposite hydrogel 16.4.1 Drug delivery applications 16.4.2 Tissue engineering 16.4.3 Wound dressing 16.4.4 Other applications 16.5 Future directions and conclusion References Further reading

17.

249 250 251 251 251 252

18.

252

252 252 253 255 256 257 257 259

17.3

17.4

266 266 267 267 267 267

Preclinical and clinical study of polysaccharide-based hydrogels

18.1 18.2 18.3

Introduction Market trend Hydrogels in drug delivery system 18.3.1 Hydrogel products used in the diseases of oral cavity 18.3.2 Hydrogels in oral delivery system 18.3.3 Hydrogels in controlled release dosage form 18.3.4 Injectable hydrogels 18.3.5 Hydrogels used in the diseases of the skin 18.3.6 Hydrogels in the diseases of the eye 18.4 Tissue engineering 18.5 Wound repair 18.6 Conclusion References Further reading

Jieyu Zhu, Yijun Pan, Haizhou Peng, Jinzhang Fang, Guoxin Du, Akshaya Tatke and Bo Liang Introduction Mechanisms of gellan gum-based in situ gels Compatibility of gellan gum in different formulations 17.3.1 Gellan gum used in aqueous-based formulations 17.3.2 Gellan gum used in lipid-based formulations Applications of gellan gum as in situ gels

266

Bijaya Ghosh, Moumita Das Kirtania and Ranjan Kirtania

In situ gel based on gellan gum

17.1 17.2

ix

261 261 262

262

273 274 274

275 276 276 277 280 281 283 288 289 289 292

263 263

Index

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Contributors Jyoti Ahirwar, Daksh Institute of Pharmaceutical Sciences, Chhatarpur, Madhya Pradesh, India Natasha Aquinas, Department of Biotechnology, Manipal Institute of Technology, Manipal Academy of Higher Education (MAHE), Manipal, Karnataka, India Ali Özhan Aytekin, Genetics and Bioengineering Department, Engineering Faculty, Yeditepe University, Istanbul, Turkey Hemant Ramachandra Badwaik, Shri Shankaracharya Institute of Pharmaceutical Sciences and Research, Bhilai, Chhattisgarh, India

Bijaya Ghosh, NSHM College of Pharmaceutical Technology, NSHM Knowledge Campus, Kolkata Group of Institutions, Kolkata, West Bengal, India Tapan Kumar Giri, Department of Pharmaceutical Technology, Jadavpur University, Kolkata, West Bengal, India Álvaro González-Garcinuño, Department of Chemical Engineering, University of Salamanca, Salamanca, Spain Anca Giorgiana Grigoras, "Petru Poni" Institute of Macromolecular Chemistry, Iasi, Romania

Anwesha Barik, Department of Pharmaceutical Technology, School of Medical Sciences, Adamas University, Kolkata, West Bengal, India

Vishal C. Gurumukhi, Department of Quality Assurance, Shreeyash Institute of Pharmaceutical Education and Research, Aurangabad, Maharashtra, India

Ramananda Bhat M, Department of Biotechnology, Manipal Institute of Technology, Manipal Academy of Higher Education (MAHE), Manipal, Karnataka, India

Yachen Hou, First Affiliated Hospital of Zhengzhou University, Zhengzhou, China

Roberta Cassano, Department of Pharmacy, Health and Nutritional Science, University of Calabria, Arcavacata, Italy Federica Curcio, Department of Pharmacy, Health and Nutritional Science, University of Calabria, Arcavacata, Italy Pankaj V. Dangre, Department of Pharmaceutics, Datta Meghe College of Pharmacy, DMIHER (DU), Wardha, Maharashtra, India Sujit Kumar Debnath, Department of Biosciences and Bioengineering, Indian Institute of Technology Bombay, Mumbai, Maharashtra, India Monalisha Debnath, Department of Biosciences and Bioengineering, Indian Institute of Technology Bombay, Mumbai, Maharashtra, India Sanjay Dey, DmbH Institute of Medical Science, Dadpur, West Bengal, India Guoxin Du, iVIEW Therapeutics (Zhuhai) Co. Ltd., Zhuhai, Guangdong, China Pallobi Dutta, Department of Pharmaceutical Technology, Jadavpur University, Kolkata, West Bengal, India Jinzhang Fang, iVIEW Therapeutics (Zhuhai) Co. Ltd., Zhuhai, Guangdong, China

Nancy Kahali, School of Pharmacy, Sister Nivedita University, Kolkata, West Bengal, India Neslihan Kayra, Genetics and Bioengineering Department, Engineering Faculty, Yeditepe University, Istanbul, Turkey Moumita Das Kirtania, School of Pharmacy, Techno India University, Kolkata, West Bengal, India Ranjan Kirtania, School of Pharmacy, Techno India University, Kolkata, West Bengal, India Leena Kumari, School of Pharmacy, Techno India University, Kolkata, West Bengal, India Jingan Li, School of Materials Science and Engineering & Henan Key Laboratory of Advanced Magnesium Alloy & Key Laboratory of Materials Processing and Mold Technology (Ministry of Education), Zhengzhou University, Zhengzhou, China Bo Liang, iVIEW Therapeutics Inc., Cranbury, NJ, United States; iVIEW Therapeutics (Zhuhai) Co. Ltd., Zhuhai, Guangdong, China Eva M. Martín del Valle, Department of Chemical Engineering, University of Salamanca, Salamanca, Spain; Institute of Biomedicine of Salamanca, IBSAL, Salamanca, Spain

xi

xii Contributors

Satish S. Meshram, Department of Pharmacognosy, University Department of Pharmaceutical Science, RTMNU, Nagpur, Maharashtra, India

Rohit Srivastava, Department of Biosciences and Bioengineering, Indian Institute of Technology Bombay, Mumbai, Maharashtra, India

Kaushik Mukherjee, Department of Pharmaceutical Technology, Jadavpur University, Kolkata, West Bengal, India

Antonio Tabernero, Department of Chemical Engineering, University of Salamanca, Salamanca, Spain

Abdelwahab Omri, The Novel Drug & Vaccine Delivery Systems Facility, Department of Chemistry and Biochemistry, Laurentian University, Sudbury, ON, Canada Yijun Pan, iVIEW Therapeutics Inc., Cranbury, NJ, United States Haizhou Peng, iVIEW Therapeutics (Zhuhai) Co. Ltd., Zhuhai, Guangdong, China Kalyani Sakure, Rungta College of Pharmaceutical Sciences and Research, Bhilai, Chhattisgarh, India Subbalaxmi Selvaraj, Department of Biotechnology, Manipal Institute of Technology, Manipal Academy of Higher Education (MAHE), Manipal, Karnataka, India Roberta Sole, Department of Pharmacy, Health and Nutritional Science, University of Calabria, Arcavacata, Italy Sandeep Kumar Sonkar, Rungta College of Pharmaceutical Sciences and Research, Raipur, Chhattisgarh, India

Akshaya Tatke, iVIEW Therapeutics Inc., Cranbury, NJ, United States Amrita Thakur, Department of Pharmaceutics, School of Pharmacy, Vishwakarma University, Pune, Maharashtra, India Sonia Trombino, Department of Pharmacy, Health and Nutritional Science, University of Calabria, Arcavacata, Italy Vinay Sagar Verma, Shri Shankaracharya Technical Campus, Faculty of Pharmaceutical Sciences, Bhilai, Chhattisgarh, India Sankalp M. Zade, Department of Pharmaceutics, Datta Meghe College of Pharmacy, DMIHER (DU), Wardha, Maharashtra, India Jieyu Zhu, iVIEW Therapeutics Inc., Cranbury, NJ, United States

Chapter 1

Microbial, animal, and biotechnologically originated polysaccharides Moumita Das Kirtania1, Nancy Kahali2, Anwesha Barik3, Sanjay Dey4 and Ranjan Kirtania1 1

School of Pharmacy, Techno India University, Kolkata, West Bengal, India; 2School of Pharmacy, Sister Nivedita University, Kolkata, West Bengal,

India; 3Department of Pharmaceutical Technology, School of Medical Sciences, Adamas University, Kolkata, West Bengal, India; 4DmbH Institute of Medical Science, Dadpur, West Bengal, India

1.1 Introduction Polysaccharides obtained from natural resources consist of monosaccharide units (3e9 carbon atoms) which are linked to each other by covalent bonds and glucosidic links [1]. They are found in nature abundantly and are having several beneficial properties like biodegradability, biocompatibility, and low immunological activity. The polysaccharides can be modified by methods using chemicals or enzymes and are utilized in several drug delivery devices, especially for targeted drug delivery due to favorable physicochemical properties; for example, the formation of gels or polymeric networks or formation of pH-sensitive devices. They can form complex with various proteins or similar macromolecules due to the availability of a range of functional groups [2]. The polysaccharides may be obtained from various natural resources as given below: l l l l

Plant-derived polysaccharides including starch or cellulose Algae-derived polysaccharides like alginate and carrageenan Animal-derived polysaccharides, for example, hyaluronic acid (HA), chitin, and chitosan Microbial organisms-derived polysaccharides like gellan gum (GG), xanthan gum, and pullulan [3] (Fig. 1.1).

The structures of different polysaccharides are described in figures [4e16]. Each of these categories have been extensively studied by researchers for various applications, may it be the pharmaceutical industry or in cosmetics, or in food industry.

FIGURE 1.1 Classification of polysaccharides.

Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine. https://doi.org/10.1016/B978-0-323-95351-1.00008-9 Copyright © 2024 Elsevier Inc. All rights reserved.

1

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Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine

1.1.1 Hydrogel-forming ability of polysaccharides The polysaccharides have been used to prepare hydrogels which can entrap medicinal agents for their controlled delivery. In many of such cases, the polymers have been cross-linked with different cross-linking agents for rigidizing their internal structure and controlling the release of drugs. The polysaccharides when cross-linked extensively can hold huge amounts of water-forming hydrogels due to their macrostructure. Further due to their high molecular weight, they are used to increase the viscosity of solutions or suspensions and other liquid dosage forms to improve their stability. Hydrogels based on HA and chondroitin sulfate (CS) have been used for tissue engineering purposes as well as in wound healing. As such their applications are huge.

1.2 Natural polysaccharides: Their sources and applications The next section would highlight and elaborate on the physicochemical characteristics, sources, structure, and application in drug delivery, especially the formation of hydrogels by various natural polysaccharides.

1.2.1 Polysaccharides derived from animal sources The polysaccharides described in the next section are obtained from various animal sources.

1.2.1.1 Glycogen The polymer belongs to the class of glucose polysaccharides and is commonly found in animal cells including mammals and non mammals, few plant species, and various kinds of microorganisms. It is also known as the most important source for glucose which can be used to generate glucose quickly. For vertebrates, liver is the main storehouse for the polysaccharide to be supplied to other tissues whenever required based on the blood glucose levels. It is also found in fat and muscle tissues where through metabolism glucose-6 phosphate is produced responsible for energy supply in cells. However, the production and breakdown of glycogen depends on the action of various enzymes and hormones and thus may vary from one person to another according to their lifestyles including food habits, physical activity, and stress levels. The beta particles of the polymer are composed of chains of alpha-D-glucose residues which are linked to each other by alpha-d bonds and contain branching at regular intervals due to which the structure appears bushy. The structure of such particles adjoined on a single protein is called glycogenin. Several such units link to form the alpha-particle ranging in diameter from 100 to 150 nm. The polysaccharide can be thus be occupied within the cells in huge amounts being osmotically inactive due to its structure. Further due to extensive branching, the solubility is high and yields sugar residues quickly due to high numbers of non reducing ends in its structure through which a number of reactions can be catalyzed by enzymes at a time [17]. The polymer has high amounts of branching due to which the solubility is higher too. It is also called animal starch. The polymer is so branched and compact that the molecular weight is around 108 Da and is equivalent to 60,000 glucose units. The supercoiled polymer chains are responsible for the compact structure of the polysaccharide allowing loads of carbon energy entrapped within small volume [18]. Due to its branched structure and multiple sites for bond formation, the polymer has been used for hydrogel formation. For example, collagen and nanohydroxyapatite have been cross-linked using glycogen [19] (Fig. 1.2).

1.2.1.2 Hyaluronic acid The polysaccharide with a high molecular weight has been found in various sources like vitreous humor, rooster comb, umbilical cords, etc. Further also found in some groups of streptococci from which the polymer has been extracted in huge amounts as it is abundantly available in such species [20]. The polymer has been found not only in the extracellular matrix (ECM) but also in the intracellular matrix and on surface of cells and has been known to be involved in a wide range of physiological functions. It makes up the ECM together with proteins and proteoglycans and plays a significant role in homeostasis of tissues. Therefore, the function of the polymer toward hydrating and lubricating the tissues or diffusing various solutes in the extracellular space is majorly due to its interaction with cells or components of the ECM. The polymer has been extensively utilized in various skin care products for its activity in hydrating the tissues, especially in the antiageing category. Hyaluronic acid (HA) has been observed to form bonding with receptors on the cell surface and activating signal pathways which affect regular functioning of cells, development of tissues, progression of tumors, inflammation, wound

Microbial, animal, and biotechnologically originated polysaccharides Chapter | 1

3

FIGURE 1.2 Structure of glycogen. Reprinted from Engelking LR, Glycogen. Textbook of veterinary physiological chemistry, Copyright (2003) with permission from Elsevier.

healing, etc. The polymer being biodegradable and biocompatible and being capable of modification through chemical means has been extensively utilized by scientists around the world for the development of biomaterials, tissue scaffolds, hydrogel-based systems, etc., which have huge clinical benefits [21]. The polysaccharide essentially consists of alternating chains of beta-(1,4)-glucuronic acid and beta-(1,3)-N-acetyl glucosamine which are not branched [20]. Hydrogel formation with HA has been especially of benefit due to its versatile properties being nonimmunogenic, biodegradable, and biocompatible rendering the required stiffness and morphology to the hydrogel along with added biological activity. The hydrogels synthesized with HA possess a complex structure together with anisotropy and viscoelasticity properties which can be tuned to requirements. The hydrogel particles especially in micro- or nanogels reach to the microlevel, whereas the cross-linked structure of the polymer in the hydrogel reaches the macrolevel. Further the HA-based hydrogels have been extensively researched for tissue repair and regeneration as well as for wound-healing properties [22]. The structure of the polymer has been studied extensively to modify the functional groups causing cross-linking by various techniques to result in hydrogels to be utilized for tissue engineering; for example, the Dielse Alder chemistry, photo cross-linking, and Schiff base chemistry techniques to prepare hydrogels [23] (Fig. 1.3).

1.2.1.3 Chitosan Chitin is found abundantly in nature within fungi cell walls or as basic structural component in the skeleton of insects, shrimps, crabs, and other crustaceans and thus it is known as the second most abundant polymer after cellulose [24]. Chitin obtained from insects, crustaceans, and fungi has huge amounts of acetylated groups in the structure which renders rigidity to the structure with low solubility and thus not suitable for direct application. Chitosan is obtained from chitin’s partial deacetylation and thus called deacetylated chitin. It is a linear natural polysaccharide with polycationic character. Chitosan has greater number of amino groups and solubility higher compared to chitin [25]. pKa of the polymer is 6.5 giving a weak basic character. Thus, it can dissolve in dilute acids. The polymer consists of amino and hydroxyl groups in the structure leading to hydrogen bond formation and crystalline structure. The molecular weight ranges from 50 to 2000 KD. The polymer is biodegradable in the body due to the action of enzymes in the gastrointestinal tract, lysozyme, gastric acid, and resident bacteria in the colon [25].

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FIGURE 1.3 Structure of hyaluronic acid. Reprinted from Dovedytis M, Liu ZJ, Bartlett S. Hyaluronic acid and its biomedical applications: A review. Eng Regen 2020;1:102e113, with permission from Elsevier.

The structure essentially consists of cross-linked chains of b-(1e4)-linked D-glucosamine and N-acetyl-D-glucosamine. The polymer is unique in its cationic character, whereas majority of the natural polysaccharides are negatively charged or neutral in character in acidic medium. Thus, the polymer can form complexes with such oppositely charged natural or synthetic polymers. The polymer possesses all the benefits of natural polysaccharides and is non toxic, biodegradable, biocompatible and possess antioxidant, antitumor, and antimicrobial properties which are again dependent on the extent of deacetylation and molecular weight of the polymer. Further, the polymer has found its applications in tissue engineering, wound-healing delivery systems, and in controlled drug delivery systems as a drug carrier [24]. Ionic interactions, hydrogen and hydrophobic bonding and forces exist within the polysaccharide chains in chitosan structure, and they are dependent on the ionic strength and molecular weight of chitosan. Chitosan cross-linking with suitable agents renders greater stability and stiffness to the polymer for better drug delivery properties [25]. Chitosan cross-linked with benzaldehyde was used to prepare and characterize freeze-dried hydrogel particles for the delivery of three model drugs like 5-fluorouracil, caffeine, and ascorbic acid [26]. Photo cross-linking of chitosan was described in yet another study, and the hydrogels produced were utilized for biomedical applications [27] (Fig. 1.4).

FIGURE 1.4 Structure of chitosan. Reprinted from Chatterjee R, Maity M, Hasnain MD, Nayak AM. Chitosan: source, chemistry, and properties, Chitosan in Drug Delivery, 2022;1e22, with permission from Elsevier.

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1.2.1.4 Heparin Heparin, a sulfated glycosaminoglycan, has been used as an anticoagulant since many decades and has been used for several indications. It is present within the mast cells and released into circulation only at the sites of injured tissues. There was also a proposition given that the main function of heparin is to fight against the bacteria and pathogens at those sites where the tissues are injured. The molecule also possesses the highest negative charge density [28]. A specific pentasaccharide sequence in the heparin molecule has been found to act together with antithrombin (an inhibitor of plasma coagulation factor) which renders it with anticoagulant activity. Heparin has also been found to possess antiviral, antimetastatic, or antiinflammatory properties [29]. The molecular weight of heparin ranges from 3 to 30 KDa [30]. The molecular structure has been found to consist of repetitive sulfated disaccharide units such as IdoA(2S)-GlcNS(6S), 6-O-sulfated, N-sulfated glucosamine, and 2-O-sulfated iduronic acid [31]. Heparin has been mainly collected from animal tissues of porcine mucosa and bovine lungs. However, biotechnological and chemical methods have been followed as alternatives [32]. Several applications of heparin-based hydrogels include tissue engineering, implantation, controlled delivery of drugs, or as biosensors because of their three-dimensional structure. The major challenges in using heparin were of some adverse effects like bleeding or thrombocytopenia, its supply and safety due to its animal sourcing and collection. Therefore, synthetic forms have been studied upon recently [33]. In a study, such hydrogels were prepared which would cross-link in situ and contained actives like arginine-glycine-aspartate and heparin with good mechanical properties [34]. In yet another study, heparin was combined with functionalized polyethylene glycol to form hybrid hydrogel network. The effect of the changes in the molar ratio of the two components was further observed on mechanical strength, swelling, and pore size of the polymer network [35] (Fig. 1.5).

1.2.1.5 Chondroitin sulfate Chondroitin sulfate (CS), the polysaccharide, is present in the composition of ECM associated with connective tissues like cartilage, bones, tendons, ligaments, and skin. Therefore it has been utilized in the management of autoimmune diseases like osteoarthritis, in which the modifications in metabolism and tissues of the joints are observed [36,37]. CS is a polymer that has good solubility in water. Early attempts to use CS in colon targeting of drugs as a carrier resulted in varying degrees of cross-linking of the polymer, resulting in a biodegradable system which can achieve controlled release of biologically active molecules. A linear relationship was discovered between the drug release rate and the amount of crosslinking, implying that the degree of cross-linking of the polymer can be utilized to control drug delivery kinetics [38e40]. The polysaccharide carries negative charges and is made up of repeated units of alternately arranged N-acetylhexosamine and glucuronic acid that carries a sulfur group in one of two positions, based on the source of collection. Chondroitin sulfate A is sulfated at the 4-position and comes mostly from bovine cartilage, whereas chondroitin sulfate B comes mostly from shark cartilage and is sulfated at the 6-position. As CS covers a wide range of polymers, has low UV absorbance, and further is very ionic, precise quantification and analysis are difficult [36,41]. Moreover, CS is a major component of joint ECM regulation and has been linked to the development and improvement of joint or cartilage-related diseases. Therefore, substantial work has gone into developing injectable CS hydrogels for osteochondral tissue engineering [36,42,43]. CS, like HA, possesses hydroxyl and carboxyl groups on its backbone, allowing methacrylate groups to be added. Methacrylic anhydride and glycidyl methacrylate are two typical reagents that include methacrylate [43]. Shin and his co-researchers developed CS and catechol (CA)-loaded hydrogels to find out different factors such as adhesive properties, swelling ratio, mechanical strength, and degradation properties [44]. They observed that CS-CA hydrogels markedly enhanced modulus with value w10 kPa, and better adhesiveness of w3 N as compared to conventional hydrogels prepared with CS (w100 Pa; 0.05 N). Furthermore, by establishing a cartilage-like milieu, CS-CA

FIGURE 1.5 Structure of heparin. Reprinted from Pawar R, Jadhav W, Bhusare S, Borade R, Farber S, Itzkowitz D, Domb A. Polysaccharides as carriers of bioactive agents for medical applications. Natural-Based Polymers for Biomedical Applications, 2008;3e53, with permission from Elsevier.

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hydrogels comprising decellularized tissue slices of cartilage enhanced chondrogenic development of human stem cells derived from adipose tissues. Finally, due to the expression of good biological, mechanical, and physical qualities for cartilage production, the translocation of the CS-CA hydrogels with tissue adhesiveness improved compatibility and integration with the cartilage tissues of the host and further helped in cartilage generation [44]. Another study used green chemistry concepts and natural polysaccharides to make chitosan/chondroitin sulfate (CHT/ CS)-based hydrogels as reported by Nunes and his coworkers [45]. Because of its well-known ability as a solvent for polysaccharides with high molecular mass, like CHT, ionic liquid (IL) was employed instead of water. CHT solution was prepared at concentrations of 4% w/v. In aqueous solutions, an 87 kDa CHT did not reach that level. At any pH range studied, hydrogels based on CHT/CS were found to be stable and possessed good swelling properties, which were proven primarily through scanning electron microscopy (SEM) studies and release studies. The hydrogels produced in an aqueous solution (0.57 mol/L) were neither stable nor swelled freely. The results of cell viability tests showed that hydrogels made from IL had no toxicity. These discoveries pave the way for the development of novel green chemistry-based technologies for creating biopolymer-based hydrogels [46] (Fig. 1.6). CS has been used recently for the production of scaffolds for cartilage regeneration by tissue engineering. Further the polymer is biocompatible and non toxic. Further the tensile strength, stability under in vivo conditions, and nonimmunogenic characteristics go in favor of the polymer. However, the major limitation holds in the degradation of the scaffolds at a high rate due to which the therapeutic efficacy suffers. Therefore in a study, the CS was used together with a highly branched PEG copolymer (HB-PEG) with multifunctional groups. The combination greatly affects the therapeutic efficacy in terms of cartilage regeneration and repair. The scaffolds containing CS-SH (CS with thiol groups) and HB-PEG demonstrated the formation of gels at a rapid rate, excellent tensile strength, and degradation at a slower rate. The efficacy of the hydrogels was confirmed by the studies conducted on stem cells derived from adipose tissues of rats and shown to reduce inflammation and increase chondrogenesis [47]. Significant advantages of the hydrogel-based scaffolds used for tissue engineering include the tensile strength and mechanical properties, the rate of its degradation, capacity for mass transport, etc., which is mainly due to the threedimensional interpenetrating polymeric network (IPN) of the hydrogels. Hydrogel-based scaffolds resemble the structure of the cartilage’s ECM and thus are much more beneficial than conventional gels. In a study, hydrogel-based scaffolds were produced for utilization in tissue engineering using two simultaneous approaches. The hydrogels were prepared with collagen, HA, and chondroitin sulfate. The method of preparation included the self-assembly of collagen molecules together with polymerization of methacrylate derivatives of chondroitin sulfate and HA by cross-linking technique. The rate of degradation, the swelling characteristics, or the modulus of compression was observed to be dependent on the extent of the methacrylation of chondroitin sulfate. The IPN hydrogels were found to be cytocompatible when studied using a culture of chondrocytes collected from rabbit. The hydrogels were also observed to be able to regulate cartilage gene

FIGURE 1.6 Swelling of CH-CS hydrogels. Reprinted from Nunes CS, Rufato KB, Souza PR, de Almeida EAMS, da Silva MJV, Scariot DB, Nakamura CV, Rosa FA, Martins AF, Muniz EC. Chitosan/chondroitin sulfate hydrogels prepared in [Hmim][HSO4] ionic liquid. Carbohydr Polym 2017;170:99e106, with permission from Elsevier.

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FIGURE 1.7 Structure of chondroitin sulfate.

expression and improve the secretion of collagen II and glycosaminoglycan released from chondrocytes. Thus the IPN hydrogels were proven to be promising tools for tissue engineering purposes [48] (Fig. 1.7).

1.2.2 Microbial and biotechnologically derived polysaccharides The polysaccharides described in the next section are derived from various cellular structures of microorganisms or produced through biotechnological procedures.

1.2.2.1 Xanthan gum The polymer is produced by fermentation from the bacterial species Xanthomonas campestris which releases it as an extracellular polysaccharide. The polymer exhibits good solubility in cold aqueous solutions which can be stabilized with it. Further the aqueous solutions express pseudoplastic flow properties which are enhanced with galactomannans like guar gum and glucomannans. Xanthan and guar gums at percentage of 25%e30% and 70%e75%, respectively, are regarded to exhibit the best synergy in between them [49]. The polymer has excellent flow characteristics and retains its effectiveness at varying temperatures, pH (1e13), or ionic strength. Due to its several beneficial properties, the natural gum finds huge applications in pharmaceutical as well as the food industry, especially in bakery, salad dressings, dairy products as well as in cosmetics. Concerning drug delivery applications, the gum has been utilized in nasal drug delivery, wound healing, drug delivery in the brain tissues, tissue engineering, and in various liquid and solid oral dosage forms, and it can stabilize dispersions avoiding the settling of the dispersed phase for prolonged periods of time [50]. The polymer basically consists of units of D-glucose and possesses a side chain of trisaccharide containing guluronic acid in between two mannose moieties. The viscous aqueous solution of the polymer when solubilized in water is due to the presence of anionic carboxyl groups in the side chain. The viscosity is however temperature-dependent and reduces with rise in temperature and becomes insignificant at >65
C. However on reducing the temperature, the viscosity is restored [51]. Due to such beneficial physicochemical characteristics, xanthan gum has been extensively researched for the preparation of hydrogels. In such a study, xanthan gum was cross-linked with starch to yield acrylic acid grafted hydrogels using irradiation in microwave which were promising for controlled drug delivery. The swelling of the hydrogels was found to be higher than the hydrogels prepared with xanthan gum alone [52]. In yet another research study, aligned hydrogels were prepared which contained pores in an aligned structure using xanthan gum by irradiation with photo initiated free radicals. The template used contained the crystals of sodium acetate [53] (Fig. 1.8). FIGURE 1.8 Structure of xanthan gum.

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1.2.2.2 Gellan gum Gellan gum (GG) is an extracellular polysaccharide, anionic in nature and made up of repetitive units of roughly 60% glucose, 20% glucuronic acid, 20% rhamnose, as well as acetate, glycerate, and two acyl groups, bonded to the glucose residue proximal to the glucuronic acid. GG can form gels which undergoes sol to gel transformation under different temperature conditions with high gel strength and outstanding stability. GG produces translucent, flexible, elastic, and soft gels in its high acyl structure, whereas the low acyl structure produces brittle gels. These characteristics have sparked the development of hydrogels to be delivered in the injectable form which can be used in regenerating as well as repairing the cartilage which has been damaged. Further the hydrogel has also been used in chemically modified intervertebral disc (IVD) regeneration [54]. The structure is mainly linear with repetitive tetra-saccharide units and one acetyl group and a carboxyl group in its native state. It is thus calcium-sensitive, yet possesses rheological properties that are similar to xanthan gum having similar density of charges on the molecule [55,56]. Xu and his coworkers developed hydrogel with GG and successfully altered by treating it with methacrylic anhydride [57]. By changing the proportion of methacrylic anhydride, low-modified and high-modified GG hydrogels were created. The hydrogels were made using chain-growth, step-growth, and mixed-model methods, all of which took use of the thioleene click chemistry. The compressive modulus of the resulting hydrogels can be dramatically altered by adding calcium and changing the thioleene ratio. Hydrogel swelling ratios were shown to be inversely related to hydrogel compression modulus. The creation of physical cross-links by adding calcium to the network reduced network flexibility and inhibited hydrogel expansion. The hydrogels showcased proliferation of NIH/3T3 fibroblasts which demonstrated a favorable connection with stiffness character of the substrate. The spreading behavior and morphology of the cells were shown to be connected to the structure and functional groups of the substrate. An optimal ratio of thiol-ene, calcium concentration, and cross-linking characteristics of mixed mode, in general, can result in hydrogels which are much stiffer and which are linked to proliferating fibroblasts. Overall, they concluded that chemical alteration and the addition of divalent ions can be used to adjust the mechanical properties of GG hydrogels. The information offered here will aid in the development of a GG platform for future tissue engineering applications [57]. Biomaterials that meet the specific necessities of human tissues and neural cells are needed for neural tissue engineering and three-dimensional in vitro tissue modeling. Koivisto and her coworkers devised an alternate method for producing biomimetic hydrogels instead of the conventional cross-linking process using bioamines such as the spermine and spermidine [58]. These bioamines have been shown to act as cross-linking agents for producing GG hydrogel at a temperature of 37
C, which allowed human neurons to be encapsulated. The produced hydrogel has undergone the mechanical and rheological testing, and they demonstrated properties which mimicked biologically and could be compared to the behavior of brain tissue of naive rabbit under situations of stress or strain. Based on the above trials, they concluded that the addition of laminin to GG hydrogels caused cell type-specific activity, such as neuronal cell maturation and neurite movement [58] (Fig. 1.9).

1.2.2.3 Dextran Dextran is a common term for a group of glucans produced by the polymerization of a-D-glucopyranosyl moiety of sucrose, which is catalyzed by the enzyme dextransucrase. A preponderance of (1/6)-linked a-D-glucopyranosyl units is a typical trait. The polymer has been found to be synthesized from various species of microorganisms, and a number of

FIGURE 1.9 Chemical structure of gellan gum.

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molecular weights and chemical structures have been discovered with different branching characteristics. The nonpathogenic organism Leuconostoc mesenteroides NRRL B-512 is responsible for commercial dextran biosynthesis. [n sucrose / (a-D-glucopyranosyl unit)n þ n fructose]dthis basic reaction is catalyzed by dextransucrase. Branching has been observed at O-3 of the glucosyl unit. Branching of nearly 5% has been observed for dextrans that are available commercially. Around 40% in the side chains of the polymer consisted of single a-D-glucopyranosyl units, around 45% was made up by two units, and around 15% was made up by more than two units. The polymer differs from other high molecular weight polysaccharides in that it creates low viscosity solutions. It is a non toxic polymer having antithrombotic and antiinflammatory properties [59,60]. By photo crosslinking modified dextran, a polysaccharide-based hydrogel was created by Kim and his co-researchers [61]. Bromoacetyl bromide was used to bromoacetylate dextran, which was then treated with sodium acrylate to include a vinyl group. After that, a long-wave UV lamp was used to photo crosslink the acrylated dextran. Elemental analysis, fourier transform infrared (FTIR), 1H-NMR, and 13C-NMR were used to characterize reaction products (bromoacetyl dextran, acrylated dextran, and hydrogel). Depending on the degree of substitution in bromoacetyl dextrans, they concluded that the produced dextran hydrogels exhibited a wide range of swelling behavior in different pH conditions [61]. Dex-SS, the dextran hydrogels consisted of Schiff base links as well as disulfide bonds and were created using a simple approach consisting of oxidation of dextran and thereafter producing of Schiff base bonding between cystamine and polyaldehyde dextran referred to as Schiff base reactions containing disulfide. Studies related to swelling, viscosity determination, and 13C CP/MAS NMR results revealed that the degree of cross-linking in Dex-SS hydrogels was greatly influenced by -CHO/-NH2 ratio. SEM, rheology investigation, and Ellman’s assay were used to demonstrate degrading behaviors of Dex-SS hydrogels which were dependent on the reductive and acidic environment. Furthermore, doxorubicin (DOX) was injected into the matrix of the hydrogel and released in a manner dependent on pH/GSH. Transwell assays with HepG2 cells confirmed that Dex-SS hydrogel was cytocompatible and had excellent cell absorption of the released drug [62] (Fig. 1.10).

1.2.2.4 Scleroglucan Scleroglucan is a name that refers to a group of glucans generated by fungi, particularly those belonging to the genus Sclerotium. Scleroglucan is the brand name for the commercial product. Halleck, from “The Pillsbury Company”, was responsible for its early research. He sparked an immense interest in this polysaccharide [63]. Scleroglucan has been utilized in various industrial applications such as thickening drilling muds, fracturing and completion fluids, and increased oil recovery, due to its outstanding rheological qualities and resilience to electrolytes, temperature, and hydrolysis. Scleroglucan is made by growing a specific strain of Sclerotiutn rolfsii in an aerobic submerged culture. The fermentation is done in a standard fermenter, and batch culture is utilized. As a carbon source, the medium contains roughly 3% Dglucose, corn-steep liquor, nitrate, and mineral salts. The fermentation may be done using a variety of carbohydrate sources, and the organism is so stable that it can even be cultivated in sea water. There are two types of scleroglucan in the market: native scleroglucan, which contains both the polysaccharide and the mycelium, and refined scleroglucan, which does not. Direct precipitation of the sterilized broth with oneetwo volumes of isopropanol yields native scleroglucan. The mycelium is removed from purified scleroglucan by filtering the dilute broth in the presence of a filtration aid. Isopropanol is then used to precipitate the scleroglucan. Scleroglucan comes in two grades, each of which is sold as a white, odorless powder that is around 100 microns in size. The separation of scleroglucan from its broth has been proposed using poly(ethyleneglycol) [64].

FIGURE 1.10 Chemical structure of dextran.

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Scleroglucan is a branched homopolysaccharide that when fully hydrolyzed yields D-glucose. The polymer is made up of a primary chain of (1/3)-connected units of b-D-glucopyranosyl along with one unit of b-D-glucopyranosyl linked (1/6) to every third unit. Periodate oxidation studies and hydrolysis by specific glucanases were used to determine the chemical structure. When Basidiomycetes QM806 exo-b-l,3-glucanase hydrolyzed scleroglucan, one mole of gentiobiose was produced for every two moles of glucose, confirming the ratio of (l/3) and (l/6) glycosidic linkages. D-glucose, gentiobiose, and laminaribiose were formed after partial hydrolysis by Rhizopus arrhizus endo-b-l,3-glucanase, but no laminaritriose, suggesting that the b-(l/6) linkages are evenly dispersed on every third glucose unit of the main chain [64]. A novel hydrogel has been created employing scleroglucan (Sclg) and borax as a cross-linker by Coviello and his coworkers [65]. The release of the medication from the physical gel was tested after it was encapsulated with theophylline (TPH). Drug delivery in simulated gastric and intestinal fluids was determined using the same technique that was used to make tablets. The release profiles were subjected to a current theoretical technique, which found a satisfactory agreement with the experimental data. Besides, the diffusion coefficient was calculated using a suitable technique based on two sets of data gathered using distinct setups obtained through diffusion and permeation studies. A very simple mathematical technique permitted the Fick’s second law’s two-dimensional problem to be reduced to a one-dimensional system, resulting in significantly easier data management without sacrificing the correctness of the original two-dimensional problem. The dynamics of the swelling characteristics based on the uptake of water was also followed in the characterization of the gel. As seen in the obtained findings when the hydrogel was impregnated with the drug, the hydrogel consisting of borax and Sclg was observed to be acceptable for long-term drug delivery [65]. The physicochemical features of a semi-IPN consisting of intermediate alginate chains in the hydrogel network of scleroglucan/borax, as well as its applicability for modified drug release formulations, are discussed by Matricardi and his coworkers [66]. The network ability to differentiate the drug release with varied steric hindrance was examined in terms of the practicality of a drug delivery device consisting of the innovative matrix structure of polysaccharide. The semi-IPN hydrogel’s mechanical properties revealed that alginate has a significant effect on the system consisting of borax/scleroglucan. Studies in the regime of shear oscillation revealed that the viscosity characteristics of the system of polymers were more than additive; in fact, alginate causes an order of magnitude raise in the storage modulus of the hydrogel. Optical data gathered in circular dichroic studies revealed alginate and scleroglucan were compatible at the molecular level, regardless of the presence of borax. In the temperature range investigated, the observed semi-IPN is thermally irreversible [66] (Fig. 1.11).

1.2.2.5 Pullulan Pullulan, an extracellular polysaccharide was discovered in Aureobasidium pullulans (syn. Pullularia pullulans) for the first time [67]. It is a natural polysaccharide obtained through fermentation of starch and is soluble in water. Pullulan is a white-colored, odorless, flavorless, and extremely stable powder [68]. Pullulan (Pullulan PI-20) is a commercially available polysaccharide with a purity of about 90%. Mono-, di-, and oligosaccharides are the most common contaminants found in it. Due to the film-forming capabilities, pullulan is an excellent alternative of gelatin for application in tablet coatings or the production of capsule shells and edible flavored films. Since 1976, it has been utilized as an additive and culinary component in Japan [69]. FIGURE 1.11 Chemical structure of scleroglucan.

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Pullulan, a homopolysaccharide of linear structure, consists of maltotetraose and maltotriose units linked in a regular pattern by a-(1/6) and a-(1/4) connections. Based on the environment of culture and differences in strain, it may also contain a tiny amount of a-(1/3) linkages [40]. The molecular mass of pullulan ranges from around 10 to 3000 kDa, depending on the production parameters such as time, pH, and phosphate concentration. Pullulan is structurally identical to starch amylopectin and maltodextrin, both of which contain glucose units consisting of a-(1/6)- and a-(1/4) glucosidic linkages, just like pullulan. The variation in the amounts of these links is the most significant distinction. Maltodextrin consists of a-(1/6)-glucosidic linkages of nearly 20%, whereas pullulan contains around 30%. Another polymer, corn starch, on the other hand, comprises 95% a-(1/4)-glucosidic bonds and 5% a-(1/6)-glucosidic bonds. Furthermore, structure of the molecule is being tertiary, as well as the extent and process of material degradation in the gastrointestinal tract, differs between pullulan and these glucans [69]. Wound dressings made of carbohydrate polymers have been around for a long time. Pullulan is a carbohydrate polymer made by Aureobasidium strains that is abundant in nature. Priya and her coworkers employed pullulan hydrogel (10%) in their study to examine its healing efficiency on wounds in rats which were sutureless [70]. They utilized 150e200 gm weighing male Wistar rats and separated them into three groups. Each group consisted of six rats, and a cut on the dorsum was made which was 3 cm in thickness. 500 mL of hydrogel was administered to the wounds of Group I rats, and then fingers were clamped for 2 min. Authors administered povidine-iodine to Group II (positive control) rats, but not to Group III (control) rats. The medication was inserted once a day until the sores were totally healed. The number of days necessary for complete healing was used to determine the healing rate. The incision made in Group I rats in which hydrogel was applied, healed in six days, whereas the wounds of the positive control and control rats took 11 and 15 days to heal, respectively. The animals were sacrificed, and the wound-breaking strength of incision was determined [70]. Hydrogels were prepared with poly(vinyl alcohol) (PVA) and pullulan (HP) by means of conventional cross-linking by chemical means with the cross-linker sodium trimetaphosphate (STMP) and by dual means using both chemical and physical cross-linking processes including the technique of freeze-thaw [71]. The gels produced including the cryogels and hydrogels that resulted were created with tissue engineering in mind. Two different grades of PVA were used with molecular weights of 47,000 (PVA47) and 125,000 (PVA125) g/mol and varied weight ratios of HP and PVA. FTIR spectroscopy and SEM were used to characterize the hydrogels. In simulated fluids such as phosphate buffer of pH 7.4, the swelling kinetics, dissolving behavior, and degradation profiles were studied. The degradation profiles, swelling, and pore size were all affected by the pullulan concentration and cross-linking procedure. Cryogels have a smaller capacity for swelling than ordinary hydrogels, but they are more resistant to hydrolytic breakdown. LDH (lactate dehydrogenase) and MTT (3-(4,5-dimethylthiazol-2-yl)-2,5 diphenyltetrazolium bromide) assays were used to determine the biocompatibility of the hydrogels. It was observed that the biocompatibility of the HP/PVA125 scaffolds in 75:25 w/w ratio which were cross-linked dually was higher and enhanced L929 murine fibroblast cell adhesion and proliferation to a larger extent compared to HP/PVA47 scaffolds in 50:50 w/w ratio cross-linked only by chemical means, according to MTT and LDH experiments. Furthermore, the HP/PVA125 cryogel exhibited the best ability to differentiate cells into adipogenic differentiation. Overall, the HP/PVA composite hydrogels or cryogels were found to be excellent biomaterials for tissue engineering applications [71] (Fig. 1.12).

1.2.2.6 Levan Levan is synthesized in the form of an exopolysaccharide (EPS) and obtained from the ECM of different bacteria such asdBacillus, Streptococcus, Acetobacter, Erwinia, Aerobacter, Zymomonas, Gluconobacter, Azotobacter, Mycobacterium, Corynebacterium, and Pseudomonas [72]. In addition, Poli and her coworkers discovered Halomonas sp., a potential source of levan [73]. Halomonas levan was explored further as a bioflocculating agent [74], in drug delivery devices for proteins and peptides [75], biocompatible material [76,77], sticky multilayer film [78], and a heparin-mimicking glycan [79,80]. The homopolymer of fructose, levan, is a naturally occurring fructan. Its primary chain is made up of five-member fructofuranosyl rings that are linked together by b-(2/6) links. The major chain is branched by the fructofuranosyl ring b-(2/1) linkage [69]. Effective pathogen-capturing materials are necessary to stop the transmission of pathogenic microorganisms such as the influenza virus. Using levan polysaccharide, Kim and his coworkers presented a new pathogen-capturing and recovery material. They made hydrogels by combining levan and PVA and cross-linking them using glutaraldehyde. The authors observed that the water solubility and adsorption ability of fabricated levan-PVA hydrogels are high. Levan-PVA hydrogels exhibited a 3D porous structure, according to SEM measurements. The effectiveness for the capture of influenza virus of levan-PVA hydrogels is higher than that of commercial cotton swabs, according to RT-PCR research. Furthermore, they

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FIGURE 1.12 Chemical structure of pullulan.

confirmed that bioaerosol samples were effectively caught using levan-PVA hydrogels on gauze as a filter medium. As a result, levan-PVA hydrogels could be simple and effective pathogen capture and recovery materials [81]. By mixing levan with pluronic and carboxymethyl cellulose, Choi and his co-researchers were able to assess the utility of levan as a novel polymer to be used as dermal filler and create a hydrogel consisting of levan which can be applied as injection as well as physically [82]. The hydrogel was prepared at an increased temperature of 37
C after mixing the polymers to produce a sol state in a specified ratio at 4
C for two days. The levan hydrogel had a greater elastic modulus than an HA-based hydrogel. The levan hydrogel had an interconnected porosity structure, comparable to the HA hydrogel, as seen in SEM pictures. In human dermal fibroblast cells, Levan demonstrated noncytotoxicity, increased cell proliferation, and increased collagen synthesis when compared to HA. In contrast to the Pluronic F127 hydrogel or the HA hydrogel, the levan hydrogel which was injected was found to be stable for two weeks as well as biocompatible in vivo. Additionally, in case of in vivo studies, the levan hydrogel produced more collagen than the HA hydrogel. More crucially, the levan hydrogel outperformed the HA hydrogel in terms of antiwrinkle efficacy in a wrinkled model mouse. As a result, the injectable, biocompatible, and anti-wrinkle levan hydrogel has a lot of potential as a replacement for existing commercial dermal fillers [82] (Fig. 1.13).

1.2.2.7 Schizophyllan Schizophyllan (SPG), an extracellular natural polysaccharide with a,b-1,3-D-glucose structure, occurs in water in the form of triple helix which is stiff in nature and possesses a pitch of 0.30 nm per residue, a diameter of 2.6 nm, and persistence length of 200 nm, when the average molecular mass is greater than 9  104. Sarcoma 180, a mouse tumor, responds to a schizophyllan aqueous solution with host-mediated anticancer action. As schizophyllan is a rod-like rigid polymer when dissolved in water, a concentrated aqueous solution of schizophyllan may be observed as two parts. In the cholesteric phase, the triple helix structures of schizophyllan align in a direction parallel with the isotropicecholesteric interface [83]. Schizophyllan, on the other hand, dissolves in dimethyl sulfoxide as a single chain. The triple helix breaks down into three

FIGURE 1.13 Chemical structure of levan.

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coiled chains in water and dimethyl sulfoxide solutions in an all-or-none fashion. Schizophyllanepolynucleotide macromolecular complexes can be generated when a single chain of schizophyllan is combined with polynucleotide [84]. This molecule acts like a complimentary polynucleotide chain for the matching polynucleotide. A triple helix is formed by two molecules of schizophyllan and one molecular chain of poly(cytosine) of poly(adenine) [85e88]. Several research groups have investigated the effect of sorbitol on the formation of gel with the polysaccharide. According to rheology, DSC, and ORD [89], the formation of gel is caused by the conversion of the triple helix structure from level of II to I, thus resulting in 3D structure of interpenetrated chains of triple helix I which renders the necessary stiffness to the structure. Furthermore, neither the region of the junction points zone nor the accumulation of SPG triple helices together were found to be responsible for the formation of gels of SPG-sorbitol [90]. Schizophyllan is a water-loving b-glucan derived from the Schizophyllum commune. Due to its unique features carrying a structure of triple helix, immunological modulation and efficacy against tumor progression, SPG has been extensively studied. Lee and her coworkers attempted to make an SPG hydrogel using photopolymerization method with thiol-ene. They used ultrasonicated SPG to make SPG-norbornene and SPG-thiol [91]. The authors used thiol-ene photo-click reaction to create two different forms of SPG hydrogels like SPG as well as PEG/SPG hybrid. The Young’s modulus and the swelling characteristics of SPG hydrogels may be regulated within 20e60 and 0.5e10 kPa, respectively, by changing SPG concentration and amount of thiol and norbornene groups. The authors discovered that the backbone of SPG as a triple helix had an effect for the creation of a physical network by cross-linking with thiol-ene groups in a PEG/SPG hybrid hydrogel. The breakdown of the hydrogel can also be controlled by adjusting the formulation. From the above observations, they have concluded that highly controllable SPG hydrogels could be used in a variety of applications [91]. Utilizing the interaction between glucose and boronic acid, Shingo and his co-researchers demonstrated that new hydrogel materials can be prepared from made from poly(acrylic acid) modified with boronic acid and SPG. Moreover, they observed that SPG’s natural property of wrapping as single-walled carbon nanotube (SWCNT) can be used to create a hydrogel containing carbon nanotubes. Due to the high stability of the SWCNT-based hydrogel, it has a lot of potential for biosensors, biomineralization, artificial muscles, and other applications. As, SPG may entrap a variety of other guest molecules [92,93], this stimuli-responsive hydrogel can be combined with additional capabilities intrinsic to the wrapped guest molecules, resulting for its application to be used as chemical sensors, or in drug delivery, or in self-healing products, etc. [94] (Fig. 1.14).

1.2.2.8 Curdlan It is the third polysaccharide generated by fermentation that has been approved for use in food in the United States. Curdlan is already used in a wide range of commercial food products in countries like Taiwan, Korea, and Japan. Curdlan is made by fermenting a culture of Alcaligenes faecalis var. myxogenes in a culture medium that contains some trace minerals, glucose, and suitable nitrogen sources. Raw curdlan is dissolved in alkali, removed from the microbe, purified, and dried into powder [45]. Curdlan belongs to the (1,3)-b-glucans family of compounds. The structure consists of carbons at the first and third positions of the glucose residues being interlinked with a b linkage repetitively. Structurally curdlan is a long chain, but intramolecular and intermolecular hydrogen bonding causes it to create more complicated tertiary structures [55]. A successful development of hydrogel coated with polydopamine and curdlan is reported by Michalicha and her coworkers. They observed that curdlan hydrogel generated through thermal gelling is nimble and has a high water

FIGURE 1.14 Chemical structure of schizophyllan.

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FIGURE 1.15 Chemical structure of curdlan.

absorption capacity. However, its modification which leads to the upraise of biofunctionality do very often affect the mechanical strength and solubility. As a result, polydopamine was deposited on the curdlan hydrogel to increase its ability to establish bonding with the free amine groups of drugs. As before, it had similar mechanical characteristics and structure as well as soaking capacity. A model drug, an aminoglycoside antibiotic (gentamicin), was successfully adsorbed on modified curdlan structure through the quinine group of polydopamine. Following Fickian diffusion, around half of the immobilized medication was released, inhibiting bacterial growth in the matrix-surrounding medium for a long time. Even after the entire mobile drug had been released, the residual drug quantity was securely bonded to the hydrogel, preventing bacterial adherence. As a result, polydopamine-modified curdlan hydrogel has the potential for developing various functional materials, like drug-loaded biomaterials [95] (Fig. 1.15). Lin and his co-researchers established two-step self-assembly technique to develop triple helix curdlan preparation procedure for creating a highly stretchy hydrogel containing silver and being transparent. They also observed that the internal nanofibril network of this hydrogel, which comprises nanofibrils with a diameter of 20 nm, reinforces it. Due to the unique structure of the hydrogel as a triple helix, this hydrogel has a 350% of fractured strain and 0.2 MPa of fractured stress, and a water content of 97%, which is significantly improved compared to hydrogels of other polysaccharides having similar percentage of water. From the results of animal studies, it can be clearly observed that this curdlan hydrogel containing silver also boosts fibroblast growth, efficiently causes the inhibition of infections due to bacteria, reduces the levels of inflammatory factors, and improves healing of wounds topically. These qualities encourage the hydrogel’s continued development as a prospective material for wound healing for clinical purposes [96]. In another study, curdlan hydrogels were developed with excellent stretching properties by the use of chemical crosslinking agents like 1,6-hexanediol diglycidyl ether, 1,4-butandiol diglycidyl ether, and ethylene glycol diglycidyl ether. The hydrogels thus produced were subjected to tests for evaluating their tensile strength. They were found to extend by 600%e900%. When the stretched films of the hydrogels were dried and evaluated, the values for their tensile strength ranged from 117 to 148 MPa and the values obtained for Young’s modulus was 1.6 GPa which were considerably higher than the films obtained from nonstretched hydrogels. During the processing conditions of drying and stretching, molecules of curdlan in the hydrogels were crystallized under strain and achieved regular orientation which resulted in the excellent tensile strength of the hydrogels. This was confirmed by X-ray crystallographic studies which concluded about the crystalline structure of the curdlan molecules in the dried and stretched hydrogels with a degree of orientation of about 80% [97] (Fig. 1.16).

1.3 Preparation of polysaccharides through biotechnological approach Biopolymers have been observed as the natural polymers that are produced and accumulated within the cells of organisms during their normal life cycle as metabolic products like glycogen, carbohydrates, or starch. Few biopolymers have also been extracted from algal species [98,99]. Recently biopolymers have also been synthesized without cells with the help of enzymes in vitro [100]. They have been produced through fermentation. The accumulated metabolic products extracted as biopolymers from the intracellular or extracellular components can be done efficiently through biotechnological production methods [101,102]. Polysaccharides obtained from microbial sources or the microbial polysaccharides can again be classified into three types namely: (a) the structural polysaccharides which are present in the cell wall structure of the microorganism like peptidoglycans or teichoic acids; (b) intracellular/cytosolic polysaccharides (IPSs) which function as carbon source and release energy in cell, and (c) exopolysaccharides or EPS that are released by cells as capsules or as biofilm in the extracellular space [103e107] (Fig. 1.17).

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FIGURE 1.16 Preparation of (A) dried gel films and (B) stretched dried gel films. Reprinted from Carbohydr Polym, 269, Matsumoto Y, Enomoto Y, Kimura S, Iwata T, Highly stretchable curdlan hydrogels and mechanically strong stretched-dried-gel-films obtained by strain-induced crystallization, 118312, Copyright (2021), with permission from Elsevier.

FIGURE 1.17 Biotechnological production of exopolysaccharides.

Recent research has been focused on biotechnological production of microbial derived polysaccharides, especially the exopolysaccharides. The biotechnological technique has many benefits compared to chemical or plant-based methods due to higher energy efficiency, rapid volumes of production which are not dependent on the location of the sources, or the season during which the production is done. Further industrial wastes as well as agricultural wastes may be utilized in the process as substrate which adds extra value. A major challenge in this regard may be to reduce the production cost which is generally on the higher side [108e110]. Few exopolysaccharides discovered through recent research include Microbactan [111], or the EPS extracted from Rhodotorula mucilaginosa [112] or from Lactobacillus planta [113]. The process by which each EPS is produced by the microorganisms vary and depend on many factors. However generally the EPSs have been observed to be produced and stored in the cells following the growth phase [114]. Various types of enzymes are involved in most of the steps for the biosynthesis of EPS such as the intracellular enzymes like hexokinase, enzymes present in the cell membrane like glycosyltransferases, or extracellular enzymes [106].

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The amount of EPS produced is directly linked to the process conditions and medium composition [106]. Availability of carbon and nitrogen sources, temperature, pH, growth phase, agitation, and aeration plays major influence on EPS production by microorganism, especially bacteria [115,116]. The optimization of the process and media conditions is further required for maximum yield and higher production rates [117]. Thereafter the extraction of the EPS from the culture medium becomes important, and there exists physical and chemicals processes for the purpose. However there exists limitations in each case. For example, for extraction of EPS by chemical means use of cation exchange resins or ethylene diamine tetraacetic acid may be done. However, contamination of the polysaccharide has been observed in such case [118]. By physical means like application of heat, centrifugal force, or using ultrasonicator, the yield may reduce [119,120]. In recent years, several techniques have been devised for the extraction of EPS. In a study, the extraction has been carried through three basic steps which included both the physical and chemical methods namely the removal of the cells from the culture medium usually by applying centrifugal forces, precipitation of the polysaccharide through the use of precipitating agent like ethanol, acetone, methanol, isopropylalcohol, etc., which are miscible with water and finally drying of the polymer [121]. Unless the EPS is heat labile, the use of heat can be done to decompose the enzymes responsible for degradation of the polysaccharides [121,122]. However, few other methods may be used like deproteinization or reprecipitation when a higher level of purified product is important [123]. Therefore, optimization of the extraction technique is important. One can choose a process which would involve least amounts of lysis of cells, avoid the disruption of EPS, and prevent higher quantities of proteins or DNA to be released [124,125]. Whatever may be the technique, the extraction of polysaccharides is a bit complicated and 100% of extraction for the EPS may not be possible. However, proper optimization techniques may increase the yields which again depend on the type of the polysaccharide and the ultimate objective of the process with respect to the quality or the quantity [126]. The extracted product may however contain some other components apart from the polysaccharide like proteins, chemicals, or DNA fragments. Therefore, it becomes necessary to purify the extracted product. Thus, reprecipitating of the EPS from its solution may be considered from the extract. One may consider and optimize a suitable process for purifying the polysaccharide based on the effect of the methods on percentage purity, recovery of the product, or characteristics of the polymer. Few processes may also reduce the yield of the product or affect the properties of the polymer in a negative way [121]. Considering the above facts, the method for purification should be carefully chosen which would not only increase the yield but also assure purity and quality of the product [127].

1.4 Future prospect of the polysaccharide hydrogels in drug delivery and regenerative medicine From the above account, it is clearly understood that hydrogels based on natural polysaccharides from animal and microbial sources as well those obtained through biotechnological procedures have significant advantages. In the last decade, much progress has been made with respect to drug delivery, wound healing, and tissue engineering utilizing them. However, a lot more remains to be done especially compared with their tremendous potential.

1.5 Conclusion The above account on the various polysaccharides derived from animal, microbial, and biotechnological sources can easily provide an impression about the tremendous potential of the polymers for drug delivery and biomedical applications. However, challenges remain in certain cases to provide delivery systems with better stability and drug targeting potential. Further, the production of polysaccharides using microorganisms by adopting biotechnological techniques was also discussed. The factors affecting the production, the extraction, and purification techniques were also summarized. In the next few chapters, we shall find a detailed account of each of the polysaccharides, their properties, applications, sources, extraction methods, and their utilization to produce interconnected hydrogels which have tremendous possibilities in drug delivery and biomedical applications.

Acknowledgments The authors acknowledge the support of the seed fund granted by Adamas University, Kolkata, West Bengal, India.

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Highly stretchable curdlan hydrogels and mechanically strong stretched-dried-gel-films obtained by strain-induced crystallization. Carbohydrate Polym 2021;269:118312. ISSN 0144-8617. [98] Mahmoud YA, El-Naggar ME, Abdel-Megeed A, El-Newehy M. Recent advancements in microbial polysaccharides: synthesis and applications. Polymers 2021;13(23):4136. [99] Mayara CSB, Vespermann KAC, Pelissari FM, Molina G. Current status of biotechnological production and applications of microbial exopolysaccharides. Crit Rev Food Sci Nutr 2019;60:1475e95. [100] Steinbüchel A. Perspectives for biotechnological production and utilization of biopolymers: metabolic engineering of polyhydroxyalkanoate biosynthesis pathways as a successful example. Macromol Biosci 2001;1:1e24. [101] Pandey S, Shreshtha I, Sachan SG. Microbial exopolysaccharides as novel and significant biomaterials. Amsterdam, Netherlands: Springer; 2021. [102] Wang J, Salem DR, Sani RK. Microbial and natural macromolecules. Amsterdam, The Netherlands: Elsevier; 2021. [103] Bergmaier D. Production D’exopolysaccharides par fermentation avec des cellules immobilisees de LB. Rhamnosus RW-9595M D’un milieu A Base de permeat de lactoserum. Resume 2002;2:2e3.

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[104] Boels IC, Kranenburg RV, Hugenholtz J, Kleerebezem M, De Vos WM. Sugar catabolism and its impact on the biosynthesis and engineering of exopolysaccharide production in lactic acid bacteria. Int Dairy J 2001;11(9):723e32. [105] Lahaye E. Role structurant des exopolysaccharides dans, un biofilm bacterien. 2006. http://www.theses.fr/2006LORIS071. [106] Kumar AS, Mody K, Jha B. Bacterial exopolysaccharides e a perception. J Basic Microbiol 2007;47(2):103e17. [107] Ruas-Madiedo P, Gavilan CG. Invited review: methods for the screening, isolation, and characterization of exopolysaccharides produced by lactic acid bacteria. J Dairy Sci 2005;88(3):843e56. [108] Lopez GCV, Fernandez FGA, Sevilla JMF, et al. Utilization of the Cyanobacteria anabaena Sp. ATCC 33047 in CO2 removal processes. Bioresour Technol 2009;100(23):5904e10. [109] Rutering M, Jochen S, Broder R, Martin S, Volker S. Controlled production of polysaccharides-exploiting nutrient supply for levan and heteropolysaccharide formation in Paenibacillus Sp. Carbohydrate Polym 2016;148:326e34. [110] Thompson JC, He BB. Characterization of crude glycerol from biodiesel production from multiple feedstocks. Appl Eng Agric 2006;22(2):261e5. [111] Camacho-Chab CJ, Castaneda-Ch R, Jes M, Noem R, Galaviz-Villa I. Biosorption of cadmium by non-toxic extracellular polymeric substances (EPS) synthesized by bacteria from marine intertidal biofilms. Int J Environ Res Publ Health 2018;25(314):1e11. [112] Vazquez-Rodriguez A, Vasto-Anzaldo XG, Perez DB, Chapoy-Villanueva H, Rivas GG, Garza JA, et al. Microbial competition of Rhodotorula mucilaginosa UANL-001L and E. coli increase biosynthesis of non-toxic exopolysaccharide with applications as a widespectrum antimicrobial. Sci Rep 2018;8:798. [113] Sasikumar K, Vaikkath DK, Devendra L, Nampoothiri KM. Bioresource technology an exopolysaccharide (EPS) from a Lactobacillus plantarum BR2 with potential benefits for making functional foods. Bioresour Technol 2017;241:1152e6. [114] Becker A. Challenges and perspectives in combinatorial assembly of novel exopolysaccharide biosynthesis pathways. Front Microbiol 2015;6:687. [115] Barbosa AM, Paulo DTC, Mariane MP, Lourdes C. Production and applications of fungal exopolysaccharides. Semina Ciências Exatas Tecnol 2004;25(1):29e41. [116] Czaczyk K, Myszka K. Biosynthesis of extracellular polymeric substances (EPS) and its role in microbial biofilm formation. Pol J Environ Stud 2007;16(6):799e806. [117] Chug R, Gour VS, Mathur S, Kothari SL. Optimization of extracellular polymeric substances production using Azotobacter beijreinckii and Bacillus subtilis and its application in chromium (VI) removal. Bioresour Technol 2016;214:604e8. [118] Comte S, Baudu M. Relations between extraction protocols for activated sludge extracellular polymeric substances (EPS) and EPS complexation properties: Part I. Comparison of the efficiency of eight EPS extraction methods. Enzym Microb Technol 2006;38:237e45. [119] Liang Z, Wenhong L, Shangyuan Y, Ping D. Extraction and structural characteristics of extracellular polymeric substances (EPS), pellets in autotrophic nitrifying biofilm and activated sludge. Chemosphere 2010;81(5):626e32. [120] Liu H, Fang HHP. Extraction of extracellular polymeric substances (EPS) of sludges. J Biotechnol 2002;95(3):249e56. [121] Freitas F, Alves VD, Reis MAM. Advances in bacterial exopolysaccharides: from production to biotechnological applications. Trends Biotechnol 2011;29(8):388e98. [122] Singha TK. Microbial extracellular polymeric substances: production, isolation and applications. IOSR J Pharm 2012;2(2):276e81. [123] Roca C, Alves VD, Freitas F, Reis MAM. Exopolysaccharides enriched in rare sugars: bacterial sources, production, and applications. Front Microbiol 2015;6:288. [124] Gehr R, Henry JG. Removal of extracellular material techniques and pitfalls. Water Res 1983;17(12):1743e8. [125] Liu H, Fang HHP. Extraction of extracellular polymeric substances (EPS) of sludges. J Biotechnol 2002;95(3):249e56. 1983. [126] Donot FA, Baccou FJC, Schorr-Galindo S. Microbial exopolysaccharides: main examples of synthesis, excretion, genetics and extraction. Carbohydrate Polym 2012;87(2):951e62. [127] Sõlvia A, Freitas F, Sevrin C, Grandfils C, Reis MAM. Production of FucoPol by enterobacter A47 using waste tomato paste by-product as sole carbon source. Bioresour Technol 2017;227:66e73.

Chapter 2

Glycogen-based hydrogels Bijaya Ghosh1 and Tapan Kumar Giri2 1

NSHM College of Pharmaceutical Technology, NSHM Knowledge Campus, Kolkata Group of Institutions, Kolkata, West Bengal, India;

2

Department of Pharmaceutical Technology, Jadavpur University, Kolkata, West Bengal, India

2.1 Introduction The importance of polymers as biomaterials can never be overestimated. As drug delivery techniques get more and more sophisticated, the use of target-oriented polymers, especially biodegradable and biocompatible ones, keeps rising. The last few decades had seen a steady demand for biofunctional polymers in the field of drug delivery and tissue engineering. Initially, the biomedical field was more interested in developing custom-made synthetic polymers rather than the ones available from natural sources. However, lately the use of natural biopolymers, especially polypeptides and polysaccharides, has increased significantly [1]. Polymers of natural origin bring with them the advantages of both low temperature synthesis and comparatively free from compounds that contribute to environmental hazards. The frequency of oral administration was really high till the 1950s. Most of the drugs needed three to four times administration a day [2]. This resulted in patient noncompliance as well as systemic side effects caused by unavoidable peaks and troughs [3]. The discovery of the release sustaining polymers partially solved that problem. Most of the polymer obtained for drug delivery research has been derived from enzymatic or synthetic laboratory manufacturing pathways. In this context, there is a general interest to prepare new materials using natural polymer as a starting point. Interest has been drawn to glycogen too [4]. Glycogen, a natural constituent of the body, is an excellent biomaterial. Because of its “hyperbranched structure,” it can pack a very high number of glucose entities in a small space. The cross-linked polymer constructs developed with glycogen retain its high water imbibing capacity which is desired in tissue engineering biomaterials [5].

2.2 Glycogen 2.2.1 Source Glycogen is a versatile polymer available in nature in abundance. In nature, it serves as a storehouse for energy. It is present in almost all living beings starting from simple bacteria to the most complicated life forms, that is, humans [6,7]. In mammals, glycogen deposits are mainly found in the liver and skeletal tissue. Approximately 5%e10% hydrated weight of the liver is contributed by glycogen. Compared to that, the proportion in the skeletal tissue is less: approximately 2%. Glycogen is produced by prokaryotes also. Many strains of Streptomyces, Rhizobium, and Methanococcus produce glycogen, especially during the stationary phase of their growth [8,9]. Accumulation of glycogen was also noticed in some opportunistic pathogens like Enterobactor and Escherichia [10]. However, not all the sources are suitable for mass production of glycogen. Aikawa et al. attempted to produce glycogen using bacteria, but the efficiency of the process was low [11]. The main difficulty lay in the extraction of the synthesized glycogen. The product accumulated within the cell, and the researchers had to disrupt the cells to extract it [11]. For successful bacterial production of glycogens, two factors are important: the species utilized for production should be nonpathogenic and the polymer should be secreted extracellularly. Researchers have also demonstrated the extracellular production of glycogen from nonpathogenic Pseudomonas by culturing it in lysogeny broth (0.5 g/L) [12,13]. However, more research is required to make bacterial production of glycogen a commercial reality. Glycogen-like polymers derived

Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine. https://doi.org/10.1016/B978-0-323-95351-1.00015-6 Copyright © 2024 Elsevier Inc. All rights reserved.

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from plant is known as phytoglycogen. It is available in many plants such as rice, sorghum, and barley, especially in the sugary-1 mutants of maize [14,15].

2.2.2 Chemical structure and composition Microscopic studies indicate that glycogens have dendrimer-like structure with a dense core and lighter periphery. Structurally glycogens are compared to a hyperbranched tree-like structure, where the core is denser than the periphery [16]. Based on size and shape, glycogens can be classified into three forms. The simplest level has glucose units randomly joined to form chains. The next higher level has got the smaller chains linked to form greater branches and folded into a spherical shape, popularly known as beta particles. The size of the beta particles ranges between 20 and 50 nm in diameter. At the third level, much bigger structures of 150 nm in diameter are found. These particles with rosette shapes are known as alpha particles. Alpha particles are poly disperse in nature and remain in association with beta particles. Under acidic conditions, alpha particles dissociate into beta particles [17]. The near spherical shape [18] of glycogen macromolecules makes them ideal for use in nanostructured drug delivery products [19]. The location of the functional moieties is of utmost importance in nanoparticulate target release dosage forms. The surface of such systems is often functionalized to equip them with various properties. The targeting moieties are usually attached to the surface of nanoparticles. On the other hand, when the purpose is imaging, attachment should be done at the deeper layers because dye moieties usually have toxic effects and must be hidden in the core. The hyperbranched structure of glycogen can be used for both purposes. For the development of biological constructs, organic soft particles are thought to be convenient [20] as functionalization can take place both at the surface and core. In tissue engineering, matching of the physicochemical properties with the actual tissue is important. Glycogen, because of its ability to provide multiple sites for reaction, can be used to develop composite hydrogels of the desired physicochemical properties. Zhang et al. used glycogen as the connector between collagen and hydroxyapatite to develop a composite hydrogel for bone tissue repair [21]. Composition wise phytoglycogen has similarity with amylopectin (has both a-1,4 linkages and a-1,6 linkages) but has more branch points [22]. Phytoglycogen is bigger than glycogen having a size distribution in the range of 30e100 nm venom [23]. Various models have been developed to explain the structure of glycogen. A well-accepted model compares the glycogen structure with a series of tiers made from two type’s glucose chains. The outermost tier consisted of unbranched glucose chains, whereas each inner chain has two branches. In this arrangement, 50% of the total glucose residues are present as unbranched chains at the surface of the molecule [24]. Theoretically, a full-sized glycogen would have 55,000 glucose units arranged in 12 tiers and with a size of 44 nm. In practice, glycogen molecules are available in much smaller sizes. Glycogen isolated from skeletal muscle has a size of 25 nm, which can justify a structure of seven tiers [25]. Some glycogens obtained from the heart and liver show a more complicated structure. Here the small beta particles join up to form a more complex structure known as alpha particles. Beta particles are extremely small and can be classified as nanoparticles (20e50 nm in diameter). In contrast, alpha particles are bigger (150 nm) and show a wider particle size distribution. Morphological study of negatively stained particles using transmission electron microscopy also revealed a rosette-like structure composed of small alpha of dimension 20e30 nm [26].

2.2.3 Physicochemical properties In solution, glycogen mainly exists as beta particles, which are nothing but randomly joined glucose chains [27]. The chains consisting of beta particles have an average length of around 10 to 15 glucose units [28]. Polar functional groups of glycogen attract water. This property can be used in moisturizing and emulsifying poorly soluble agents, where the strength of binding depends upon the chain arrangement and branching of the polysaccharides. Glycogen extracted from biological sources is polydisperse in nature. Moreover, uniformity of mass does not guarantee structural uniformity. The location of branch points may differ resulting in different three-dimensional structures [6]. Commercial products often carry a small negative charge due to the presence of residual phosphate. Glycogen has three active eOH groups each of which possess different reactivity. Hydroxyl groups present at the C2 and C3 positions act as secondary alcohols whereas those present at the C6 position act as primary alcohol [29]. In the body, the glycoside bonds of glycogen are broken down under the influence of glucagon which is produced by the islets of Langerhans [30].

2.2.4 Hydrogel-forming ability Hydrogels are increasingly getting investigated in the field of controlled drug delivery. They allow easy adjustment of drug loading as well as release rates through manipulation of structural parameters like porosity and cross-link densities. The

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polymer solvent interaction also affects the extent of the swelling rate and thus affects the capacity of drug loading in hydrogels. The effort was made to customize the glycogen polymer for these specific purposes. Patra et al. developed a stable hydrogel by cross-linking glycogen with N-isopropylacrylamide [31]. In 2020, the same research group [32] had shown the cross-link modification of polymers containing glycogen and glycine can be a useful tool to modify its swelling pattern and mechanical properties [33,34]. Hydrogels in spite of their multiple utility suffer from a major disadvantage. Being soft materials, they may undergo physical damage which reduces their functional efficiency. Self-healing properties are highly desirable in hydrogels that are made for use in active tissues which are under continuous stress (heart, muscle, bone, etc.) Hence, the development of self-healing hydrogels or hydrogels which can recover their structure is an area of active research [35,36]. One of the shortcomings of glycogen hydrogels is their low tensile strength (0.128 MPa) compared to hydrogels made from other polysaccharides [37]. However, these same hydrogels can show superior elongation at break and significant self-healing properties; both parameters show a strong correlation with hydrogen-bonding capacity. Hydrogels made with glycogen and polyvinyl alcohol (PVA) in a ratio of 1:1 showed 96% shape recovery after cutting with a knife [38]. This opened up the possibility of the preparation of self-healing gel by incorporating glycogen as one of the ingredients. The effort has been made to produce glycogen-based customized hydrogel by other research groups too [39]. The mechanical strength of the glycogen network was significantly enhanced through the insertion of ferric ions. Ferric iron (Feþ3) can form coordination bonds with the functional groups associated with glycogen and acrylic acid. A new type of dual cross-linked hydrogel was developed by manipulating the proportion of glycogen and acrylic acid in a mixture and loading it with various concentrations of Feþ3 [38]. On evaluation, the prepared hydrogels displayed high stretching ability and desired mechanical strength. In the formation of a hydrogel network, cross-linking agents play a substantial role. It is thought that glycogen with its hyperbranched structure and plethora of functional groups can be an effective cross-linker. Zhang et al. produced microspheres using guanido decorated glycogen [21].

2.3 Drug delivery applications Though glycogen was discovered in 1857 [40], its therapeutic potential is not yet fully explored. An unmet need is the lack of suitable delivery systems for diseases that affect the distal part of the gastrointestinal system. Smart polymers that can release their drug load selectively in the colon were much sought after. Patra et al. developed a grafted hydrogel network equipped with the property of colon-targeted release [31]. They used glycogen as the basic ingredient and combined it with PNIPAM which can show temperature-driven phase change. PNIPAM dissolves in water, but above its lower critical solution temperature undergoes phase transition [41]. The lower critical solution temperature of PNIPAM is 32 C, while the internal environment of the human body has a temperature of around 37 . NIPAM changes from a soluble state to an insoluble dehydrated state at this temperature. Glycogen was attached to PNIPAM using ethylene glycol dimethacrylate as a linker. The developed network was successful in delivering five ASA and ornidazole using the synthesized hydrogel. The synthesized hydrogel showed poor swelling in acidic media, indicating that there is considerable resistance toward the release of entrapped drugs in the stomach. The author suggested that reduced swelling was due to the protonation of hydrophilic groups of the hydrogel network, which prevented the formation of H-bonds with the surrounding water [31]. In contrast, the swelling was much higher in alkaline media in which hydrophilic groups were unprotonated. In alkaline media, H-bonding is favored and entrapped drugs can be easily released. Zhang et al. utilized the location of multiple functional groups of glycogen as binding sites to develop collagenhydroxyapatite hydrogels. The gel showed high efficiency in mesenchymal stem celldosteoblast differentiation [21]. Water-soluble polysaccharides, due to their potential functionalities, are considered to a desirable biomaterial for nanostructured delivery systems [42]. A research group extracted soluble phytoglycogen from Su-1 maize and reacted the same with octenyl succinic anhydride to produce hydrophobic nanoparticle. They used these nanoparticles as emulsion stabilizers to produce o/w Pickering emulsion having medium chain triacylglycerol as the oil phase. The study claimed that nanoparticles formed a three-dimensional network around the disperse phase to confer long-term stability indicating derivatives of phytoglycogen can be used as stabilizers for Pickering emulsion [43]. The study highlighted two important pointsdfirst glycogen, derived from plant sources, can be reacted with external groups to form customized moieties and second, they can serve as strong emulsifiers. Glycogen itself is not totally bioinactive. It is postulated that body recognizes implanted glycogen as a product of hepatocyte damage and responds by activating the immune system. This property is actually the basis of its use for the activation of polymorph nuclear neutrophils [44]. Another promising research involving glycogen was glycogen-based pH-sensitive nanosystem that was developed for targeted delivery to liver cancer cells [45]. Glycogen particles were decorated with b-galactose through asialoglycoprotein

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[45]. It is known that asialoglycoprotein receptors in the Kupffer cells present in the lumen of hepatic sinusoids recognize the galactose moieties and mark them for degradation. Researcher had successfully targeted the liver cancer cells decorating glycogen with galactose and was able to target liver cancer in a mouse model. Hydrogels made with glycogens are already undergoing clinical trials [5]. SARS-CoV-2 infection had created a dire necessity for vaccines that can generate immediate response. When it was found mRNA vaccines can fulfill that need, the developers faced a problem. Basically, mRNA vaccines are negatively charged large molecules which are produced in vitro. However, delivering these molecules into cells where they can produce the desired protein is a challenge. The effort is to develop nonviral delivery systems for mRNAs as they are safer and more versatile compared to their viral counterparts [46]. Glycogen nanoparticles with their rosette-like dendrimer structure are thought to be a potential carrier for the same. Glycogen is also used for the stabilization of silver nanoparticles. The antimicrobial activity of Ag is due to Agþ ions. In nanoparticles, surface to volume ratio is high. However, nanoparticles tend to grow in size. They combine to form microparticles. Hence stabilization is a big problem. Glycogen with its multiple OH groups complexate easily with metallic ions. The resulting supramolecular constructs can serve as a template for nanoparticle growth. Using this assumption, Bozniak et al. prepared Ag-glycogen hybrid nanoconstructs and tested it against Staphylococcus aureus, Escherichia coli, and Candida albicans [47]. Significant reduction of the growth rate was observed for all these species indicating glycogen can be an effective stabilizer in nanoemulsions of metallic moieties. Effort has been made to develop a delivery system for melittinda putative antidiabetic drug present in bee venom [48]. Hanna et al. synthesized a set of glycogens, and functionalized them with different acid groups to different degrees. Glycogen conjugates with a high number of acidic groups showed high electrostatic attraction for cationic melittin resulting in its encapsulation [49].

2.4 Tumor targeting Glycogen is important in tumor targeting as glycogen conjugates show enhanced permeability and lower lymphatic drainage in the tumor vessels. They don’t need any ligands [50,51].

2.5 Tissue engineering applications Tissue engineering has advanced speedily because the supply of organ and tissue transplants is less than the demand [52e55]. The use of hydrogels as scaffold materials for tissue engineering has been extensively studied [56e59]. For the growth and proliferation of cells in the scaffolds, they require nutrients, vitamins, minerals, a suitable osmotic pressure and pH, adequate oxygen, and a water environment. Hydrogels are porous, cross-linked, hydrophilic structures that can absorb many times their dry weight in water. Because of the hydrogels’ high water content and porous structure, oxygen and nutrients can diffuse quickly within the scaffolds. For tissue engineering, hydrogels are the best option because of their similarities to the natural microenvironment of cells.

2.5.1 Bone tissue regeneration Bone damage from fractures, trauma, and bone tumors cannot self-heal, and existing surgical treatment methodologies produce pain, inflammation, and infection. These conditions necessitate advanced treatment methodologies to restore normal and healthy life. Recently, tissue engineering has been widely used as a technique for the regeneration of damaged tissue [60e63]. Bone tissue engineering is a suitable therapy for bone defects, and hydrogels are appropriate scaffold materials [37,64e68]. Hydrogel scaffolds have biodegradability and biocompatibility and provide other factors for the proper regeneration of cells. Glycogen has several advantages, including better biocompatibility, more cross-linking points, and a stereochemical structure. Hydrogels based on glycogen, collagen, and hydroxyapatite were prepared for bone tissue regeneration [21]. First, glycogen was conjugated with ethylenediamine and then grafted with the guanido group. The grafted particle was oxidized with sodium periodate to provide an aldehyde group. The composite hydrogel was prepared by mixing nanohydroxyapatite, collagen, and modified glycogen. The bond (Schiff base) between collagen and glycogen was created through electrostatic interaction. The cross-linking degree of the composite hydrogel was determined through a swelling study in PBS buffer. It was observed that increases in nanohydroxyapatite, collagen, and glycogen decrease the swelling ratio. Moreover, with a fixed amount of glycogen, an increased concentration of collagen decreased the swelling rate. Additionally, the addition of nanohydroxyapatite decreased the swelling rate. Hydrogel network with less swelling

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indicates enhanced cross-linking [69]. The biodegradation of hydrogels plays a crucial role in tissue regeneration [70]. The differentiation of MSCs was induced by the proper mechanical strength of the hydrogel, which maintained solid morphology for several days. Moreover, the degradation of hydrogel is required to create enough space for the regeneration of bone tissue. The biodegradation study of the prepared hydrogel was performed in pH 7.4 phosphate buffer solution that simulates physiological environment. About 70% of the prepared hydrogel mass was remaining after 14 days of incubation in a pH 7.4 phosphate buffer solution. Therefore, the original appearance of the hydrogel was maintained during the attachment, proliferation, and differentiation of cells. The observed Young’s modulus and compressive modulus of hydrogels were within the range necessary for stem cell differentiation. Composite hydrogels loaded with mesenchymal stem cells exhibited desired cell growth, viability, and adhesion. The differentiation of bone mesenchymal stem cells into osteoblast or chondrocytes depends on the amount of glycogen, collagen, and hydroxyapatite present in the hydrogel. All of the results showed that the developed hydrogel with substantial mechanical and biological properties has huge potential as scaffolds for reconstruction of bone.

2.5.2 Wound healing and skin tissue regeneration Wound healing is a multifaceted procedure that comprises the matched interactions of chemokines, growth factors, various cells, and cytokines [71e74]. Failure of any process leads to scar formation, and wounds become chronic. Chronic wounds necessitate repetitive treatments, which enhance medical costs and disturb the quality of life. Conventional wound dressing materials are generally used, which provide a constructive environment for wound healing and control of wound infection. However, the effectiveness of these treatment modalities is limited. Recently, a tissue engineering approach has emerged to restore skin structure and improve the healing of wounds [75e78]. Hydrogel-based wound dressings have attracted extreme attention owing to their three-dimensional structure and similarity to skin tissue. Moreover, it can be applied as a liquid and converted to a hydrogel dressing at the wound site [79e82]. Novel biodegradable microfibrous hydrogels were prepared using glycogen as direct contact dressings/interface materials for wound healing [83]. Initially, the modification of glycogen was carried out through alkylation using allyl and propargyl bromides. The alkylation of glycogen’s eOH group was carried out in sodium hydroxide solution, and the resultant product allylated and propargylated glycogen (APG) was formed (Fig. 2.1). The existence of double and triple bonds is responsible for further cross-linking and functionalization. Then the sponges, like microfibers, were prepared by freeze drying aqueous solutions of APG. The adhesion and cell growth in the developed microfibers were studied using human osteoblast-like MG-63 cells. The culture dish containing glycogen-based materials exhibited a higher number of initially adhered cells in comparison to the polystyrene culture dish (control) one day after seeding. Similar findings were reported on the third and seventh days of the experiment (Fig. 2.2A). The glycogen-based materials showed significantly smaller cell spreading areas in comparison to polystyrene dish (control) throughout the investigation (Fig. 2.2B). The proliferation of cells diminished in culture for seven days with the glycogen sample. However, the shape of cells (polygonal or spindle like) in a glycogen-based culture dish is similar to that in a control culture dish. The cell numbers were augmented with increasing time and produced confluent layer at the end of the

FIGURE 2.1 Preparation of allyl and propargyl derivatives of glycogen. Reprinted from Rabyk M, Hruby M, Vetrik M, Kucka J, Proks V, Parizek M, Konefal R, Krist P, Chvatil D, Bacakova L, Slouf M, Stepanek P. Modified glycogen as construction material for functional biomimetic microfibers, Carbohydr Polym 2016;152:271e279, Copyright (2016), with permission from Elsevier.

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FIGURE 2.2 Number of human osteoblast-like MG 63 cells on Days 1, 3, and 7 after seeding on the bottom of polystyrene culture wells with or without the presence of the tested glycogen fibrous material (n ¼ 8) (A). Cell spreading area on the first day after seeding (n ¼ 47 and 77 for APG-0.5 and polystyrene, respectively) (B). Note: Analysis of variance, Student-Newman-Keuls method. *P  .001 in comparison with control polystyrene. Reprinted from Rabyk M, Hruby M, Vetrik M, Kucka J, Proks V, Parizek M, Konefal R, Krist P, Chvatil D, Bacakova L, Slouf M, Stepanek P. Modified glycogen as construction material for functional biomimetic microfibers. Carbohydr Polym, 2016;152:271e279, Copyright (2016), with permission from Elsevier.

experiment. The physiological growth of cells was observed in the existence of a glycogen-based sample devoid of evident cell damage. Next, the RGD-modified glycogen fibers were developed and tested for cell growth. The developed material exhibited enhanced cell growth. The RGD peptide is recognized by cellular receptors and enhances the cell’s affinity and adhesion [84]. Generally, RGD peptides are extensively used in biomedical fields for improving the strength and duration of cell adhesion [85]. The number of cells was higher in RGD-modified materials compared to pure polystyrene dishes, but no significant difference was observed. RGD-APG-0.5 culture dish showed a higher number of adhered cells in comparison to control dish under full exposure time (Fig. 2.3A). The number of adhered cells in the RGD-APG-5 dish was analogous to the control dish. However, the cell spreading area in the control dish was considerably higher compared to the RGDmodified glycogen dish on the first day after seeding (Fig. 2.3B).

FIGURE 2.3 Number of human osteoblast-like MG-63 cells on Days 1, 3, and 7 after seeding on polystyrene culture wells with or without RGDmodified glycogen fibrous material (n ¼ 20) (A). Cell spreading area on the first day after seeding (n ¼ 181, 152, and 46 for RGD-0.5, RGD-5 and polystyrene, respectively) (B). Note: Analysis of variance, StudenteNewmaneKeuls method. *P  .001 in comparison with control PS. Reprinted from Rabyk M, Hruby M, Vetrik M, Kucka J, Proks V, Parizek M, Konefal R, Krist P, Chvatil D, Bacakova L, Slouf M, Stepanek P. Modified glycogen as construction material for functional biomimetic microfibers. Carbohydr Polym 2016;152:271e279, Copyright (2016), with permission from Elsevier.

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2.5.3 Self-healing hydrogel in tissue engineering Because of the complex mechanical and physiological milieu inside the human body, some hydrogel defects were observed after implantation. Conventional hydrogel exhibited irreversible damage after the destruction and gradually increase to big fissures resulting failure of the material [86e88]. Therefore, self-healing hydrogel is needed to enhance the practical life of implanted hydrogel. It can heal and repair damage spontaneously. Moreover, premature failure was not observed after implantation due to mechanical damage. Self-healing of hydrogels was achieved by covalent and noncovalent interactions of hydrogel functional groups or by external stimuli including pH, ionic strength, light, and heat [89e91]. The self-healing hydrogel showed an extended life span since it has the ability to repair itself after being exposed to external forces [92,93]. Moreover, it improved the safety of the material by inhibiting failures produced by wear and tear. Hussain et al. created an independent self-healing hydrogel based on a dual physically linked (DPC) hydrogel by simply combining varying amounts of acrylic acid (AA) with glycogen (Gly) solution with continuously stirring, along with the initiator ammonium persulfate (APS) [37]. Finally, different quantities of ferric ions were introduced into these hydrogels to produce ionic cross-linking sites for coordination bonding with Gly and AA functional groups. Gly/PAAFe3þ is the name given to the resulting hydrogel. The hydrogels created have two forms of noncovalent sacrificial bond interactions. AA carboxylic groups and Gly hydroxyl groups form coordination links with ferric ions (Fe3þ), whereas Gly hydroxyl groups and AA carbonyl groups link via hydrogen bonds. By altering several chemical factors, the produced hydrogels displayed exceptional self-healing capabilities and adaptable mechanical properties. Hydrogel samples with definite dimensions were split into two pieces using a knife and subsequently assembled at their freshly sliced interfaces to study self-healing effectiveness. To prevent the hydrogels from evaporation, the samples were placed in a Petri dish and covered in a polyethylene sheet. Tensile strength evaluations of the healed samples were performed at ambient temperature after the specified time intervals. Fig. 2.4A depicts a distinctive tensile stress-strain curves of the whole and cut hydrogels. The cut samples demonstrated outstanding self-healing capabilities in regard to time, resuming the original shape of the uncut hydrogel by exhibiting analogous stress-strain behavior. The stress-strain profiles including all healing times closely coincided with the actual uncut hydrogel specimens, demonstrating that mechanical characteristics recovered with time. The tensile stress of the cut hydrogel samples increased with healing time and surpassed 0.36 MPa within 24 h. Fig. 2.4B depicts the comparable fracture tensile stress. The stress and strain percentage selfhealing effectiveness was assessed by relating the stress and strain intensity of the healed specimens to that of the original sample, and the findings are shown in Fig. 2.4C. After only 6 h of healing time at room temperature with no extrinsic stimulation or healing agent, hydrogel exhibited a 58% restoration in stress and a 72% restoration in strain strength. The self-healing efficacy enhanced with time, and after 24 h, it demonstrated an optimum restoration of 83% stress recovery and 93% strain elongation recovery compared to the initial uncut hydrogel specimen. Similarly, they recovered around 44% hardness within 6h in comparison to the uncut hydrogel. The toughness percent recovery of the healed hydrogels increased and reached approximately 72% toughness within 12 h after healing. Following 24 h, the recovered sample had 83% toughness compared to the original uncut hydrogel specimen (Fig. 2.4D), demonstrating that the Gly-PAA-Fe3þ hydrogel is highly recoverable and capable of self-healing through sacrificial bond interactions. The coupling between Fe3þ migration and polymer segment reconfigurations resulted in the self-healing represented in Fig. 2.5 and the self-healing behavior of the Gly-PAA-Fe3þ hydrogels. The immobilized Fe3þ ions significantly improved the mechanical strength of the hydrogel, while the free Fe3þ ions played an important part in the self-healing process. Following mechanical injury, the macromolecular chain breaks (PPA and Gly) resulted in the creation of reactive end groups (-COOH and -OH) on neighboring fracture surfaces, whereas free Fe3þ ions diffused toward the freshly cut interface (Fig. 2.5B), where they might form new metal-carboxylate complexes with Fe3þ to reconstruct the Gly-PAAFe3þ network when the fracture surfaces came into contact (Fig. 2.5C). As a result of the synergistic interactions of dynamical coordination bonds and hydrogen bonds, the segmental variants of entangled PAA and Gly chains had been further facilitated, culminating in network reorganizations and consequently encouraging self-healing behavior, as shown in Fig. 2.5D. In addition to self-healing potential, the Gly/PAA-Fe3þ hydrogel was reshapable, as illustrated in Fig. 2.5. A thin film was split into two halves with a razor to test its reshapability and self-recovery. For better viewing, the two sliced pieces were colored with two dissimilar dyes, rhodamine B and methylene blue. The cut pieces were assembled in an irregular pattern and squeezed between two glass plates, as shown in Fig. 2.5A. Following 10 min, they were withdrawn from the glass plates, and it was seen that they had self-healed and produced a new film that could survive without visible fracture. Following 10e20 min, the film was mounted without any visible damage, as seen in the images in Fig. 2.5B. The development of a fresh layer with no obvious cracks (Fig. 2.5B) demonstrated that the resulting Gly-PAA-Fe3þ hydrogels exhibited excellent self-recovery and reshapability. The breaking of sacrificial links caused the hydrogel network to

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FIGURE 2.4 (A) Strain-stress curves of original and healed hydrogels, (B) fracture stress of healed and original hydrogels, (C) stress and strain healing efficiency of healed hydrogels, and (D) toughness of original and healed hydrogels. Reprinted from Hussain I, Sayed, SM, Liu S, Oderinde O, Yao F, Fu G. Glycogen-based self-healing hydrogels with ultra-stretchable, flexible, and enhanced mechanical properties via sacrificial bond interactions. Int J BiolMacromole 2018;117:648e658, Copyright (2018), with permission from Elsevier.

deconstruct when it was cut into distinct parts. Following the reconnecting of the numerous chopped pieces, the mobilized free Fe3þ ions and polymer chains resulted in the creation of new sacrificial bonds through the diffusion process. It is appropriate for producing innovative soft materials with prospective applications in biomedical domains due to its straightforward technique of synthesis at moderate settings and inexpensive raw materials, as well as its great mechanical properties, high stretchability, and good self-healing efficiency. Hussain et al. created a glycogen-based usable hydrogel as a sensing element for detecting human mobility using a simple and unique method, emphasizing its potential mechanical characteristics, self-healing ability, and ionic conductivity [94]. The Gly chain was physically cross-linked with the PAA-co-acrylamide (AAM) chain in the hydrogel, and afterward, iron(III) (Fe3þ) ions were inserted into the 3D network system, resulting in ionic and metaleligand interactions with polymer chain functional groups. At 25 C, the proposed hydrogels demonstrated high mechanical strength and intrinsic self-healing capabilities without the application of any external chemicals. These hydrogels also displayed ionic conductivity as a result of the free mobility of metal ions and the dynamic character of polymer chains, which makes them an excellent choice for conductive and self-healing biomaterials. The Gly-elastic sensor’s nature allowed them to recognize different human movements, and they exhibited good sensing and exceptional recoverability characteristics, making them a suitable choice for use in wearable electronics and e-skin. As an example, the Gly-hydrogel was employed as a resistivitytype strain sensor, with great sensitivity and precise and constant identification of physiological signals such as the bending

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FIGURE 2.5 Schematic illustration of the self-healing process (A) shows cutting hydrogels in half with a knife, (B) shows active hydrogel components migrating to the cut surfaces, (C) shows sacrificial hydrogen and metal-ligand connections reforming, and (D) shows the healed hydrogel. Reprinted from Hussain I, Sayed, SM, Liu S, Oderinde O, Yao F, Fu G. Glycogen-based self-healing hydrogels with ultra-stretchable, flexible, and enhanced mechanical properties via sacrificial bond interactions. Int J Biol Macromole 2018;117:648e658, Copyright (2018), with permission from Elsevier.

of various human joints. Such hydrogels could directly transport electrons and transform electrical energy into mechanical or chemical energy. The mechanical characteristics, stretchability, and notch-insensitive characteristics of the hydrogels were also manually tested, and the results are shown in Fig. 2.6A, which demonstrated that a 200 mm strip length could be overextended greater than 2000 mm devoid of breaking and that after stress releasing, the hydrogel reverted to its original form. Subsequently, the hydrogel was knotted and overextended, and it was remarkable to discover that the knotted sample could likewise be extended to a length of 2000 mm devoid of fracture. This stretching resulted in approximately a tenfold augment in length (Fig. 2.6B). These hydrogels also exhibited ultrastretchability while possessing a notch, confirming their excellent notch insensitivity. A tiny opening was created in the center of the hydrogel specimen with a 5 mm syringe needle for the notch insensitivity experiment, and it was then stretched to determine its notch insensitivity. As illustrated in Fig. 2.6C, a notch in the middle with a diameter of 2 mm might be overextended to 40 mm devoid of fracture, indicating a 20-fold stretching and a 2000% increase in strain over its initial size. Furthermore, the hydrophilicity and flexibility of the Gly and PAA chains across the notched region allowed them to reorganize themselves, resulting in resistance to notching. The associations between the components of the developed hydrogel were strong enough to prevent the notch from breaking at 2000% strain, making them a good hydrogel material. These properties have made such hydrogels a promising candidate for bioelectrodes, bioactuators, and biosensors. This is a novel method for improving the self-healing of hydrogel materials by employing natural polymers as skeletons, which could be valuable for flexible electronics and e-skin purposes. In order to create a hydrogel with capability for self-healing and possessing adjustable mechanical properties, Hussain et al. introduced bonding interactions in a hydrogel made from a hybrid of polymers of natural and synthetic origin [38]. A synthetic polymer, that is, PVA, rendered flexibility and stretchability to the prepared hydrogels, while glycogen, a natural polymer, provided the required strength. The hydrogels were created at room temperature by the free radical polymerization method using AA as monomer and with APS as the initiator. Fe3þ ions were used as an ionic cross-linker in the

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FIGURE 2.6 Visual photos depicted (A) hydrogel stretchability and flexibility; (B) knotted hydrogel stretchability; and (C) Gly-hydrogel notch insensitivity. Reprinted from Hussain I, Ma X, Luo Y, Luo Z. Fabrication and characterization of glycogen-based elastic, self-healable, and conductive hydrogels as a wearable strain-sensor for flexible e-skin. Polymer 2020;210:122961, Copyright (2020), with permission from Elsevier.

polymeric mixture, which enhanced the mechanical strength and self-healing proficiency of the resultant hydrogels. The physical cross-linking of the polymeric chains to trivalent metal ions through noncovalent interactions led to the creation of hydrogels. By adjusting the synthetic conditions, such as glycogen (Gly) and PVA concentrations, as well as the metal ion (Fe3þ) concentration, it was possible to produce Gly-PVA/PAA-Fe3þ hydrogels which demonstrated customizable stretchability and mechanical behavior, as well as remarkable self-healing capacity. Sacrificial hydrogen bonds between functional groups of polymers and metal ions, as well as sacrificial coordination contacts of varying strength, result in adjustable sacrificial bonds. When external stress is applied, weaker sacrificial connections first break, dissipating the energy and giving the hydrogel adaptable mechanical and self-healing capabilities. The new hydrogel thus produced can be said to have a wider range of applications due to its outstanding self-healing efficiency and customizable mechanical properties. This theory was utilized by Hussain et al. in the preparation of the hybrid hydrogel. PVA assembly typically produces soft hydrogels with low mechanical qualities, which restricts their industrial applications. The fostering of ionic coordination during the addition of metal ions to the PVA hydrogel network increased the designed hydrogel’s mechanical properties. Furthermore, the introduction of the polymeric network of glycogen demonstrated a significant impact on the hydrogel network’s mechanical strength. By variations of PVA or glycogen concentrations, it was possible to tailor the mechanical properties and self-healing efficacy of the intended hydrogels. The mechanical strength of hydrogels typically rose with the addition of glycogen, but the stretchability and flexibility of the hydrogels increased with the increase in the concentration of PVA. The hydrogels demonstrated a 24-hour self-healing efficacy of roughly 98.78% without any outside assistance. They were thus found to be appropriate for the mass production of advanced soft materials for diverse applications in the biomedical domains owing to the simple and economical technique of synthesis, affordable raw materials, adequate mechanical properties, and excellent self-healing capability.

2.6 Conclusions As the methods of tissue engineering find application in the medical field, the search for hydrogel-forming materials intensified. Glycogen is a material, gifted with this property. It is ubiquitous in nature starting from mammalian tissues to simple plants and can also be synthesized from sucrose. Being a biomolecule, it is biodegradable, the most desired property in drug carriers. However, the polymer can stimulate the immune system which is the basis of its use for the activation of polymorphonuclear neutrophils. Research is on to develop glycogen as a functionally advanced material for drug delivery and tissue engineering through chemical modification.

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The effect of RGD density on osteoblast and endothelial cell behavior on RGD-grafted polyethylene terephthalate surfaces. Biomaterials 2009;30:711e20. [85] Wohlrab S, Muller S, Schmidt A, Neubauer S, Kessler H, Leal-Egana A, et al. Cell adhesion and proliferation on RGD-modified recombinant spider silk proteins. Biomaterials 2012;33:6650e9. [86] Guadagno L, Vertuccio L, Naddeo C, Calabrese E, Barra G, Raimondo M, et al. Self-healing epoxy nanocomposites via reversible hydrogen bonding. Compos B Eng 2019;157:1e13. [87] Andersen A, Krogsgaard M, Birkedal H. Mussel-inspired self-healing double-cross-linked hydrogels by controlled combination of metal coordination and covalent cross-linking. Biomacromolecules 2018;19:1402e9. [88] Wang XH, Song F, Xue J, Qian D, Wang XL, Wang YZ. Mechanically strong and tough hydrogels with excellent anti-fatigue, self-healing and reprocessing performance enabled by dynamic metal-coordination chemistry. Polymer 2018;153:637e42. [89] Miyamae K, Nakahata M, Takashima Y, Harada A. Self-healing, expansion-contraction, and shape-memory properties of a preorganized supramolecular hydrogel through host-guest interactions. Angew Chem Int Ed Engl 2015;54:8984e7. [90] Takashima Y, Yonekura K, Koyanagi K, Iwaso K, Nakahata M, Yamaguchi H, et al. Multifunctional stimuli-responsive supramolecular materials with stretching, coloring, and self-healing properties functionalized via hosteguest interactions. Macromolecules 2017;50:11.

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[91] Xia NN, Xiong XM, Rong MZ, Zhang MQ, Kong F. Self-healing of polymer in acidic water toward strength restoration through the synergistic effect of hydrophilic and hydrophobic interactions. ACS Appl Mater Interfaces 2017;9:37300e9. [92] He J, Shi M, Liang Y, Guo B. Conductive adhesive self-healing nanocomposite hydrogel wound dressing for photothermal therapy of infected fullthickness skin wounds. Chem Eng J 2020;394:124888. [93] Li S, Wang L, Zheng W, Yang G, Jiang X. Rapid fabrication of self-healing, conductive, and injectable gel as dressings for healing wounds in stretchable parts of the body. Adv Funct Mater 2020;30:2002370. [94] Hussain I, Ma X, Luo Y, Luo Z. Fabrication and characterization of glycogen-based elastic, self-healable, and conductive hydrogels as a wearable strain-sensor for flexible e-skin. Polymer 2020;210:122961.

Chapter 3

Hydrogel based on hyaluronic acid Roberta Cassano, Federica Curcio, Roberta Sole and Sonia Trombino Department of Pharmacy, Health and Nutritional Science, University of Calabria, Arcavacata, Italy

3.1 Introduction Hydrogels are three-dimensional materials, formed by cross-linked polymers, with a high affinity for water and biological fluids and capable of absorbing from 10% up to thousands of times their dry weight in water (Fig. 3.1) [1,2]. They can be developed using both natural and synthetic polymers and are active in biomedical areas, especially for drug delivery applications [3e11]. Hydrogels are generally produced with two cross-linking methods [12]: one physical and one chemical. Physical hydrogels, based on noncovalent interactions, are characterized by deformation, low mechanical strength, and poor stability [13]. Chemical hydrogels characterized by covalent bonds have better mechanical stability but are fragile and irreversible when the network is formed. Among the various polymers used for the formation of hydrogels, the natural ones, like the polysaccharides, are preferred [14,15]. They have been considered for the development of drug delivery systems (DDSs) not only for their high biocompatibility but also because they can be cross-linked with various molecules capable of improving the physical and chemical properties of DDS and to increase the bioavailability of carrier drugs (Fig. 3.2).

FIGURE 3.1 Hydrogel structure.

FIGURE 3.2 Hydrogel as drug delivery system. Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine. https://doi.org/10.1016/B978-0-323-95351-1.00007-7 Copyright © 2024 Elsevier Inc. All rights reserved.

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One relevant member of the polysaccharide biomaterials is hyaluronic acid (HA) [16]. It is a hydrophilic biopolymer composed of repeated units of b-1,4-glucuronic acid and b-1,3-N-acetyl-D-glucosamine (Fig. 3.3). Hyaluronic acid is present in almost all body fluids and tissues [17] and plays an important role in the human body, for example, in the extracellular matrix (ECM), it helps to maintain homeostasis and viscoelasticity thanks to its high molecular weight and its absorption capacity of a large amount of water [18]. Even at the intracellular level, it performs important functions being able to regulate thanks to its link with specific cell surface receptors, cell adhesion, migration, proliferation, and differentiation, thus slowing down processes such as inflammation, tumor progression, metastasis and promoting wound healing [19]. The purpose of this chapter is to discuss the performance of HA-based hydrogels for drug delivery with particular regard to dermatological, ocular, injectable, and inhalable applications.

3.2 Dermatological applications Since hyaluronic acid is an important component of the skin’s ECM, it is involved in various physiological functions of the skin. Thanks to its high affinity for water, it is able to maintain skin hydration and improve both its permeability and viscoelasticity [20,21] (Fig. 3.4). Furthermore, being equipped with carboxylic groups, the HA can ionize at pH 6e7 and release the drug carried by the electrostatic repulsion. Due to these properties, this material is also used in pH-sensitive DDSs. In this regard, Kim and coworkers have created both pH and temperature-sensitive hydrogel, consisting of poly(N-isopropylacrylamide) (PNIPAM) and HA, for the transdermal release of luteolin, a drug capable of inhibiting the hyperproliferation of keratinocytes in psoriasis. In vitro skin permeation experiments have shown that the obtained hydrogel effectively delivers luteolin to the epidermis and dermis, increasing the potential relief of psoriasis, making it an attractive carrier for transdermal release [21]. A transdermal delivery system was developed by Zhang et al. for the nonsteroidal antiinflammatory drug (NSAID), ibuprofen which, due to its short biological half-life, must be administered frequently orally and is highly irritating to the digestive tract. In this regard, a stable microemulsion was used as a carrier of the dispersed drug, in turn, in a hydrogel based on hyaluronic acid to increase the percutaneous absorption of the drug, thus bypassing the gastrointestinal tract. The transdermal release studies compared the hyaluronic acid-based microemulsion/hydrogel formulation with the commercial cream and hydrogel formulations and the results showed that the transdermal release of ibuprofen was significantly higher than that of the commercial hydrogel and cream. The cumulative amount permeated in 24 h was 1.42 times and 2.52 times

FIGURE 3.3 Hyaluronic acid structure.

FIGURE 3.4 Hyaluronic acid hydrogel and skin.

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that of the commercial hydrogel and cream, respectively. Furthermore, in vitro studies conducted on cell lines cellular keratinocytes (HaCaT) demonstrated the reduced cytotoxicity of the drug administered through the microemulsion/ hydrogel formulation based on hyaluronic acid toward these cells. Therefore, the results obtained by Zhang and coworkers indicate that microemulsion/hydrogel formulation can be an effective tool for the transdermal administration of ibuprofen [22]. Hasan et al. developed composite hydrogels using sodium alginate (SA), pectin (P), and HA loaded with clindamycin (cly), a semisynthetic derivative of lincomycin which exhibits antibacterial activity, and has been shown to be useful for the treatment of wounds infected with methicillin-resistant Staphylococcus aureus (MRSA) [23]. The composite hydrogels loaded with cly were evaluated for their physical properties such as pH, surface morphology, liquid absorption, etc. In vitro studies were conducted to evaluate cly release, bactericidal effects, and in vivo wound-healing activity in MRSA-induced wounds in mouse models. All hydrogels, including those based on HA, released cly within 12 h and also exhibited potent antibacterial activity against MRSA. In a mouse model, the hydrogel also significantly accelerated healing and reepithelialization of MRSAinfected wounds. Therefore, cly-loaded composite hydrogels could be an attractive strategy for treating infected skin wounds. Recently Yuan and coworkers have formulated a new carrier for indometacin (IND) [24], a NSAID, often used in the treatment of chronic pain and inflammatory conditions [25]. To improve IND skin permeation and reduce adverse effects that occur after oral administration, this drug was encapsulated in HA-modified transfersomes (IND-HT) which showed spherical structure, nanosize, and good encapsulation efficiency. Subsequently, with the aim to prolong the IND skin retention, the IND-HTs were incorporated into a Carbopol 940 hydrogel (IND-HTs/Gel). Instrumental analyzes revealed that HA enhanced the transdermal effect by modifying the microstructure of the skin layers and decreasing the skin barrier function. Furthermore, the gel obtained showed a remarkable analgesic activity and no irritation harmful to the skin. Hence IND-HT/Gel could be considered a promising noninvasive approach to improve skin permeation and transdermal drug release.

3.3 Ophthalmic applications Hyaluronic acid is also widely used in ophthalmic formulations both for its protective effect against damage caused by preservatives and for its pseudoplastic behavior which ensures good diffusion of the polymer on the ocular surface [26,27] (Fig. 3.5). Zhu and coauthors developed a heat-sensitive in situ gelling formulation based on PNIPAM, a stimulus sensitive polymer, and hyaluronic acid for the ophthalmic delivery of ketoconazole [28]. Ketoconazole is an antifungal and it was inserted into the gels, and the percentage incorporated was found to be between 91% and 96%. In vitro tests have shown a moderate and no bursting release of ketoconazole. In addition, in in vivo studies, the gels were found to be well tolerated and no signs of irritation or toxic effects were observed, while a good cure rate and growth inhibition of Candida albicans were found. Therefore, this new formulation could prove to be an interesting new ocular dosage form capable of prolonging the residence time of drugs administered in the eye. In 2021, Jeremy Chae and his colleagues designed and produced a hyaluronic acid hydrogel, useful in the treatment of glaucoma, to be injected in situ, through a microneedle, into the suprachoroidal space (SCS) of the eye [29]. The idea is that the expansion of SCS can increase the drainage of aqueous humor from the eye, resulting in the reduction of intraocular pressure (IOP). This should allow for a reduction in IOP for four months. Safety studies conducted by clinical ophthalmic examinations indicated that the treatment is well tolerated. Additionally, histopathological examinations

FIGURE 3.5 Administration of hyaluronic acid hydrogel in the eye.

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performed at the injection site showed only mild bleeding and fibrosis. While further analyzes, using ultrasound biomicroscopy, showed a strong correlation between IOP reduction and SCS expansion. Therefore, this hyaluronic acid-based hydrogel injection strategy could prove useful for treating ocular hypertension and glaucoma without resorting to conventional drugs or surgery. In the same year, Bao et al. have created a hydrogel film based on oxidized chitosan glycol/hyaluronic acid useful for the double ophthalmic administration of dexamethasone (Dex), a powerful glucocorticoid and levofloxacin (Lev), an antibiotic belonging to the class of third generation quinolones [30]. The goal was to obtain a system to be used in postoperative endophthalmitis, an inflammatory condition due to the invasion of bacteria into the eye during surgery. Using different degrees of oxidation of hyaluronic acid, different hydrogel films were obtained, and as the degree of oxidation increased, the swelling ratio of the hydrogel films decreased. Then the hydrogels showed a release pattern where the Lev was released quickly, followed by a slower release of the Dex. Furthermore, the obtained hydrogels have been shown to possess both a powerful ability to inhibit bacterial growth in different bacterial strains, and an effective antiinflammatory activity tested in RAW264 macrophages. Overall, the resulting hydrogel film could be a promising vehicle for the treatment of postoperative endophthalmitis. Recently Ilochonwu et al. developed an intravitreal in situ-forming hydrogel based on HA and polyethylene glycol (PEG) polymers potentially useful as an extended-release system for the bevacizumab monoclonal antibody [31]. After the intravitreal injection, through a 29G needle, the formulation underwent a solegel phase transition at the administration site with the achievement of an intraocular deposition system for a delayed release of bevacizumab. Swelling and degradation studies have indicated that the release of the antibody from the hydrogel can last as long as 400 days. A biological test showed that the released bevacizumab remained bioactive. Therefore, the HA and PEG hydrogel offers a high potential for the sustained release of therapeutic antibodies in the treatment of ocular diseases.

3.4 Hyaluronic acid injectable hydrogels In recent years, HA-based injectable drugs (Fig. 3.6) are proving very popular for drug delivery because, after their injection into the human body, their half-life is improved, with longer residence times than HA solutions [32]. The three-dimensional, highly hydrophilic, elastic, and deformable structure of the polymeric network, in fact, offers an extraordinary swelling capacity that allows the absorption and trapping of large quantities of water with the drugs dissolved in it, thus guaranteeing its subsequent release in specific targets therapeutics. In order to obtain biocompatible and antibacterial injectable hydrogels, Andrade del Olmo and collaborators prepared hyaluronic acid solutions cross-linked with divinyl sulfone (DVS). The hydrogels were loaded with acetylsalicylic acid (AAS) and showed antiinflammatory properties in vitro, while those loaded with the antibiotics cefuroxime (CFX), tetracycline (TCN), and amoxicillin (AMX) showed in vitro antibacterial efficacy against staph. aureus. The combined use of antibiotics and AAS showed a synergistic effect which effectively reduced the S. aureus population. The results showed that HA-DVS hydrogels loaded with antibiotics/AAS could be used effectively to fight S. aureus-borne infections [33]. As far as onco therapeutic drug release is concerned, an interesting strategy has been proposed by Fiorica and colleagues to maximize drug concentration in the area close to the tumor: the development of an injectable in situ formed hydrogel, obtained by reaction between an amino derivative of hyaluronic acid and functionalized b-cyclodextrins vinyl sulfone complexing doxorubicin [34]. In vitro experiments revealed that the hydrogel, being able to release the anticancer

FIGURE 3.6 Injectable hyaluronic acid hydrogel.

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drug in a prolonged manner could block the growth of colorectal cancer without causing cytotoxic side effects in other areas of the body such as the heart. Therefore, the resulting injectable hydrogel can be considered an interesting device useful for localized chemotherapy of solid tumors. Ma and colleagues prepared injectable hydrogels by self-cross-linking of hyaluronic acid aldehyde (HA-CHO) and hydrazide-modified poly(g-glutamic acid) (g-PGA-ADH) for proteins delivery [35]. These HA/g-PGA hydrogels showed good mechanical properties. For release studies, bovine serum albumin (BSA) was chosen as the model protein and these studies showed that it was released either by diffusion or relaxation. The obtained results indicated the possibility of using HA/g-PGA hydrogels as attractive strategies for the controlled release of proteins. The performance of an injectable hydrogel obtained from the self-assembly of hyaluronic acid and kappa-carrageenan gelling agent, useful for wound healing and the controlled administration of meropenem, a beta-lactam antibiotic belonging to the carbapenem class, has recently been designed and evaluated [36]. The hydrogel was characterized by nuclear magnetic resonance, Fourier transform infrared spectroscopy, thermogravimetric analysis, and scanning electron microscopy (SEM). Viscosity studies indicated the hydrogel’s ability to gel in situ and injectable at various temperatures. Meropenem was loaded into the hydrogel, and its release at pH 7.4 was found to be 96.12%. Antibacterial studies on Pseudomonas aeruginosa, S. aureus, and Escherichia coli with the meropenem-laden hydrogel showed higher zones of inhibition. In vivo studies in SpragueeDawley (SD) rats showed accelerated healing.

3.5 Inhalable hyaluronic acid hydrogels The inhaled route of administration has attracted great attention in recent years as the lungs can absorb drugs thanks to their remarkable permeability and the large surface area that allows for the systemic and local delivery of drugs [37e39]. It is also a noninvasive route of administration, which avoids first pass metabolism (Fig. 3.7). In this regard, Athamneh and collaborators designed aerogel hybrid microspheres based on alginate and hyaluronic acid potentially useful for the administration of pulmonary drugs, obtained using a gelation process of the emulsion followed by drying by supercritical CO2. The microspheres exhibited low density, high porosity, and good aerodynamic properties in vitro compatible with pulmonary drug release [40]. With the aim of prolonging the pharmacological effect of budesonide, a water-insoluble corticosteroid antiinflammatory drug, without affecting its percentage of dissolution, Liu and collaborators studied its pulmonary retention and pharmacokinetics, after loading it, in nanocrystalline particles prepared by wet ball grinding and subsequently inserted in inhalable mucoadhesive microparticles composed of hyaluronic acid [41]. All obtained particles were characterized by dynamic light scattering (DLS), laser diffraction, SEM, X-ray powder diffraction (XRD), and differential scanning calorimetry (DSC).

FIGURE 3.7 Inhalable hyaluronic acid hydrogels.

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In vivo studies were also conducted in rats to evaluate dissolution, release, aerosol performance, mucoadhesion, and pharmacokinetics. The inhaled microparticles showed a remarkable prolonged pharmacological effect compared to the suspension of budesonide nanocrystals suggesting that hyaluronic acid, thanks to its mucoadhesiveness, could be able to prolong the pharmacological effect of the active ingredient by extending its residence time in the lung. In 2021, Yan and colleagues fabricated a microgel using cross-linked carboxymethyl chitosan and hyaluronic acid. Zinc chloride (ZnCl2) was also used to promote precipitation [42]. The numerous properties of the microgel thus obtained make it suitable for administering drugs into the lungs. These properties include particle size (4 mm), a good degree of swelling, and porous structure. Furthermore, in vitro studies have demonstrated the potential ability of microgel to escape macrophage phagocytosis. Assessment of release behavior was performed using BSA as a model drug, and sustained release over 24 h was observed in vitro. The microgel remained intact in the mouse lung after injection for at least 36 h. The data obtained underline the potential of the microgel to favor local and prolonged pulmonary release. Nikjoo and colleagues proposed a new methodology for chemical cross-linking and spray drying the preparation of powders intended to the inhalation for sustained pulmonary drug delivery [43]. For the formulation of the hydrogel, various cross-linkers such as urea and glutaraldehyde have been used. The powders obtained were characterized. The analysis of the powders indicated a cross-linked structure of hydrogel with sufficient thermal stability to resist spray drying. The microparticles obtained had a spherical shape with an average particle diameter size between 2 and 3 mm, suitable for inhalation The particles formed by glutaraldehyde cross-linked HA e exhibited a more adequate aerosolization performance than those cross-inked with urea. Also swelling and stability in water were found to be greater for glutaraldehyde than for urea. The obtained results indicated that spray drying can be utilized to formulate inhalation powders of hyaluronic acid hydrogels as promising scaffolds for pulmonary sustained drug delivery.

3.6 Hyaluronic acid hydrogels and their applications in tissue engineering Hyaluronic acid hydrogels are one of most functional natural biomaterials in the field of tissue engineering not only due to their ability to adhere and cell proliferation but also to sufficient biological activity to stimulate a microenvironment for cell survival [44] (Fig. 3.8). FIGURE 3.8 Hyaluronic acid hydrogels applications in tissue engineering.

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3.7 Cartilage and bone regeneration Parck et al. have created an injectable hydrogel based on methacrylate glycol chitosan (MeGC) and hyaluronic acid, by means of photo cross-linking, using riboflavin as photoinitiator in visible light [45]. A minimum irradiation time of 40 s was required to produce stable gels useful for the encapsulation of chondrocytes, cartilage cells. Most of the chondrocytes encapsulated in MeGC-based hydrogel after 300 s of irradiation maintained a rounded shape with high cell viability for 21 days. The incorporation of hyaluronic acid into MeGC hydrogels increased the proliferation of chondrocytes and the deposition of the cartilage ECM. Thus, MeGC/HA composite hydrogels have been shown to be potentially useful in cartilage repair. Jin and coworkers produced an injectable epigallocatechin-3-gallate (EGCG) hydrogel to control inflammation and improve cartilage regeneration. The typical complication of osteoarthritis is one of the most common cartilage disorders. Having EGCG the property of modulating inflammation and inhibiting the formation of radical species, it was combined with HA conjugated with tyramine and gelatin to create a composite hydrogel with a concentration of 50 Mm EGCG and 5% w/v HA. In vitro studies have shown that the composite hydrogel thus obtained was able to protect the chondrocytes against the pro-inflammatory factor, IL-1b with chondrogenic regeneration. In vivo histological analysis, in a model of surgically induced osteoarthritis, showed that the EGCG-HA hybrid hydrogel containing gelatin can minimize cartilage loss. In order to create a hyaluronic acid hydrogel with excellent mechanical properties and, therefore, applicable in the field of cartilage tissue engineering (CTE), Zhao and collaborators have devised a strategy that integrates hydrogels and nanomaterials to form a system with mechanical properties suitable for the production and the recombination of cartilage tissue [46]. In particular, cellulose nanofibrils (CNFs) were chosen for their high mechanical strength and excellent biocompatibility. Therefore, functionalized CNFs in methacrylate suitable for photoreticulation with HA methacrylate to produce HA/CNF nanocomposite hydrogel have been made. The latter have adequate compressive strength (0.198  0.009 Mpa) and restorability, such as to make them useful as a load-bearing tissue for articular cartilage. Furthermore, this nanocomposite hydrogel could provide a good microenvironment for bone marrow mesenchymal stem cell proliferation and chondrogenic differentiation. Furthermore, in the rat models that had cartilage defects, the composite hydrogel showed an important repair effect in the full thickness defect model. Chen et al. fabricated an injectable adhesive hydrogel of aldehyde and methacrylate (AHAMA)-modified HA with multiple anchoring mechanisms such as amide bond through the dynamic reaction of the Schiff base, hydrogen bond, and physical interpenetration [47]. AHAMA hydrogel showed significantly improved durability and stability for at least seven days in a humid environment and higher adhesive strength (43 Kpa on skin and 52 Kpa on glass), compared to commercial fibrin glue (nearly 10 Kpa) and to the HAMA hydrogel (almost 20 Kpa). In addition, the AHAMA hydrogel was biocompatible and improved the proliferation (1.2-fold after three days) and migration (1.5-fold after 12 h) of bone marrow stem cells (BMSCs) in vitro. In a rat osteochondral defect model, the implanted AHAMA hydrogel significantly improved cartilage regeneration within weeks of implantation. Atoufi and collaborators created injectable heat-sensitive hydrogels based on PNIPAM and hyaluronic acid containing various amounts of PLGA micro/nanoparticles (ACH-PLGA) coated with chitosan-g-acrylic acid, to facilitate the regeneration of cartilage tissue [48]. The ACH-PLGA particles were inserted into the hydrogels for the following reasons: first of all, to improve the mechanical properties of the hydrogel. Then, the PLGA core acts as a vector for the controlled release of small molecules of chondrogenic melatonin. Furthermore, the particles could reduce the syneresis of the heatsensitive hydrogel during the gelling process. Finally, the adhesion tests showed the great ability of the hydrogel to integrate with natural cartilage. Thanks to improved mechanical properties, low syneresis, and sustained drug release capability, this injectable hydrogel is a promising material for cartilage regeneration. Whereas mesenchymal stem cells or hydrogel-supplied chondrocytes, after being implanted in damaged joints, could be exposed to high levels of reactive oxygen species (ROS) in the inflammatory microenvironment, resulting in inhibition of the regeneration process of cartilage tissue. To mitigate ROS-induced side effects, Shi and colleagues therefore developed a multifunctional antioxidant hydrogel. Hyaluronic acid was grafted, through a dynamic covalent bond, with phenylboronic acid (HA-PBA) and poly(vinyl alcohol) and was further stabilized by reaction with thiolated gelatin [49]. The obtained hydrogel was cytocompatible, injectable, and promoted cell adhesion and chondrogenic differentiation of mesenchymal stem cells. Furthermore, intra-articular injection of the hydrogel in mice revealed adequate stability and good biocompatibility in vivo. Therefore, this hydrogel can be applied for the regeneration of cartilage tissue in a chronic and elevated inflammatory ROS microenvironment. Ziadlou and coauthors designed and manufactured, via enzymatic cross-linking, injectable hydrogels for drug delivery and cartilage tissue repair by combining different concentrations of hyaluronic acid-tyramine (HA-Tyr) with silk fiber (SF)

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solutions regenerated [50]. To evaluate the influence of the hydrogel on the ECM deposition, bovine chondrocytes were incorporated into HA-Tyr/SF composites, HA/SF, or HA-Tyr. All hydrogel formulations were cytocompatible and promoted the expression of cartilage matrix proteins by allowing chondrocytes to produce ECM; the most important chondrogenic effects were observed in hydrogels with HA20/SF80 polymer ratios. The latter, loaded with anabolic and antiinflammatory drugs, also showed a longer and prolonged release profile over time, compatible with the typical duration of treatment for osteoarthritic joints and, therefore, suitable for CTE and, at the same time for the administration of drugs. Phan and collaborators also obtained injectable hydrogels, consisting of silk fibroin and hyaluronic acid, capable of restoring cartilage defects, with a minimally invasive surgery, characterized by a highly porous network and, therefore, useful for adhesion and cell proliferation and soft tissue regeneration [51]. In addition, methylprednisolone (MP) was inserted into the obtained hydrogel to relieve the inflammatory process. The SF/HA hydrogel scaffold was prepared by chemical cross-linking between the lysine residues of the silk fibroin, through the formation of Schiff bases. Compared to HA-free hydrogels, SF/HA hydrogels provide a more controlled release of MP. In vitro studies have also shown that SF/ HA hydrogels do not alter angiogenesis and blood vessel formation, making them suitable for cartilage regeneration and exhibiting controlled biodegradation. Overall, SF/HA hydrogels may look potential to be effective for regenerating articular cartilage lesions.

3.8 Wounds treatment Yang and collaborators have designed and manufactured hydrogels of bionic ECM based on poly(g-glutamic acid) modified with thiol (g-PGA-SH) and oxidized hyaluronic acid (HA-CHO) with properties of adaptability and ability to promote the initial coverage of the wound and prolongation of the duration of action of the dressing [52]. Furthermore, these hydrogels have been shown to possess viscoelastic characteristics like those of the natural ECM, the ability to degrade both in vitro and in vivo, and the scavenging activity of free radicals. Properties such as: the gelation time, the rheological and mechanical behavior, and the porous structure could be adjusted according to the hyaluronic acid content. Furthermore, in in vivo studies, they have shown that the obtained hydrogels were able to significantly improve the wound-healing process compared to the commercial dressing (Tegaderm), also favoring angiogenesis. To promote vascular cell growth and wound closure, Ying et al. have created an injectable hydrogel based on collagen and hyaluronic acid (COL-HA) simulating the ECM [53]. The preparation of the hydrogel (COL-HA) was carried out by in situ coupling of phenolic portions of collagen I-hydroxybenzoic acid (COL-P) and hyaluronic acid tyramine (HA-Tyr) by means of horseradish peroxidase (HRP). The hydrogel obtained has a porous structure that favors the exchange of gases, fluids, and nourishment. Endothelial cells and fibroblasts were grown within this hydrogel and showed important cell proliferation and possibility of vascular regeneration. Wound healing, treated with the COL-HA hydrogel, was greater than both the commercial formulation and the COL-P and HA-Tyr hydrogels alone. This latest result indicates the importance of coupling collagen and hyaluronic acid to improve wound healing.

3.9 Hyaluronic acid hydrogel and neuroregeneration Xu and coworkers prepared injectable chitosan (CS)-HA hydrogel, inserted into nerve ducts based on poly(D, L-lactic acid) (PDLLA)/b-tricalcium phosphate (b-TCP), for the prolonged release of nerve growth factor (NGF) and the consequent repair of damaged peripheral nerves in rats [54]. All CS-HA hydrogels have porosities greater than 80%, and their mass loss reached 70% in eight weeks. Furthermore, after in vivo studies, they proved to be suitable for the adhesion, diffusion, and differentiation of neuronal cells. Further studies on sciatic nerve defect repair have shown that the CS-HA/NGF hydrogel-loaded PDLLA/b-TCP nerve ducts have a significant positive effect on axon regeneration and myelination, compared to the same empty nerve ducts. These results show that CS-HA/NGF injectable hydrogel can effectively promote nerve regeneration. Angiogenesis plays an important role in brain injury repair as it contributes to axonal regeneration in the injury area. Since hyaluronic acid is the main component of the brain’s ECM, Jiaju Lu and collaborators prepared a hydrogel by modifying hyaluronic acid with a VEGF mimetic peptide from KLT (KLTWQELYQLKYKGI) [55]. The hydrogel obtained showed a porous and three-dimensional structure, with a large specific surface compatible with cell adhesion and interaction. Compared with unmodified hyaluronic acid hydrogel, HA-KLT could effectively promote endothelial cell attachment, spread, and proliferation. Furthermore, their pro-angiogenic efficacy was evaluated in vivo, after being implanted in the brain of an injured rat. HA-KLT hydrogels could potentially be useful in repairing brain defects, promoting angiogenesis and inhibiting scar tissue formation.

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3.10 Hyaluronic acid hydrogel and stem cells HA-based hydrogels also represent a promising strategy in stem cell delivery that have numerous potentials for the repairation of damaged or defective organs (Fig. 3.9). They, in fact, can restore the normal functions of a diseased tissue, avoiding its replacement. Normally, in therapeutic stem cell treatments, damaged cells are replaced or restored by regulating inflammation and the immune system. The low survival rate and local retention of transplanted cells makes necessary the development of biocompatible materials capable of improving the efficacy of stem cells and, in this sense, hyaluronic acid-based hydrogels meet the desired requirements [56]. Dong et al. have developed an advanced approach to delivering adipose-derived (ASC) stem cells, for the treatment of burn wounds, using an in situ-formed hydrogel system composed of hyperbranched poly(ethylene glycol) diacrylate (HBPEGDA) polymer, a thiol-functionalized hyaluronic acid (HA-SH) and a short RGD peptide [57]. The combination with the RGD peptide to increase cell adhesion not only improved cell proliferation but also increased the paracrine activity of angiogenesis and growth factors. Hydrogel-ASC treatment significantly improved neovascularization, accelerated wound closure, and reduced scar formation. Hamilton and collaborators developed a new microencapsulating formulation based on degradable hyaluronic acid acrylate (AHA) as a cell-controlled release vehicle of stem cells. These microspheres were compared with those based on poly(ethylene glycol) diacrylate (PEGDA), a durable hydrogel [58]. AHA-based microspheres have shown to possess higher swelling ratio, faster degradation rate, lower retention modulus, and larger mean diameter than PEGDA-composed microspheres. Additionally, in vitro cell viability and release and short-term in vivo biocompatibility studies, using SpragueeDawley rats, revealed that, compared to PEGDA, AHA-composed microspheres resulted in a significantly lower in vivo. foreign body response. Fat-derived stem cells (ADSCs) show potential skin regeneration. In this context, stem cell therapies are showing encouraging results in accelerating tissue regeneration. Unfortunately, the low survival rate of transplanted cells, caused by the lack of protection during and after transplantation, negatively affects efficacy. To overcome this problem, Gong et al. obtained a dopamine-methacrylate hyaluronic acid (DA-MeHA) [59]. In particular, the authors first covalently cross-linked the hyaluronic acid with methacrylic anhydride, and, in a second step, the obtained product was again covalently crosslinked with dopamine. The obtained DA-MeHA hydrogel adhered firmly to the skin wound defect and promoted both cell proliferation in vitro and regeneration of the skin defect in vivo due to the high levels of growth factors secreted by the ADSCs of the DA-MeHA hydrogel. In order to improve the efficiency of mesenchymal stem cells for the treatment of inflammatory lesions related to atopic dermatitis, Lee et al. have recently developed injectable self-cross-linkable hydrogels using thiol-functionalized hyaluronic acid (HA-SH) [60]. The gelling ability and the mechanical properties of the obtained hydrogels were easily adjusted by varying the concentration of the polymer in the precursor solution before injection. HA-SH hydrogels loaded with mesenchymal stem cells showed high cell viability (>80%) for seven days and good biocompatibility in vivo after subcutaneous implantation in mouse skin. There was also increased expression of antiinflammatory cytokines, which can alleviate the immune response. In an animal model of atopic dermatitis, a reduction in epidermal thickness and mast cell infiltration was achieved by applying an HA-SH solution loaded with mesenchymal stem cells. This HA-based injectable hydrogel represents a potential stem cell carrier and has great potential in treating inflammatory diseases up to inflammation.

FIGURE 3.9 Hyaluronic acid hydrogel and stem cells.

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Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine

3.11 Conclusions In recent decades, the development of biopolymer-based hydrogels, hyaluronic acid-based hydrogels, applicable in prolonged drug release and tissue engineering, is becoming increasingly important. These materials are, in the controlled release of drugs, useful because they allow for dose control, local administration, and the reduction of side effects with a consequent increase in the efficacy and safety of drugs. Regarding the employment of hyaluronic acid hydrogels in the field of tissue engineering, they are particularly interesting due to their ability to adhere and to induce cell proliferation but also to their biological activity allowing for the stimulation of a microenvironment supporting cell survival. In this context, this chapter was intended to highlight and discuss recent advances in the design of chemical and physical products based on hyaluronic acid hydrogels useful for the controlled temporal and spatial specific release of drugs and for tissue regeneration.

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[26] Aragona P, Papa V, Micali A, Santocono M, Milazzo G. Long term treatment with sodium hyalu-ronate-containing artificial tears reduces ocular surface damage in patients with dry eye. Br J Ophthalmol 2002;86:181e4. [27] Zhang X, Wei D, Xu Y, Zhu O. Gel-based materials for ophthalmic drug delivery. Gels 2021;7:130e45. [28] Zhu M, Wang J, Li N. A novel thermo-sensitive hydrogel-based on poly(N isopropylacrylamide)/hyaluronic acid of ketoconazole for ophthalmic delivery. Artif Cell Nanomed Biotechnol 2018;46(6):1282e7. [29] Chae JJ, Jung JH, Zhu W, Gerberich BG, Bahrani Fard MR, Grossniklaus HE, et al. Drug-free, nonsurgical reduction of intraocular pressure for four months after suprachoroidal injection of hyaluronic acid hydrogel. Adv Sci 2020;2:2001908e20019020. [30] Bao Z, Yu A, Shi H, Hu Y, Jin B, Lin D, et al. Glycol chitosan/oxidized hyaluronic acid hydrogel film for topical ocular delivery of dexamethasone and levofloxacin. Int J Biol Macromol 2021;167:659e66. [31] Blessing C, Ilochonwu M, Mihajlovic RF, Rousou C, Tang M, et al. Hyaluronic acid-PEG-based diels-alder in situ forming hydrogels for sustained intraocular delivery of bevacizumab. Biomacromolecules 2022;23:2914e29. [32] Andrade Del Olmo J, Pérez-Álvarez L, Sáez Martínez V, Benito Cid S, Pérez González R, Vilas-Vilela JL, et al. Drug delivery from hyaluronic acidBDDE injectable hydrogels for antibacterial and anti-inflammatory applications. Gels 2022;8:223e43. [33] Andrade Del Olmo J, Alonso JM, Sáez Martínez V, Ruiz-Rubio L, Pérez González R, et al. Biocompatible hyaluronic acid-divinyl sulfone injectable hydrogels for sustained drug release with enhanced antibacterial properties against Staphylococcus aureus. Mater Sci Eng C 2021;125:112102e1121020. [34] Fiorica C, Palumbo FS, Pitarresi G, Puleio R, Condorelli L, Collura G, et al. A hyaluronic acid/cyclodextrin based injectable hydrogel for local doxorubicin delivery to solid tumors. Int J Pharm 2020;589:119879e89. [35] Ma X, Xu T, Chen W, Qin H, Chi B, Ye Z. Injectable hydrogels based on the hyaluronic acid and poly (g-glutamic acid) for controlled protein delivery. Carbohydr Polym 2018;179:100e9. [36] Ijaz U, Sohail M, Usman Minhas M, Khan S, Hussain Z, Kazi M, et al. Biofunctional hyaluronic acid/k-carrageenan injectable hydrogels for improved drug delivery and wound healing. Polymers 2022;14:376e86. [37] Liang Z, Ni R, Zhou J, Mao S. Recent advances in controlled pulmonary drug delivery. Drug Discov Today 2015;20:380e9. [38] Loira-Pastoriza C, Todoroff J, Vanbever R. Delivery strategies for sustained drug release in the lungs. Adv Drug Deliv Rev 2014;75:81e91. [39] Sheth P, Myrdal PB. Polymers for pulmonary drug delivery. In: Smyth HD, Hickey AJ, editors. Controlled pulmonary drug delivery. 2nd ed. Germany: Springer: Berlin/Heidelberg; 2011. p. 265e82. [40] Athamneh T, Amin A, Benke E, Ambrus R, Leopold CS, Gurikov P, et al. Alginate and hybrid alginate-hyaluronic acid aerogel microspheres as potential carrier for pulmonary drug delivery. J Supercrit Fluids 2019;150:49e55. [41] Liu T, Han M, Tian F, Cun D, Rantanen J, Yang M. Budesonide nanocrystal-loaded hyaluronic acid microparticles for inhalation: In vitro and in vivo evaluation. Carbohydr Polym 2018;181:1143e52. [42] Yan Y, Wu Q, Ren P, Liu Q, Zhang N, Ji Y, et al. Zinc ions coordinated carboxymethyl chitosan-hyaluronic acid microgel for pulmonary drug delivery. Int J Biol Macromol 2021;193(Part B):1043e9. [43] Nikjoo D, van der Zwaan I, Brülls M, Tehler U, Frenning G. Hyaluronic acid hydrogels for controlled pulmonary drug delivery-A particle engineering approach. Pharmaceutics 2021;13:1878e98. [44] Li H, Qi Z, Zheng S, Chang Y, Kong W, Fu C, Yu Z, Yang X, Pan S. The application of hyaluronic acid-based hydrogels in bone and cartilage tissue engineering. Adv Mater Sci Eng 2019;2019:12. Article ID 3027303. [45] Park H, Choi B, Hu J, Lee M. Injectable chitosan hyaluronic acid hydrogels for cartilage tissue engineering. Acta Biomater 2013;9:4779e86. [46] Zhao H, Zhang Y, Liu Y, Zheng P, Gao T, Cao Y, et al. In situ forming cellulose nanofibril-reinforced hyaluronic acid hydrogel for cartilage regeneration. Biomacromolecules 2021;22(12):5097e107. [47] Chen J, Yang J, Wang L, Zhang X, Heng BC, Wang D, et al. Modified hyaluronic acid hydrogels with chemical groups that facilitate adhesion to host tissues enhance cartilage regeneration. Bioact Mater 2021;6(6):1689e98. [48] Atoufi Z, Kamrava SK, Davachi SM, Hassanabadi M, Garakani SS, Alizadeh R, et al. Injectable PNIPAM/Hyaluronic acid hydrogels containing multipurpose modified particles for cartilage tissue engineering: synthesis, characterization, drug release and cell culture study. Int J Biol Macromol 2019;139:1168e81. [49] Shi W, Fang F, Kong Y, Greer SE, Kuss M, Liu B, Xue1 W, Jiang X, Lovell P, Mohs AM, Dudley AT, Li T, Duan B. Dynamic hyaluronic acid hydrogel with covalent linked gelatin as an anti-oxidative bioink for cartilage tissue engineering. Biofabrication 2022;14:014107e26. [50] Ziadlou R, Rotman S, Teuschl A, Salzer E, Barbero A, Martin I, et al. Optimization of hyaluronic acid-tyramine/silk-fibroin composite hydrogels for cartilage tissue engineering and delivery of anti-inflammatory and anabolic drugs. Mater Sci Eng C 2021;120:111701. [51] Phan VH, Murugesan M, Nguyen PPT, Luu CH, Le NH, Nguyen HT, et al. Biomimetic injectable hydrogel based on silk fibroin/hyaluronic acid embedded with methylprednisolone for cartilage regeneration. Colloids Surf B Biointerfaces 2022;219:11285. [52] Yang R, Liu X, Ren Y, Xue W, Liu S, Wang P, et al. Injectable adaptive self-healing hyaluronic acid/poly (g-glutamic acid) hydrogel for cutaneous wound healing. Acta Biomater 2021;127:102e15. [53] Ying H, Zhou J, Wang M, Su D, Ma Q, Lv G, et al. 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[56] Fang Y, Shi L, Duan Z, Rohani S. Hyaluronic acid hydrogels, as a biological macromolecule-based platform for stem cells delivery and their fate control: a review. Int J Biol Macromol 2021;189:554e66. [57] Dong Y, Cui M, Qu J, Wang X, Kwon SH, Barrera J, et al. Conformable hyaluronic acid hydrogel delivers adipose-derived stem cells and promotes regeneration of burn injury. Acta Biomater 2020;108:56e66. [58] Hamilton M, Harrington S, Dhar P, Stehno-Bittel L. Hyaluronic acid hydrogel microspheres for slow release stem cell delivery. ACS Biomater Sci Eng 2021;7(8):3754e63. [59] Gong M, Yan F, Yu L, Furong L. A dopamine-methacrylated hyaluronic acid hydrogel as an effective carrier for stem cells in skin regeneration therapy. Cell Death Dis 2022;13:738e49. [60] Lee H, Lee TW, Chandrasekharan A, Sung SE, Yim SG, Kim S, Seong KY, Seo MS, Yang SY. Injectable self- crosslinkable thiolated hyaluronic acid for stem cell therapy of atopic dermatitis. ACS Biomater Sci Eng 2022;11(8):1613e22.

Chapter 4

Hydrogels based on chitosan Sujit Kumar Debnath1, Monalisha Debnath1, Rohit Srivastava1 and Abdelwahab Omri2 Department of Biosciences and Bioengineering, Indian Institute of Technology Bombay, Mumbai, Maharashtra, India; 2The Novel Drug & Vaccine

1

Delivery Systems Facility, Department of Chemistry and Biochemistry, Laurentian University, Sudbury, ON, Canada

4.1 Introduction Hydrogel or aquagel, a polymer matrix, draws significant interest in biomedical fields due to optimum biocompatibility, retention ability, permeability, and desired release kinetics. Hydrogel is a three-dimensional (3D) macromolecular polymer system containing hydrophilic chains where water is present as a dispersion medium. These polymer matrices seem soft tissue due to high water contentment, low modulus, and elasticity. Hydrogels are stable upon swelling in water and absorb a lot of water 10e1000 times its volume [1]. Hydrogels possess many applications, specifically in drug delivery and tissue engineering [2]. These materials can crosslink with different substances, including small molecules, proteins, nucleic acids, etc. These materials have demonstrated several characteristics like controllable degradation, protect labile drugs, physicochemical tolerability, etc. Due to their versatile properties, hydrogels are successfully employed in many local drug delivery systems. In this contrast, thermoplastic and hydrogel-based scaffolds have drawn significant interest due to control release of drug delivery, safe implantation, and degradation. These materials can be administered in different routes like enteral, parenteral, in situ implantation, topical, and ocular [3].

4.2 Classification of hydrogel These can be broadly categorized into the synthetic and natural-based hydrogel. Synthetic-based hydrogel has demonstrated some disadvantages like inflammatory reactions, difficulty in removing unreacted precursors, material migration, etc. Therefore, much attention has been given to biodegradable hydrogels prepared with natural polymers that undergo enzymatic degradation [4]. Natural hydrogel mainly contains polysaccharides or protein chains. The presence of the hydrophilic group in the polysaccharides is a crucial element in preparing hydrogel. Many raw materials like cellulose, alginate, chitosan, chitin, hyaluronic acid, pectin, xanthan gum, starch, and dextran are used in the hydrogel preparation. Collagen, gelatin, resilin, elastin, keratin, and silk are examples of protein chains responsible for the formation of hydrogel lattices. Synthetic polymers like polyvinyl alcohol (PVA), polyethylene glycol, polyethylene oxide, and polyacrylamide are employed for hydrogel formulation. These synthetic polymers have hydrolyzable moieties with a slower degradation rate, resulting in more robust hydrogel than natural polymers. Natural polymer undergoes enzymatic degradation and produces biocompatible byproducts. Hence, natural polymer shows higher biodegradability than synthetic polymers [5]. Hydrogels can be temperature-sensitive. Hydrogels can be categorized as positive and negative thermos-sensitive hydrogels based on their behavior toward their microenvironment. Negative thermosensitive hydrogels (prepared of poly(N-isopropyl acrylamide), poly(N,N-dimethyl acrylamide), poly(N-ethyl methacrylamide), poly(methyl vinyl ether), poly(N-vinyl caprolactam), etc.) show lower critical solution temperature. Positive thermos-sensitive hydrogels (polyacrylic acid [PAA], poly(acrylamide-co-butyl methacrylate), polyacrylamide, pluronic, tetronics, etc.) show upper critical solution temperature. Hydrogels also can be categorized based on crosslinking, crosslinked junctions, and electric charge (Fig. 4.1).

Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine. https://doi.org/10.1016/B978-0-323-95351-1.00010-7 Copyright © 2024 Elsevier Inc. All rights reserved.

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FIGURE 4.1 Classification of hydrogel in different prospects.

4.2.1 Crosslinking This class of hydrogels is classified into two classes: chemical and physically crosslinked hydrogels. Chemically crosslinked hydrogels demonstrate permanent linkages, whereas the formation of linkage in physically crosslinked hydrogels developed naturally due to several interactions in the polymer chain. Physical crosslinking makes hydrogel temporarily exist and reversible. Polymers like PVA/starch, PVA/chitosan, and PVA/gelatin are assembled by freeze-thawing. The nonhomogeneity in this type mainly occurred due to free chain ends and chain loops. Crosslinking happens with hyaluronic acid, alginate, chitosan, and cyclodextrin-based hydrogen through H-bonding [6]. Chemically crosslinked hydrogels are formed through polymerizing end-functionalized macromeres and covalent crosslinking. Hydrogel’s ability increases after absorbing water into the hydrated porous structure. This crosslinking can also be seen in polysaccharide-based hydrogels. Chemically crosslinked chitosan hydrogel loaded with gelatin was developed to encapsulate chondrocytes [7]. Loaded gelatin molecules into this chitosan derivative hydrogel polymerized to create a semi-interpenetrating (semiIPN) polymer network. The more gelatin content decreased the pore size and collapse extent in the dry state of hydrogels. Additionally, the swelling ratio was reduced, and hydrogel strength was improved due to the gelation. In most cases, physically crosslinking is the most promising technique than chemically crosslinking in preparing the hydrogels due to nontoxicity.

4.2.2 Electric charge Electric charge can be located on the entire crosslinked chain. Hydrogels can be categorized into ionic, nonionic, and ampholytic hydrogels based on the electric charge. The ionic hydrogel can be of two types: anionic and cationic. Cationic hydrogels can conjugate with other anionic charged molecules and exhibit pH-responsive gelation. These hydrogels successfully deliver genes and are also considered alternative vectors to viruses. Anionic hydrogels contain negative ions that bind with the polymer network. In contrast, neutral hydrogels have an equal number of positive and negative ions distributed throughout the polymer matrix. The water absorptivity in hydrogels depends on the inherent elastic restoring force and osmotic pressure of hydrogels [8,9]. Chitosan-based hydrogels are an ecofriendly drug delivery system. The impact of dual ionic and covalent crosslinks was analyzed on the chitosan-based hydrogels [10]. Diiodo-trehalose derivatives and citric acid were employed as covalent and ionic crosslinkers, respectively. These hydrogels were disintegrated entirely within 96 h through the hydrolysis process of enzyme trehalase. Maximum elastic properties were achieved at high CTS concentrations and a high degree of covalent and ionic crosslinks.

4.2.3 Crosslinked junctions Hydrogels can be blended, polyelectrolyte complex, and block copolymeric hydrogels. Blend hydrogels are denoted as blending hydrogels into new materials with distinct physical properties. These hydrogels possess superior mechanical, thermal, and electrical properties to conventional hydrogels. Block copolymeric hydrogels are prepared by blocking different polymerized monomers to form 3D polymeric frames that can soak a high volume of water to retain their semisolid structure. These are successfully employed in designing nanostructured materials to reduce drug toxicity, enhance drug efficacy, and long-term therapy. These hydrogels contain repeating units of electrolyte groups, either

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polycations or polyanions. The interpolymer complexes are formed between polyions using hydrogen bonding or electrostatic interaction. For example, macroscopic and polyelectrolyte multilayers (PEMs) hydrogels were prepared by electrostatic interactions between hyaluronic acid and chitosan. This PEM was fabricated by a layer-by-layer model onto a poly(ethylene terephthalate) (PET) surface. Due to the electrostatic interactions, this material showed intrinsic self-healing ability.

4.3 Hydrogel preparation using chitosan In hydrogel preparation, polymer networks are crosslinked by chemical or physical interaction. Chemical crosslinked hydrogels are prepared by free-radical, thermal, or photopolymerization. Whereas physically crosslinked networks are formed with noncovalent interaction at the molecular level. Various properties of hydrogels depend on the crosslinking density and crosslinking ratio between crosslink molecules and polymer repeating units. Altering these parameters will modify the swelling characteristics of the hydrogel. With increasing crosslinking, hydrogels become more compact, resulting in lesser swelling. In contrast, soft polymeric systems having more swelling capability are produced with low crosslinking ratio. Therefore, altering the crosslinking patterns can easily manipulate the several properties of hydrogels [6].

4.3.1 Chemical crosslinking Chemical crosslinking hydrogels are prepared with reversible covalent bonding between chitosan macromeres (Fig. 4.2). They can be formed in many ways: semi-IPN polymer networks, interpenetrating polymer networks, hybrid polymer networks, and chitosan crosslinked systems (Table 4.1). Amine and hydroxyl groups present in chitosan chains are involved in such chemical crosslinking.

4.3.1.1 Using crosslinkers Crosslinking happens due to the reaction between chitosan and crosslinking agents (CAs) leading to biopolymer preservatives. Mostly, covalent or ionic bonding is formed. Covalent CAs have two functional groups participating in the condensation reaction. The most effective CAs is glutaraldehyde (GTA) (dialdehyde). The aldehyde of this compound takes part in the reaction with primary amine groups of chitosan, resulting in the formation of Shiff’s bases [18]. Crosslinkers initiate a crosslinking reaction in the chitosan chains. This crosslinking can happen between the same polymer or different polymers with a crosslinker. Some such crosslinkers are acrylic acid, di-isocyanate, palladium cation, genipin, GTA, etc. PAA is a synthetic polymer formed with the polymerization of acrylic acid monomers. It is a high molecular weight polymer with good water solubility. This material exists in crosslinking form and is employed to fabricate

FIGURE 4.2 Different chemical interactions in the chitosan hydrogel network: (A) free-radical, (B) condensation, and (C) addition reactions [11].

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Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine

TABLE 4.1 Chitosan chemical hydrogel preparation and potential application. Type

Crosslinkers

Purpose

Application

Reference

Using crosslinkers

Poly(acrylic acid)

Assessment of electrochemical signal of Nile Blue (NB) and performance of NB delivery

1. Prepared hydrogel was suitable for the delivery of hydrophilic drugs.

[12]

Glutaraldehyde (GTA)

Coupling with GTA with covalent bond during enzyme immobilization

Glucose oxidase immobilization showed good catalytic activity for glucose oxidation

[13]

Genipin

To enhance the crosslinking and preserve the thermosensitive character

1. Less time took to form gel 2. Localized at the site of injection for more than one week

[14]

Maleic chitosan and TPEG

Improving the application of chitosan hydrogel in 3D printing

1. Improved gelling rate by w2 fold 2. Improved compressive strength by w10 fold 3. Improved scaffold stability

[15]

Synthetic poly(ethylene glycol) diacrylate

Improving the hardness of chitosan hydrogel

1. Improved mechanical and biochemical properties 2. Successfully employed in the fabrication of 3D scaffold

[16]

Methacrylate glycol chitosan, semiinterpenetrating collagen, riboflavin photoinitiator

Feasibility of photopolymerizable chitosan under visible light to sustain cell viability and promote proliferation

1. Enhanced the comprehensive modulus 2. Slow down the rate of degradation

[17]

Using photopolymerization

TPEG, thiol-terminated poly(ethylene glycol).

polymeric blends and nanocomposites [19]. Biodegradable and biocompatible pH-sensitive PAA-grafted chitosan hydrogel was prepared by inducing the acrylic acid monomer onto chitosan [12]. Nil blue was incorporated into this gel that showed a high loading yield with millions of signals. This nil blue-loaded hydrogel can be used in the early detection of cancer. GTA can conjugate/react with many groups of proteins like amine, imidazole, phenol, and thiol. It was observed that aldehyde could react with amino acids and increase the order of reactive moieties such as hydroxyl, secondary amino, guanidine, a-amino, and e-amino. Therefore, GTA can easily interact with amino groups and be an ultimate option for enzyme immobilization [13,20]. Chitosan hydrogels are frequently crosslinked with GTA, the most effective CA. This modified chitosan matrix has several biomedical applications. However, GTA (a powerful sterilizing agent) shows high cytotoxicity. It was observed that GTA at the concentration of 3.0 ppm inhibits fibroblast growth in tissue culture at 99%. Therefore, eliminating GTA after crosslinking is highly desirable [21]. Genipin is a bifunctional and naturally occurring water-soluble CA. It is obtained by hydrolyzing geniposide from the fruits of Gardenia Ellis. This compound reacts with amino groups, proteins, and amines of chitosan. It is used as reticulating agent due to its lower toxic profile than other covalent crosslinkers (1000 times lower than GTA). It can bind with a primary amine in chitosan and is successfully used in tissue engineering [22]. Using genipin-induced covalent and coupled ionic crosslinking, hydrogels based on chitosan were synthesized with varying genipin amounts [14,23]. This chemical CA was added to the neutralized chitosan solution (glycerol-phosphate complex). Likewise, the pH barrier of chitosan solution can be overcome, and its thermosensitive characteristics can be preserved. Genipin induces a chemical crosslinking process. Concurrently, this compound neutralized chitosan solution resulting in disturbance in the ionic crosslinking process. Other suitable CAs have epoxide groups like ethylene glycol diglycidyl ether and epichlorohydrin. Disruption of the epoxide ring of these compounds

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undergoes condensation reaction with chitosan’s hydroxyl groups. Generally, covalent crosslinking is an irreversible reaction and is stable in the acidic solution [18]. In addition to a covalent bond, electrovalent (ionic) bonds between chitosan and the CA are also formed. In such bonding, the charge of the ionic chitosan should be opposite to the modified polymer. The induction of ionic crosslinking effects happens due to the electrostatic interaction between CAs with polymer. These CAs should have at least two 3

functional groups that produce a negative charge in water like eO-PO2
3 , eSO , eCOO . The most commonly used ionic CAs are polyphosphates and citrates. This ionic crosslinking is less stable than the covalent bond. In a study, pluronic F-127 is widely used to develop a controlled release system due to lower toxicity and immunogenicity. The drug release from this polymer matrix is shortened in the aqueous solution. The release profile can be modulated when pluronic F-127 is added with chitosan. The physical mixture of chitosan and pluronic F-127 is extensively explored in different routes like ocular, intravesical, vaginal, intranasal, etc. [24].

4.3.1.2 Using photopolymerization Photopolymerizable hydrogels are obtained polymer from natural origin and become well accepted in tissue engineering due to superior biocompatibility, the proliferation of cell growth, and the ability to cure in situ with a minimally invasive procedure. Crosslinked chitosan hydrogels can also be prepared by photopolymerization. In this process, precursor solution converts into gel upon photoinitiators (ultraviolet [UV], or visible irradiation). This system is helpful for both in vitro and in vivo. In the photopolymerization process, visible or UV light (photoinitiators) generates free radicals that initiate crosslinking and radical polymerization. Visible light curing is more advantageous than UV curing due to less mutagenic effects, cytotoxicity, and nonthermogenic. In a study, photopolymerizable chitosan hydrogels were prepared from type 1 collagen (Col) to create a favorable microenvironment to promote bone repair [17]. With increasing the compressive modulus of methacrylated glycol chitosan, the osteogenic differentiation of the gel increased. Improved cell attachment, proliferation, and spreading were observed in mouse bone marrow stromal cells. This type of polymerization shows various advantages over other polymerization techniques. The inclusion of lactose and azide moieties in chitosan (photocrosslinkable derivative) widely applies in tissue engineering. Chitosan with vinyl benzoic acid derivatives also shows similar applications [25,26]. Generally, chitosan hydrogel employed in 3D printing usually suffers from poor mechanical and weak formability. To overcome this problem, a photopolymerization approach was adopted in the preparation of chitosan hydrogel from thiolterminated poly(ethylene glycol) (TPEG) and maleic chitosan with high acrylate group substitute [15]. This photopolymerization approach overcomes the oxygen inhibition effect. The molar ratio of thiol/acrylate is the critical parameter in regulating degradation, mechanical, microstructure, and rheological properties. Due to the photopolymerization, the gelling rate was increased by w2 folds, and the compressive strength was increased by w10 fold compared to a pure hydrogel. Photopolymerization is also possible between simple chitosan and synthetic PEG diacrylate [16]. In this synthesis, photosensitive modified vascular endothelial growth factor was added to chitosan, resulting in covalent crosslinking. This modified chitosan was found good biocompatible. The prepared 3D chitosan scaffold was successfully seeded with cells.

4.3.2 Physical crosslinking Physical crosslinking is another chitosan-based hydrogel network crosslinking (Table 4.2). The interaction can be an ioniccrosslinking hydrogel or polyelectrolyte complexes (Fig. 4.3). It is also possible for grafted or entangled chitosan hydrogels. The popularity of physical chitosan-based hydrogel is more than the chemical method due to the lower toxicity, ease of modulating the swelling rate, mechanical behaviors, and degradation patterns. However, these physical chitosan gels have demonstrated lower mechanical properties.

4.3.2.1 Ionic crosslinking Chitosan is a polycationic polymer having ionizable amine groups that form crosslinks with anions. Physical hydrogels’ sol/gel transition properties can be modulated with varying ionic strength, pH, temperature, etc. Hydrophobic and electrostatic interactions with hydrogen bonds help the gel structure’s integrity. Tripolyphosphate is such type of multivalent counterions that produce crosslinking gels and employed for delivery of low molecular weight drugs. Hydrogels are also prepared with dual or multiple crosslinking strategies. Using random copolymerization of monomers, polyampholyte hydrogels were synthesized. Hydrogen bonding and electrostatic interaction of these hydrogels helped swell the polymeric network [36]. These polyampholyte polymers contain randomly dispersed anionic and cationic groups that form

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Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine

TABLE 4.2 Chitosan physical hydrogel preparation and potential application. Type

Crosslinkers

Purpose

Application

Reference

Ionic crosslinking

Ammonium dibasic phosphate

Control bleeding in major trauma and surgeries

2% chitosane5% nanobioglass.

[27]

NaOH and gaseous NH3

Improve mechanical properties

Hydrogel accelerates resorption kinetics, macrophagic infiltration, and polarization of macrophage.

[28]

NaHCO3

Need to develop an injectable vehicle for drug delivery

The in situ solegel transition system showed a potential application in injectable drug delivery.

[29]

NaHCO3

Need to develop a bioactive injectable system as a cell carrier

Prepared hydrogel showed good dispersion, viability, and proliferation of encapsulated cells.

[30]

AgNO3, gaseous NH3

Chitosan physical hydrogel under a gaseous ammonia atmosphere showed low antibacterial activities and poor mechanical properties

Prepared chitosan hydrogel demonstrated higher tensile strength and superior antibacterial activity against negative and positive bacteria.

[31]

Sodium hyaluronate

Need to improve strong finite chain and sustain large deformation

Prepared hydrogel exhibited plastic-like behaviors with a monotonous decrease in stress.

[32]

Hyaluronic acid

Need to improve good lubrication, wettability, and biocompatibility

Prepared hydrogel successfully coated multilayered onto PET surface. The self-healing capacity is owed to the associatione dissociation ionic bonding process.

[33]

Poly(g-glutamic acid)

Need to develop porous 3D scaffolds

Chitosan/poly(g-glutamic acid) hydrogel significantly improved mechanical properties. This hydrogel demonstrated good adhesion and proliferation in mouse embryo fibroblasts.

[34]

k-carrageenan

Need to develop pH-responsive oral delivery of bovine serum albumin

The prepared PEC controlled the release of BSA in a pH-depended manner.

[35]

Polyelectrolyte complex

BSA, bovine serum albumin; NaHCO3, sodium bicarbonate; NH3, ammonia; PEC, polyelectrolyte complex.

viscoelastic and tough hydrogels with mechanical properties. Randomized ionic bonds help hydrogels strong and impart elasticity and permanent crosslinks, whereas weak bonds dissipate energy, reform, and reversibly break. Therefore, the supramolecular structure of physical hydrogel can be modulated with varying ionic combinations to manipulate the mechanical properties. This polyampholyte approach is a simple and best option for tough hydrogel. Usually, hydrogels lose their ideal properties in subzero environments. Chitosan acts as a hemostatic agent, widely used to aggregate blood cells to form a plug. To improve clotting property, nanobioglass with silica, calcium, and phosphate ions was incorporated during chitosan hydrogel preparation [27]. No CA was used in this hydrogel preparation. 2% chitosan solution (1% v/v acetic acid) solution’s pH was maintained at 6.5 using 1% NaOH under magnetic stirring to form a gel. Two steps gelation process was adopted to prepare chitosan physical hydrogels [28]. Initially, NaOH-induced gelation was followed by gaseous NH3-induced gelation. Chitosan hydrogels were obtained with ionic interaction between chitosan and a concentrated aqueous solution of NaOH. This chitosan solution was further exposed to gaseous NH3 to induce gelation. Adding NaHCO3 to the chitosan solution gives a homogeneous solution. The solegel transition of this chitosan/NaHCO3 system occurs from top to bottom after heating.

Hydrogels based on chitosan Chapter | 4

53

FIGURE 4.3 Different physical interactions in the chitosan hydrogel network [11].

The hydrogel showed porous morphology, which depended on the concentration of NaHCO3. The addition of the hydroxyapatite to chitosan indicated better protein as well as cell adhesion, higher osteogenic gene, and enhanced cell proliferation. A pH-responsive chitosan-hydroxyapatite hydrogel was prepared by NaHCO3 [30]. Under gaseous ammonia atmospheres, chitosan hydrogels demonstrated poor mechanical properties and low antibacterial activities. To overcome this problem, AgNO3 was introduced during the preparation of chitosan hydrogel under gaseous NH3 [31]. The swelling behavior and morphological characteristics of the hydrogel were improved significantly. This hydrogel was effective against both gram-negative and gram-positive bacteria.

4.3.2.2 Polyelectrolyte complex Chitosan acts as polyelectrolytes that depend on the solution pH. Polyelectrolyte complex networks are formed due to ionic interfaces between two opposite charged polymers. This interaction generally occurs at 4e6 pH. This process is fast, and modification can be done by rising crosslinker concentration and ionic strength. Polysaccharides are mainly employed due to their optimum biodegradability and biocompatibility. Chitosan is positively charged polysaccharides that can form complexes with negatively charged natural polymers like xanthan gum, carrageenan, pectin, alginate, etc. Chitosan can interact with synthetic polyphosphoric acid, PAA, polylactic acid, etc. The complex formation depends on the amount of each polymer, mixing ratio, and charge density of polymers. The net charge on the complexes is the key determining factor for their solubility. Complexes are usually insoluble or precipitate once the charge becomes zero [25,37]. Two phenomena are generally expected once chitosan is fully charged; pKa is significantly less, and all negatively charged macromolecules are incorporated in the solution. First, chitosan concentration (Cp) should be close or below to chain entanglement (C*inverse of intrinsic viscosity) for optimum liquideliquid separation. The number of macromolecules produced in the process determines the simple or complex coacervation in a precise crosslinker/chitosan molar ratio. Second, the system will make irregular macroaggregation if Cp > C*. In such conditions, inhomogeneous gelation and precipitation are irrelevant to the crosslinker amount [38]. Therefore, proper methodologies are required to synthesize homogenous ionic chitosan. The mechanically robust polyelectrolyte complex hydrogels were synthesized by tuning sodium hyaluronate (negative charge) and chitosan (positive charge) [32]. The prepared hydrogel showed a mechanically stable polyelectrolyte complex that becomes plastic-like strains and can form a flexible chain between synthetic and biological hydrogels. PET surface-bylayer-by-layer approach was adopted to form nanometric PEMs and macroscopic hydrogels [33]. In this complex, electrostatic interactions were established between the chitosan and hyaluronic acid in the development of intrinsic self-healing ability. Porous 3D constructs have a significant application in tissue engineering. A porous hydrogel based on a microstructure polyelectrolyte complex between chitosan and poly(g-glutamic acid) was prepared using a computer-aided wetspinning technique [34]. These porous constructs significantly influenced the mechanical, thermal, and swelling properties. Good adhesion and proliferation in mouse embryo fibroblasts were observed. The addition of k-carrageenan in the chitosan hydrogel formed pH-responsive delivery [35]. Chitosan-carrageenan polyelectrolyte complex loaded with BSA was

54

Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine

prepared by the incorporation method. The presence of salt interferes with the rheological properties of chitosan solution. The interaction between two polyelectrolytes is mostly delayed and designed complex demonstrated as different characteristics. This is a salt-induced impeding of the polyplex formation process.

4.3.3 Smart hydrogel Interest in smart or stimuli-responsive chitosan hydrogels is increasing in the biomedical field due to improved tissue engineering, aid in diagnostics, improved biosensing, and drug delivery (Fig. 4.4) (Table 4.3).

4.3.3.1 Temperature-sensitive hydrogels Low molecular weight heparin administration is not as frequent as heparin. Due to their interesting pharmacokinetic profile, these materials have long-term clinical applications like deep vein thrombosis. Thermoresponsive poloxamer-based systems containing LMWH/chitosan nanocomplexes were prepared to sustain LMWH release and diminish frequent subcutaneous application [39]. The mass ratio of LMWH/chitosan at 1:2 was optimized for preparing homogenized and stable nanocomplexes. This thermoresponsive hydrogel showed a gelation temperature of 28.6 C and a gelation time of 50 s. These parameters were decreased with the addition of hydroxypropylmethylcellulose. In both cases, prolonged dissolution and LMWH were achieved. Glycerophosphate (a phosphate supplement) is approved by the Medicines and Healthcare products Regulatory Agency (MHRA). After hydrolysis, glycerophosphate forms inorganic phosphate and glycerol within the body. Chitosan/bglycerophosphate thermosensitive gel was prepared by incorporating mitomycin-C for bladder cancer [40]. For bladder cancer treatment, local drug administration is needed to improve targeted delivery to urothelial malignant tissues. Chitosan formed gel in the presence of b-glycerophosphate at a body temperature of 37 C and was retained in the bladder mucosa for drug delivery to the urothelial cancerous tissues. Syringeability, gelation, drug release, and mucoadhesive properties are altered with changing chitosan molecular weight. High molecular weight chitosan (370 kDa) with b-glycerophosphate showed higher urine washout resistance than lower weight chitosan gel formulations. Chitosan/b-glycerophosphate hydrogels often showed a high degradation rate and poor mechanical properties in physiological conditions resulting in the limitation of their applications [53,54]. To overcome this problem, covalent crosslinking with double networks (DNs) is introduced. Genipin (a dialdehyde) spontaneously interacts with chitosan (specifically with primary amine groups) to form crosslinked networks and improve the mechanical properties. The swelling capacity of hydrogels can be decreased with a high degree of crosslinking, resulting in restriction of water diffusion. In such a situation, adding pullulan (a liner nonionic polysaccharide) to the chitosan solution can improve flexibility by forming semi-IPN networks [55].

FIGURE 4.4 Stimuli-responsive chitosan hydrogel and their potential application.

Hydrogels based on chitosan Chapter | 4

55

TABLE 4.3 Different types of stimuli-responsive chitosan hydrogel and their applications. Type of responsive

Drug/material used

The reason behind smart hydrogels

Name of formulation

Specification of gels

Reference

Temperaturesensitive

LMWH

Prolong release of LMWH and reduce the frequency of administration

The poloxamer-based system contained LMWH/chitosan nanocomplexes

1. Prolongation of drug release 2. Prolong gel dissolution 3. Addition of HPMC decrease gelation temperature and time

[39]

Mitomycin-C

To overcome the limitation of intravesical administration like a rapid washout and drug dilution

Chitosan/b-glycerophosphate in situ gel formulation

1. Improving the gelling properties 2. Prolong retention in the bladder

[40]

Disulfiram (DSF)

No adequate and controlled release of DSF for cancer therapy

Chitosan/DSF-loaded hydrogel

1. High cellular uptake compared to free DSF 2. Sustained delivery of DSF

[41]

Ciprofloxacin hydrochloride

Drug shows short halflife and low bioavailability

Chitosan/guar gum/PVP crosslinked hydrogel

1. Water swelling was decreased with increasing crosslinking 2. This hydrogel was stable at different temperatures

[42]

Ciprofloxacin

Drug shows short halflife and low bioavailability

N-trimethyl chitosan/sodium carboxymethyl xanthan gum hydrogel

1. Improving loading efficiency 2. Release patterns followed zero-order kinetics

[43]

Cephradine

The serum protein binding of this drug is less (8%e17%). More than 90% of administered drugs are excreted unchanged within a few hours

Chitosan/guar gum hydrogel blending with PEG

1. pH-dependent release kinetics observed 2. This hydrogel showed less swelling at basic and neutral fluid but showed higher swelling at acidic pH

[44]

Doxorubicin (DOX)

Systemic administration of DOX showed nonselective cardiac toxicity

Chitosan gel

1. Deeper skin diffusion 2. Increased DOX cytotoxicity

[45]

Lidocaine hydrochloride

This drug was taken as a model drug

Chitosan/ionic liquidbased hydrogels

1. Faster penetration 2. Higher release rates of lidocaine

[46]

Heparin and polylysine

To overcome the individual limitation of PEG and chitosan polymer in electro-responsive drug delivery

PEG-chitosan hybrid matrix hydrogel

1. Electrostatic interaction forces helped to release of anionic drug 2. Applying electric current, cationic drugs release was slowed down.

[47]

pH-sensitive

Electroresponsive

Continued

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Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine

TABLE 4.3 Different types of stimuli-responsive chitosan hydrogel and their applications.dcont’d Type of responsive

Drug/material used

The reason behind smart hydrogels

Name of formulation

Specification of gels

Reference

Magnetoresponsive

FITC-dextran, rhodamine B

Need to improve viscoelastic properties of gel and control of drug release

Chitosan hydrogels were prepared with catechol/Fe3þ ligands

1. Modulation of mechanical response was achieved by pHdependent metal/ catechol complex 2. Reduced swelling degree

[48]

Glucose oxidase

Glucose oxidase was used as a model oxidizing agent to understand the suitability of cellulose-based enzyme immobilization

Magnetic cellulosechitosan hybrid gel

1. Enzyme immobilization is possible with the glutaraldehyde technique 2. Prepared immobilized enzyme showed higher thermostable and improved storage stability

[13]

Adriamycin and rifampicin

To reduce the aggregation of magnetic nanoparticles

Magnetic chitosan hydrogels

1. Improved controlled release of drug 2. Both hydrophobic and hydrophilic drugs can be delivered by a remotely stimulated magnetic field

[49]

Bovine serum albumin

Limitation of proteinloading capacity

Chitosan-diblock copolymer of ethylene oxide and propylene oxide (CS/PEO-PPO

1. Entrapment efficiency of protein increased up to 80% 2. Continuous release was achieved for up to one week

[50]

10-hydroxy camptothecin (HCPT)

Reduction of rapid systemic elimination, systemic toxicity, and low water solubility

Chitosan nanogels

1. HCPT was entrapped into the core through facile diffusion 2. Enhanced biocompatibility was observed

[51]

Myricetin

This agent shows poor oral bioavailability. The previously developed formulation showed burst release resulting in a less therapeutic effect

Nanogels based on chitosan myricetin/HP-bCD inclusion complex

1. Myr-loaded nanogels showed sol to gel transition at physiological temperature 2. Formulation swelled in basic buffer and was eroded in acidic buffer solution

[52]

Nanogels

FITC, fluorescein isothiocyanate; HPMC, hydroxypropylmethylcellulose; LMWH, low molecular weight heparins; PEG, polyethylene glycol.

Due to its easy handling and unique properties, chitosan polymer is used to prepare hydrogel. Traditionally, disulfiram (DSF) is used for the treatment of alcoholism. Additionally, its anticancer activities have been reported earlier. For sustained effect, DSF-loaded chitosan hydrogels are prepared utilizing the self-assembly method [41]. Higher cytotoxicity was

Hydrogels based on chitosan Chapter | 4

57

observed dose-dependent compared to the free DSF solution. Additionally, these hydrogels showed better cellular uptake than the free DSF. The presence of other polymers with chitosan improves gelling characteristics. Poloxamers (triblock copolymers of propylene oxide and polyethylene oxide) are marketed with the trading name Pluronic. This is a gelation component used in the synthesis of thermosensitive chitosan hydrogels. A mixture of chitosan hydrochloride and pluronic F-127 was investigated to obtain thermosensitive hydrogels [56]. The optimum amino polysaccharide concentration was at 3%, and the chitosan:pluronic F-127 ratio of 30:70 showed a thermosensitive hydrogel. This formulation was viscous at 4 C but underwent a reversible thermal transition to a stable hydrogel at 37 C. The addition of D-ascorbic acid to hydrogel makes a different ordering in the structure of supramolecules and improves mucoadhesiveness.

4.3.3.2 pH-sensitive hydrogels pH is an essential stimulus for targeted therapeutic release, and it varies at different subcellular sites, pathological, physiology of gastrointestinal tract, or cancer site. The amine groups of chitosan help respond to chitosan against different pH. These pH-dependent properties of chitosan are helping in designing pH-stimulieresponsive hydrogels. For example, chitosan is insoluble at higher pH (oral cavity) but gets dissolves when exposed to the acidic pH of GIT. Due to this characteristic, chitosan is used in many gastric-related diseases like gastric carcinoma, ulcer, and gastritis [57]. Pure polysaccharide hydrogels show deficient mechanical strength and dissociate rapidly in body fluid resulting in the burst release of drugs. Synthetic polymers (like polyvinyl pyrrolidone, polyethylene glycol, PVA, etc.) improve the mechanical strength when added to the chitosan solution. In a study, Na-tripolyphosphate crosslinked pH-sensitive ternary composite hydrogel was prepared using natural and synthetic polymers like PVP, guar gum, and chitosan [42]. Using the combined properties, the prepared hydrogels had the ability to deliver drugs in a controlled manner. Ciprofloxacin (CFX) hydrochloride was incorporated into this hydrogel to give pH-responsive drug release at different pH concurrently to improve the bioavailability. Another pH-responsive hydrogel of this drug was prepared with N-trimethyl chitosan and sodium carboxymethyl xanthan gum [43]. The drug loading efficacy of this hydrogel was improved significantly, and the drug release from hydrogel was followed by zero-order kinetics. CFX is degraded by lysosome in conventional dosage form before killing bacteria, resulting in low intracellular drug concentration. In contrast, CFX-incorporated hydrogel showed high bactericidal potency against gram-negative and gram-positive bacteria. Chitosan-b-glycerophosphate hydrogels have been explored numerous times by altering different CAs like epichlorohydrin and GTA formaldehyde to improve drug delivery. Primarily, tetraethoxysilane (TEOS) is a silane crosslinker used in crosslinking in polymer amorphous zones. Thus, TEOS-based chitosan/guar gum eco-friendly biodegradable hydrogels synthesized by mixing with PEG to achieve pH-responsive, targeted delivery of cephradine [44]. Initially, crosslinked hydrogels showed more swelling that decreased afterward with increasing crosslinker amount. Prepared hydrogels showed maximum swelling in the acidic environment. At the same time, the swelling was slow at basic and neutral fluids.

4.3.3.3 Electro-responsive hydrogels Electro-responsive hydrogel is a smart hydrogel that transforms electric energy into mechanical energy. These hydrogels have several remarkable uses in the biomedical field, drug delivery, energy transduction, and sound dampening [57]. Iontophoresis is a tunable drug delivery to modulate drug transportation by varying variables like intensity, exposure to the electric stimulus, etc. Iontophoresis is a noninvasive drug delivery system that applies low current densities to improve drug transportation through biological barriers. The critical two mechanisms are electroosmosis and electromigration. Earlier, chitosan-based gels were employed for transdermal drug delivery systems. The properties of chitosan gels as a matrix were analyzed for hydration [58]. It was observed that the microviscosity of unhydrated gels increased with the acetic anhydride concentration, whereas it decreased in hydrated gels. Three different charged drugs were incorporated into this chitosan hydrogel. The drug release rate from the matrix was high in benzoic acid (anionic), followed by hydrocortisone (neutral) and lidocaine (cationic). Upon electric stimulus, the rate of gel weight loss was observed. Although pHsensitive, chitosan shows low mechanical flexibility in biological fluids. To resolve this problem, modulating chitosan design through chemical crosslinked blends, grafting copolymerization, and interpenetrating polymer networks. Ionic liquids (considered organic salt that is liquid below 100 C) are also introduced to functionalize chitosan. Multiresponsive semi-IPN chitosan hydrogels were prepared to improve the higher charge density and mechanical stability at a broader pH range [46]. Lidocaine was incorporated into this system as a model drug. This semi-IPN demonstrated faster penetration and a higher release rate of lidocaine hydrochloride under the external electrical stimulus. This study showed that this stimuli-responsive system has broad applications, including hemostatic, iontophoretic patches, and wound-healing dressings.

58

Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine

Percutaneous absorption and retention of doxorubicin (an anticancer drug) in the skin were investigated in the following iontophoresis using two different vehicles: cationic chitosan gel and nonionic hydroxyethylcellulose gel [45]. Electroosmosis was reduced significantly when iontophoresis was done with chitosan gel. Once the positively charged drug was incorporated into a cationic gel, transportation of the electroosmotic marker disappeared. This observation suggested that chitosan specifically interacts with negative charges of the skin. Therefore, this chitosan gel showed several advantages like reducing electroosmotic flow, the release of drug-using ionic interaction, and deeper skin diffusion. Multiple amine groups in the chitosan matrix cause extra protons in the aqueous solution resulting in high positive charges that attract negatively charged therapeutic molecules, including genes, enzymes, antibodies, and growth factors. Poly(ethylene glycol)-chitosan hydrogel is a hybrid matrix prepared by incorporating the anionic drug heparin [47]. The release of heparin was dependent on the electrostatic interaction forces. The competing negative charges of flowing electrons lead to a faster release of heparin from the positively charged network chains. The release of positively charged polylysine showed opposite release kinetics. Electric current application induced a sequence-free polylysine release slowly.

4.3.3.4 Magneto-responsive hydrogels The field of hydrogels is expanding, aiming for multifunctional, stimuli-responsive, mechanical, and tailorable physicochemical properties. The introduction of metal bonding into the hydrogels is a versatile strategy using the transition metal/ catechol moiety complexation. Magnetic hydrogels or ferrogel have wide biomedical applications like magnetic resonance imaging, tissue engineering, hyperthermia, and drug release. They often serve several additional advantages compared to other stimulation-sensitive drug release systems like easy localization of physical cues, noninvasive heat generation, remotely controlled movement, and rapid response [49]. Many magnetic-sensitive drug carriers are nonbiodegradable that often cause local inflammation. Magnetic nanoparticles (MNPs) are incorporated with chitosan or alginate due to superior compatibility. To prevent the aggregation of MNPs, magnetic chitosan hydrogels were prepared to achieve improved mechanical properties and magnetic response. Both hydrophilic and hydrophobic drugs were incorporated to check the release kinetics from the magnetic chitosan hydrogels. In both cases, stimuli drug delivery system was obtained. Mostly, the ligands are helpful to modulate and stabilize viscoelastic properties like mono-, bi-, or tridental complex with pH-dependent triggers and tuning the properties by modifying transition metal ions. Bioinspired magnetic and pHresponsive chitosan hydrogels functionalized with catechol were prepared with tunable elastic properties [48]. Catechol moiety containing hydrocaffeic acid was employed for the chitosan functionalization resulting in a viscoelastic response of gels. For expanding the multifunctionality of gel, iron oxide (g-Fe3O4) was incorporated, resulting in the modulation of mechanical response and further control over drug release. Before gelation, both high and low MW drugs were incorporated into the physical and chemical gel. Physical gels formed a compact structure with the tris-catechol/Fe3þ complexes than chemical gels. Chitosan-grafted Fe3O4 core can be prepared by electrostatic interaction between magnetic particles and chitosan with a radical polymerization initiator named [11-(2-Bromo-2-methyl)-propionyloxyl] undecyl trichlorosilane [59]. This system was stabilized with the help of acrylic acid polymerizing, followed by crosslinking with GTA. Drugs were incorporated into the electrostatic complexes containing chitosan and negative charged magnetic gel. The introduction of ionic liquids opens several new scopes for fabricating new materials like cellulose film, beads, and fibers. That’s why cellulose is widely employed in enzyme immobilization. In most cases, enzyme immobilization is carried out through the covalent attachment of GTA. This is the most convenient method for the coupling of an enzyme. That’s why magnetic cellulose-chitosan gel microspheres were prepared for enzyme immobilization [13]. Synthesized microspheres were stabilized when Fe3O4 was coated with biocompatible chitosan. The existence of the amino group in chitosan served a vital role in stabilizing magnetic composites that undergo aggregated in the fact of dipoleedipole solid attractions between particles. Enzyme immobilization of glucose oxidase showed higher thermostability, optima pH in a broader range, and improved storage stability than free glucose oxidase.

4.3.3.5 Nanogels Nanogels are the type of hydrogels with a size of 100 nm. Nanosized networks are prepared of physical or chemical crosslinked polymers resulting in drug loading of self-assembly mechanisms involving different kinds of bonding, including hydrophobic, van der Waals, electrostatic interaction between drug molecules, and polymers. Nanogels have demonstrated several advantages due to increased surface area. The potential application of nanogels has been observed in biomedical, sensor, and drug delivery. Nanogels are easy to administer and can swell up to 30 times resulting in sustaining drug release [52]. Hydrophilic nanoparticulate carriers have successfully delivered a wide range of drug molecules. In this fabrication process, organic solvents are used, which is a limitation for protein drug loading. To address this issue,

Hydrogels based on chitosan Chapter | 4

59

polysaccharide chitosan with a diblock copolymerized (propylene oxide and ethylene oxide incorporated) nanogels were prepared with size (200e1000 nm) and zeta potential of þ20 to 60þmV [50]. Significant improvement in loading efficiency and continuous release of loaded BSA was observed for one week. Chitosan-based nanogels are also designed to retard the growth of breast cancer cell lines [51]. The antitumor drug, 10-hydroxy camptothecin (HCPT), was entrapped into the chitosan-based nanogels through the facile diffusion. Myricetin (Myr) (a natural antioxidant) was incorporated into the chitosan nanogels to improve oral bioavailability [52]. The lower oral bioavailability of this material is mainly due to the poor solubility, which was boosted by forming a Myr/hydroxypropyl-beta-cyclodextrin (HP-b-CD) inclusion complex. This inclusion complex was added into B-glycerol phosphate/chitosan solution to create nanogels. These nanogels were able to transit sol to gel at physiological temperature. These nanogels swelled in the primary buffer and easily eroded in acidic buffer solutions.

4.4 Application of chitosan-based hydrogel Being a biodegradable and biocompatible polymer, chitosan has a broader application in the biomedical field. Additionally, its degradable products are nonimmunogenic and nontoxic. This material has numerous antioxidant, hemostatic, and chelating agent applications. It can control the bleeding as it acts as a procoagulant resulting in the initiation of clotting. Chitosan application is increasing gradually in multiples areas such as food industries, agriculture, cosmetics, medical, and pharmaceutics.

4.4.1 Drug delivery Chitosan-based hydrogels and their chemically altered forms are extensively explored in drug delivery applications (Table 4.4). Chitosan can adhere to mucosal glycoproteins (negative charge). Due to ionic interaction, chitosan is designated as a bioadhesive material. Bioadhesive materials increase drug residency time and provide localized drug delivery. The method of chitosan hydrogels is convenient and can be industrial scalable. Thus, different structure modifications are introduced to explore different routes of drug delivery.

4.4.1.1 Oral delivery Chitosan hydrogel-based scaffolds can be used successfully in different delivery systems like colon, intestine, stomach, and oral cavity. Mucoadhesive characteristic of chitosan is utilized for oral administration of this drug to increase absorption. After colon delivery, chitosan hydrogel can be used in inflammatory bowel or irritable diseases. In such cases, pH-sensitive chitosan hydrogels can give additional advantages of sustained release for prong periods. Methoxy poly(ethylene glycol)grafted carboxymethyl chitosan and alginate composed of hydrogel beads were prepared with an interpenetrating polymeric matrix [60]. This formulation showed slow release of BSA at pH 1.2, whereas rapid release was observed at pH 7.4. This pH-responsive hydrogel can be a promising formulation for site-specific drug delivery of protein in the intestine. pHsensitive chitosan hydrogel is also prepared with polyacrylamide using ammonium persulfate as an initiator and methylenebisacrylamide as a crosslinker for oral delivery of insulin [61]. This hydrogel successfully reduced glucose levels in diabetic mice blood with an oral bioavailability of 4.43%. Often polymer was added to chitosan to improve the polymer matrix for controlling the release kinetics of the drug. For controlled drug delivery of acyclovir, chitosan hydrogel was prepared by incorporating xanthan gum, 2-acrylamido-2methylpropane sulfonic acid with potassium persulfate (initiator), and N0 N’-methylene bisacrylamide (crosslinkers) [63]. This hydrogel synthesized by free radical polymerization technique demonstrated very less release of acyclovir at pH 1.2, but the release kinetics improved significantly with raising the media pH. The release kinetics of the drug followed the KorsmeyerePeppas model. This drug was incorporated into b-cyclodextrin-chitosan hydrogel with the same objective [64]. Here, the release kinetics followed zero-order kinetics. Additionally, pH-dependent drug release patterns and swelling behavior were observed. Stomach-specific antibiotic delivery is highly beneficial in treating Helicobacter pylori in peptic ulcers. The stability of the drug in the stomach is the leading cause of effective therapy failure. To overcome this issue, a pH-sensitive chitosanpoly(ethylene oxide)-based semi-IPN polymer network was prepared for the antibiotic delivery to the stomach [62]. The porous matrix of freeze-dried hydrogels can swell faster due to capillary action’s rapid influx of simulated gastric fluid. This semi-IPN could be a good option for drug delivery to the stomach, specifically in treating H. pylori infection.

60

TABLE 4.4 Application of chitosan-based hydrogel in drug delivery with their method of preparation. Method of preparation

Material employed

Drug used

Application

Reference

mPEG-g-CMC

Physical mixing

Methoxy poly(ethylene glycol) grafted carboxymethyl chitosan

Bovine serum albumin

Improving the release of drugs in the colon with pH-responsive hydrogel.

[60]

PAA/S-chitosan hydrogel

Chemical method using a crosslinker

Polyacrylamide, S-chitosan, ammonium persulfate, methylene bisacrylamide

Insulin

Prolonged attachment to the intestinal tissue. Extended release to reduce glucose level.

[61]

Freeze-dried Chitosan-PEO semi-IPN porous matrix

Chemical method

Chitosan, poly(ethylene oxide)

Amoxicillin, metronidazole

The entrapment efficiency of antibiotics was improved and can be delivered drug in gastric fluid.

[62]

Chitosan/xanthan gum hydrogel

Chemical polymerization method

Chitosan, xanthan gum, a-acrylamido-2methylpropane sulfonic acid, Nʹ Nʹ-methylene bisacrylamide, potassium persulfate

Acyclovir

Altering the composition of chitosan and xanthan gum drug loading and release patterns can be changed.

[63]

b-cyclodextrin chitosanbased hydrogel

Chemical polymerization method

Chitosan, b-cyclodextrin, Nʹ Nʹ-methylene bis-acrylamide, methacrylic acid

Acyclovir

Prepared hydrogel released acyclovir in a controlled manner.

[64]

C/G/GP hydrogel

Physical method

Chitosan, gelatin, glycerol 2-phosphate sodium

Latanoprost

Prepared gel decreases ocular hypertension after injection.

[65]

Chitosan/DGP thermosensitive hydrogel

Physical method

Chitosan, chicken egg albumin, a-DGlucose 1-phosphate

Levocetirizine

Prevention of allergic conjunctivitis.

[66]

In situ thermosensitive chitosan-based nasal formulation

Physical method

Chitosan, b-glycerophosphate disodium salt

Ibuprofen

Improved the solubility of ibuprofen. Accelerated ibuprofen penetrated the human nasal epithelial cells.

[67]

HTT-PEG-GP hydrogel

Chemical method

N-{(2-hydroxy-3-trimethylammonium) propyl] chitosan chloride, a-b-glycerophosphate, poly(ethylene glycol)

Insulin

This hydrogel formulation decreased the blood glucose level for 4e5 h after administration. Additionally, this hydrogel improved the absorption of hydrophilic macromolecular drug.

[68]

Chitosan/GP thermosensitive hydrogel

Physical method

Chitosan, glycerophosphate disodium

Venlafaxine hydrochloride (VH)

This thermoresponsive gel presented a sustained VH delivery.

[69]

Thermo-gelling injectable nanogels

Physical method

Carboxymethyl hexanoyl chitosan

Ethosuximide

Injectable nanogels for a depot drug delivery system.

[70]

In situ hydrogel of chitosan

Physical method

Chitosan, GP

Pingyangmycin

These hydrogels showed dual activities like interruption of nutritional supply to VM and sustained release of PYM.

[71]

Liposome-incorporated chitosan hydrogel

Physical

Chitosan, GP

Topotecan hydrochloride

Improving antitumor efficacy of TPT by reducing the release of drug.

[72]

GP, glycerophosphate disodium.

Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine

Synthesized material

Hydrogels based on chitosan Chapter | 4

61

4.4.1.2 Ocular delivery The short persistence of formulation on the eye is the main limitation of conventional ocular formulations. This issue can be resolved by incorporating gel. Concurrently, it improves drug bioavailability and drug retention in the affected site. Latanoprost is a potent ocular hypotensive compound generally recommended as first-line therapy in glaucoma. Daily installation and undesired side effects cause the treatment failure, recurring glaucoma, and prolonged persistence therapy. To overcome this problem, sustained release, the thermosensitive hydrogel was developed based on chitosan-gelatin for controlling ocular hypertension [65]. After injection, this hydrogel released the drug sustainably and decreased ocular hypertension in the glaucoma therapy. Allergic conjunctivitis is inflammation of the conjunctiva caused by an allergic reaction. To improve drug retention and patient comfort, the levocetirizine-loaded chitosan-based thermosensitive hydrogel was developed for ocular delivery [66]. Prepared hydrogel rapidly releases the drug initially, followed by an extended release in a sustained manner. The ocular retention time of the hydrogel was improved considerably compared to the eye drop.

4.4.1.3 Nasal delivery Chitosan can improve drug transportation after opening the tight junctions between the epithelial cells of the mucosal membrane. Chitosan promotes nasal drug delivery due to mucoadhesive and high water absorption capacity. The conventional drug administration route fails to deliver the drug to the brain due to restrictions in crossing the bloodebrain barrier (BBB). Recently, interest in nose-to-brain delivery has been increasing gradually for drugs targeting the lungs after bypassing the BBB [73]. This type of delivery is helpful for all kinds of brain-related diseases. Thermoresponsive nasal in situ gel formulations can be sprayed at room temperature. At physiological temperature, the solution turns into a semisolid state and adheres with nasal mucosa for the mucoadhesive property of chitosan. These in situ gels are successfully employed to deliver therapeutic molecules to the brain. For brain delivery of ibuprofen, thermosensitive chitosan hydrogel was prepared for nasal delivery [67]. With decreasing the MW of chitosan, spray patterns of hydrogel were improved. This in situ formulation enhanced ibuprofen’s solubility and accelerated its transport through human nasal epithelial cells. Thus, the chitosan-based thermosensitive hydrogel could be advantageous for intranasal administration as the spray solution is liquid at room temperature. Afterward, this sprayed solution undergoes a gelation process in the nasal cavity. Quaternized chitosan is successfully employed as an absorption enhancer and can open the tight junction between epithelial cells. Thermosensitive hydrogel for nasal delivery of insulin was prepared with chitosan chloride [68]. This formulation can spray easily to the nasal cavity and spread on the nasal mucosa. Afterward, this solution underwent a thermal transition to nonflowing hydrogel at body temperature with a reduced rate of nasal mucociliary clearance.

4.4.1.4 Parenteral delivery Chitosan is generally dissolved in acidic environments, and a phase separation observes at higher pH (above 6.5). This acidic solution is neutralized by adding polyol phosphates without phase separation at physiological pH. This combined solution remains in a liquid form at low temperatures. However, it undergoes solegel transition at body temperature. This observation opens new scope for developing a chitosan-based parenteral delivery system. In situ implants have drawn significant interest over conventional implants and microparticles. After in situ implantation, a low viscous solution converts into a solid depot or gel at the injection site, avoiding the microsurgery. For example, in situ gelling systems containing b-GP and chitosan are commercially available to administer water-soluble small-molecular drugs, peptides, and proteins. For sustained delivery of venlafaxine hydrochloride (VH-water-soluble drugs), chitosan/GP thermosensitive hydrogels were prepared for better-sustained release than the solution [69]. This combination was used to deliver multiple drugs like curcumin, melphalan, docetaxel, camptothecin, paclitaxel, and many more [74] for the inhibition of tumor growth. Another highly water-soluble drug, ethosuximide, was delivered by linking self-assembled nanocapsules of carboxymethyl hexanoyl chitosan with glycerol and glycerophosphate di-sodium salt [70]. With this encapsulation, the burst release and release rate were reduced significantly. For the treatment of vascular malformation, pingyangmycin (PYM)-loaded chitosan thermogels were prepared [71]. This hydrogel sustained the drug release for up to 12 days. In vitro study demonstrated the inhibition of proliferation and induction of apoptosis of EA.hy926 cells in a time and concentration-dependent fashion. After injection, this solution changed into a semisolid embolic agent and demonstrated dual kinetics like the sustained release of PYM and interruption of nutritional supply to VM. In the rabbit model, PYM-loaded thermogels localized drug concentration for a longer time and lowered the drug concentration in the systemic circulation.

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Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine

The release rate can be tuned by incorporating the drug into nanoparticles, microspheres, or liposomes. Extended release of drug can be possible by incorporating these nanoparticulate systems into the chitosan-based hydrogels. Topotecan hydrochloride (TPT) is a potent ovarian anticancer drug. First, TPT was incorporated into liposomes and chitosanbased hydrogels [72]. The release rate was slowed down. This hydrogel demonstrated improved antitumor efficacy of TPT.

4.4.2 Tissue engineering The tissue engineering field attends significant interest by combining different disciplines like nanotechnology, functional scaffold materials, stem cell biology, etc. Tissue engineering integrates biology with engineering to fabricate cellular components or tissues in vitro in order to upgrade knowledge of managing tissue repairing within the body [75]. In this field, bioartificial tissues are constructed in vitro as well as in vivo by altering cell functions and growth. The primary mechanism is isolating suitable cells from donor tissues and biocompatible scaffold materials and their implantation at a particular site. Biomaterials for tissue engineering are the key elements and should possess optimum surface chemistry, biodegradability, and porosity. Compared with other natural polymers, chitosan has been found advantageous in terms of its distinctive cationic amino group-containing polysaccharide structures that resemble macromolecular structures of the human body. Its biocompatibility, biodegradability, bioresorbability, low immunogenicity, and economical made chitosan a potential polymer in the tissue engineering realm [76]. Hence, it has drawn massive attention in tissue engineering (Table 4.5). The most prominent application in tissue engineering is the fabrication of cartilage, periodontal, intervertebral disc, bone tissue, skin, corneal regeneration, blood vessel tissue engineering, etc. A combination of synthetic and natural materials with chitosan has been explored in designing scaffolds in the recent era. Biopolymers like hyaluronic acid, poly(lactic acid), polycaprolactone, and alginate gelatin are added to chitosan to improve the structural integrity and mechanical strength of chitosan-based biomaterials [81]. Synthetic polymers exhibit better functionality than natural polymers and are often combined with natural polymers. However, some synthetic polymers show immune response, toxicity, and poor biocompatibility resulting in fibrous encapsulation and inflammatory reactions. Due to increasing demand, natural biodegradable polymers are extensively explored in tissue engineering. Out of them, chitosan is a unique one due to its several exciting properties, as mentioned earlier. There are several other properties of chitosan that make it more interesting, like delivery of materials/drug, high drug loading capacity, porous structure, and ability to form electrostatic interactions. Osteomyelitis is an advanced inflammatory disease that requires long-term systemic therapy of antibiotics. Rather than exposing the whole body, the locally applied antibiotics at the surgical site can prevent postoperative osteomyelitis. Chitosan-based in situ calcium phosphate composites were prepared by loading moxifloxacin hydrochloride [77]. This drug was released entirely in three days. This formulation prompted proliferation and differentiation of osteoblasts with

TABLE 4.5 Application of chitosan-based hydrogel in tissue engineering. Synthesized material

Method of preparation

Chitosan in situ hydrogel

Material employed

Drug used

Application

Reference

Chemical method

Chitosan, di-ammonium hydrogen phosphate, calcium nitrate, glutaraldehyde, ammonia solution

Moxifloxacin hydrochloride

Chitosan-calcium phosphate scaffolds boost cellular differentiation/proliferation

[77]

Chitosan in situ hydrogel

Chemical method

Chitosan, biodegradable glass nanoparticles, tetraethyl orthosilicate, triethyl phosphate, b-glycerophosphate disodium, gelatin

d

Formation of new tissues in defects treatment

[78]

Chitosan/kcarrageenan hydrogel

Chemical polymerization

Chitosan, gold nanoparticles, k-carrageenan, poly(Nisopropyl acrylamide)

d

Metal-incorporated hydrogel demonstrated enhanced cell growth and attachment

[79]

Chitosan hydrogel

Chemical method

Chitosan, gelatin, octamethacrylated polyhedral oligomeric silsesquioxane

Mesenchymal stem cells

Accelerating bone regeneration in rat calvarias defects

[80]

Hydrogels based on chitosan Chapter | 4

63

reduced inflammation, intramedullary fibrosis, and bacterial count in the bone tissue specimens in osteomyelitis. Nanocomposite hydrogels are also explored in a bone generation. Bioactive glass nanocomposite hydrogel was prepared with chitosan-gelatin that showed good biocompatibility in the rat’s model [82]. This hydrogel demonstrated regenerative potency for generating new tissue to treat defects. This thermoresponsive gel bound with the living tissues and possessed osteoblasts proliferation and angiogenesis-related gene expression resulting in bone regeneration and cell proliferation activities [78]. The introduction of metallic nanoparticles into the injectable hydrogels can facilitate bone generation applications. Gold nanoparticles, k-carrageenan, and poly(N-isopropylacrylamide) were mixed into chitosan hydrogel [79]. This metal incorporated injectable hydrogel enhanced cell growth and demonstrated its potential application in tissue engineering. DN hydrogel combines fragile polyelectrolyte integrated with a flexible neutral polymer. This DN hydrogel demonstrates a superior mechanical performance due to its asymmetric structure where the rigid network effectively disintegrates energy. However, the soft network preserves rigidity during the distortion process. That’s why biodegradable hybrid DN hydrogel was prepared to promote the bone generation [80]. This DN hydrogel had the advantages of nanocomposite and gel hydrogel, which exhibited hierarchically porous structure, controllable biodegradability, and mechanical properties. Additionally, this DN improved mesenchymal stem cell proliferation and osteogenic differentiation.

4.4.3 Commercially available chitosan-based hydrogel Several chitosan-based hydrogels are available in the market or are undergoing clinical trials. Due to its nontoxic, biocompatible, and biodegradable characteristics, chitosan is used in the delivery system and scaffolds for topical administration. The United States Food and Drug Administration already approves many chitosan products for specific applications, including wound dressing and sponges (Table 4.6). Many hemostatic bandages or gels (HemCon) and sprays (ChitoClear) are used for curing skin wounds. Gel and spray based on chitosan are also available for delivery to nasal mucosal membranes, for example, ChitoRhino [87]. The potential use of these hydrogels is in wound healing. Free and unimpeded bleeding from wounds may be contaminated with microorganisms entering the wound from the environment. HemCon is a chitosan acetate bandage preparation used to control hemostasis. It showed high wound adhesiveness than alginate [83]. The overall outcome of wound dressing is the prevention of colonization and proliferation of microorganisms. A novel next-generation wound dressing was developed with 100% chitosan fibers by Foshan United Medical Technologies Ltd. to efficiently manage wound exudate and other variables impeding wounds [84]. This chemically modified chitosan dressing demonstrated enhanced absorbency and gelling for the specific treatment of exudating chronic wounds. During the wound-healing process, morphofunctional normality can result in the formation of disoriented connective tissues that lead to reduced mechanical strength. Biomaterials can help in the proper physiological reconstruction of the skin. Chitosan membrane-based wound products exemplify such a solution [85]. Trauma and accidents are leading causes of disability and death worldwide. The deficiency of sufficient prehospital treatment and untreated bleeding are the main reasons. Chitosan-based hemostatic dressing Axiostat was developed to control hemorrhage in an ambulance setting [86]. Chitosan-based thermogelling hydrogel is prepared with b-glycerophosphate (b-GP) (a gelling agent). The chitosan/ b-GP systems have a wide application in various pharmaceutical and biomedical applications. Glycerophosphate is a biocompatible material approved by the US FDA. The in situ gelling system consisting of these materials is available in the market as implants that can deliver hydrophilic drugs, peptides, and proteins [74].

TABLE 4.6 Marketed/clinical trials of chitosan-based hydrogels. Hydrogel

Problem statement

Application

Reference

HemCon

Free and unimpeded bleeding from wounds may be contaminated with the microorganisms.

This hydrogel has demonstrated optimum hemostasis and antibacterial activities against some strains of bacteria.

[83]

KA01 chitosan wound dressing

Prevention of colonization and proliferation of microorganisms in the wound bed.

This hydrogel improved wound healing by enabling reepithelialization with reduced pain.

[84]

Chitosan mesh membrane

During the healing process, abnormal tissue architecture leads to scar formation and reduces mechanical strength.

This hydrogel-based membrane showed good adherence, hemostasis with reduced pain and itching.

[85]

Continued

64

Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine

TABLE 4.6 Marketed/clinical trials of chitosan-based hydrogels.dcont’d Hydrogel

Problem statement

Application

Reference

Axiostat

Deficiency of sufficient prehospital care and uncontrolled bleeding from the wound.

It showed rapid hemostasis, which was suitable for emergency accidents and trauma. Removal of a wound is easy without any stain.

[86]

ChitoRhino

Need to develop materials for fighting against bacterial resistance.

This gel showed adequate hemostasis characteristics and efficiently healed a wound after surgery.

[87]

ChitoHeal

Need to develop a powerful wound-healing product that forms secondary skin.

This product is helpful for rapid healing, scar reduction, and effective against diabetic foot ulcers, scratches, cuts, and burns.

[88]

KytoCel

Required product for managing moderate to excessive chronic and acute wounds.

This chitosan fiber is used as an absorbent dressing for managing serious wounds.

[89]

PosiSep

Need to minimize bleeding and edema.

This chitosan fiber is used for nasal dressing and demonstrates rapid development after hydration.

[90]

ExcelArrest XT

Need to accelerate the clotting process.

This chitosan-based patch promotes clotting to control bleeding.

[91]

ChitoClot Pad

Need to reduce the risk of complications for percutaneous coronary intervention.

This hydrogel absorbed blood and undergoes gelation.

[92]

XSTAT

Need immediate control of hemorrhage from penetrating wounds.

This hydrogel is used in gunshot wounds.

[93]

Chitoderm plus

Need to develop a wound dressing that shows a strong nonadhesive superabsorber.

This hydrogel shows good absorbent properties.

[94]

ChitoClear

Need to develop a material that can attract negatively charged RBC.

This hydrogel is used as a hemostatic agent by fascinating with negatively charged RBC.

[95]

Celox

Need for a dressing that clots hypothermic blood.

This hydrogel is used for rapid hemostatic properties and reduces the loss of blood.

[94]

4.5 Conclusion and prospects Chitosan is a widely used positive-charged and low-cost natural polymer with promising applications in drug delivery and tissue engineering. With modulating physical properties, chitosan shows good biodegradability and excellent biocompatibility. The existence of hydroxyl and amine groups is the modification site for tissue adherence, drug delivery, chemical modification, etc. Owing to unique properties, chitosan has certain restrictions for being used for controlled drug delivery and tissue engineering. These barriers can be overcome by suitable alteration in the chitosan backbone. The importance of modified chitosan hydrogel is increasing significantly in drug delivery and tissue engineering. Crosslinking hydrogels with synthetic polymers have demonstrated superior mechanical properties, including tensile stresses, Young’s modulus, compressive stresses, flexibility, and elasticity. Hydrogels can be stored in dried form. The most frequent technique used for this drying is freeze-drying. This technique is successfully employed for various hydrogel materials, including synthetic and natural hydrophilic polymers. However, hydrogel architecture is affected during the quenching process due to liquid interfacial tension occurred during the evaporation of the solvent. Injectable hydrogels are generally prepared with thermoresponsive polymers having lower critical solution temperatures (LCST). The LCST polymers exhibit sol-to-gel transition with rising temperature due to alteration of hydrophobic/hydrophilic balance within the hydrogel systems. In the preparation of in situ injectable hydrogels, the LCST of the polymer should be less than the physiological temperature. Injectable chitosan-based hydrogels have been demonstrated successfully in regenerative medicine, tissue repair, differentiation/proliferation, spreading, and cell attachments. Combining other natural or synthetic polymers with chitosanbased hydrogels is required to improve the biological, physicochemical, and mechanical characteristics. This addition will increase the potential application of injectable hydrogels in biomedical fields. However, this chitosan-based hydrogel has clinical limitations like raw material resources, commercial production, and improper standardization of molecular

Hydrogels based on chitosan Chapter | 4

65

weight distribution. High molecular weight chitosan can cause inflammation in bone grafting. There are limited studies and insufficient clinical utilities of chitosan-based hydrogels as bone scaffolds. Therefore, future research should be focused on inventing advanced technology for potential clinical applications. For clinical prospects, high loading efficiency and controlled drug release should be the primary research objectives for chitosan-based hydrogels.

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Pharmaceut Dev Technol 2014;20:812e9. [73] Kiran P, Debnath SK, Neekhra S, et al. Designing nanoformulation for the nose-to-brain delivery in Parkinson’s disease: advancements and barrier. Wiley Interdiscip Rev Nanomed Nanobiotechnol 2022;14:e1768. [74] Liu L, Gao Q, Lu X, Zhou H. In situ forming hydrogels based on chitosan for drug delivery and tissue regeneration. Asian J Pharm Sci 2016;11:673e83. [75] Mantha S, Pillai S, Khayambashi P, et al. Smart hydrogels in tissue engineering and regenerative medicine. Materials 2019;12:3323. [76] Gheorghita R, Anchidin-Norocel L, Filip R, Dimian M, Covasa M. Applications of biopolymers for drugs and probiotics delivery. Polymers 2021;13:2729. [77] Radwan NH, Nasr M, Ishak RAH, Abdeltawab NF, Awad GAS. Chitosan-calcium phosphate composite scaffolds for control of postoperative osteomyelitis: fabrication, characterization, and in vitro-in vivo evaluation. Carbohydr Polym 2020;244:116482. [78] Moreira CDF, Carvalho SM, Sousa RG, Mansur HS, Pereira MM. Nanostructured chitosan/gelatin/bioactive glass in situ forming hydrogel composites as a potential injectable matrix for bone tissue engineering. Mater Chem Phys 2018;218:304e16. [79] Pourjavadi A, Doroudian M, Ahadpour A, Azari S. Injectable chitosan/k-carrageenan hydrogel designed with au nanoparticles: a conductive scaffold for tissue engineering demands. Int J Biol Macromol 2019;126:310e7. [80] Zhang Y, Chen M, Tian J, et al. In situ bone regeneration enabled by a biodegradable hybrid double-network hydrogel. Biomater Sci 2019;7:3266e76. [81] Islam MM, Shahruzzaman M, Biswas S, Nurus Sakib M, Rashid TU. Chitosan based bioactive materials in tissue engineering applications-A review. Bioact Mater 2020;5:164e83. [82] Moreira CDF, Carvalho SM, Florentino RM, et al. Injectable chitosan/gelatin/bioactive glass nanocomposite hydrogels for potential bone regeneration: in vitro and in vivo analyses. Int J Biol Macromol 2019;132:811e21. [83] Burkatovskaya M, Tegos GP, Swietlik E, Demidova TN, Castano PA, Hamblin MR. Use of chitosan bandage to prevent fatal infections developing from highly contaminated wounds in mice. Biomaterials 2006;27:4157e64. [84] Mo X, Cen J, Gibson E, Wang R, Percival SL. An open multicenter comparative randomized clinical study on chitosan. Wound Repair Regen 2015;23:518e24. [85] Azad AK, Sermsintham N, Chandrkrachang S, Stevens WF. Chitosan membrane as a wound-healing dressing: characterization and clinical application. J Biomed Mater Res Part B Appl Biomater 2004;69:216e22. [86] Kabeer M, Venugopalan PP, Subhash VC. Pre-hospital hemorrhagic control effectiveness of Axiostat dressing versus conventional method in acute hemorrhage due to trauma. Cureus 2019;11:e5527.  [87] Hemmingsen LM, Skalko-Basnet N, Jøraholmen MW. The expanded role of chitosan in localized antimicrobial therapy. Mar Drugs 2021;19. [88] Chitotech. ChitoHeal gel. 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[89] Wound-Care. Highly absorbent Kytocel dressings. Wound-Care; 2020. p. 1. https://www.wound-care.co.uk/dressings/kytocel.html. [Accessed 29 April 2022]. [90] Hemostasisllc. Hemostasis, LLC | PosiSep C and C2. HemostassisllicCom; 2020. https://www.hemostasisllc.com/products/posisepc-c2/. [Accessed 29 April 2022]. [91] Hemostasis. ExcelArrest XT topical hemostatic patch. GlobalHemostasisllcCom; 2020. p. 1. http://global.hemostasisllc.com/excelarrest.html. [Accessed 29 April 2022]. [92] Anscate. ChitoClot pad. AnscareCom; 2020. https://www.anscare.com/en/product/detail/ChitoClot_Pad. [Accessed 29 April 2022]. [93] Revmedx. XStat e RevMedx. RevmedxCom; 2020. https://www.revmedx.com/xstat/. [Accessed 29 April 2022]. [94] Matica MA, Aachmann FL, Tøndervik A, Sletta H, Ostafe V. Chitosan as a wound dressing starting material: antimicrobial properties and mode of action. Int J Mol Sci 2019;20:5889. [95] Primex. ChitoCLEAR | ChitoClear. PrimexIs; 2020. https://www.primex.is/products-services/chitoclear/. [Accessed 29 April 2022].

Chapter 5

Hydrogels based on heparin and its conjugates Hemant Ramachandra Badwaik1, Kalyani Sakure2 and Tapan Kumar Giri3 1

Shri Shankaracharya Institute of Pharmaceutical Sciences and Research, Bhilai, Chhattisgarh, India; 2Rungta College of Pharmaceutical Sciences

and Research, Bhilai, Chhattisgarh, India; 3Department of Pharmaceutical Technology, Jadavpur University, Kolkata, West Bengal, India

5.1 Introduction Heparin (HEP) is a linear sulfated polysaccharide of glycosaminoglycan origin, comprising 1,4-linked uronic acid and glucosamine (Fig. 5.1) [1]. HEP is often extracted from pig intestine or bovine lung and has an average molecular weight of around 15 kDa. Its chemical structure and molecular weight are extremely varied [2]. HEP has a significant negative charge owing to the existence of sulfate and carboxylate groups [3]. HEP is a very acidic chemical, which is somewhat neutralized in the body by swapping acidic hydrogen atoms in sulfate groups with sodium ions due to the existence of sulfonate and carboxyl groups in the molecules. HEP is thought to work by neutralizing a variety of coagulation factors, therefore interfering with the conversion of prothrombin to thrombin. HEP is used to avoid thrombosis, thrombosis, and embolism, as well as to maintain fluid conditions in the blood during artificial blood circulation and hemodialysis [4]. HEP integration into biomaterials has shown to be extremely desirable due to its beneficial biological effects. For prolonged anticoagulant release, HEP is frequently encapsulated or covalently coupled with hydrogel (HGL) [5e7]. HEPcontaining HGLs have also been employed extensively for the effective storage and release of growth factors to stimulate angiogenesis and bone repair [8,9]. Because of their biocompatibility, low toxicity, cheap cost, and benign gelation conditions, HEP-functionalized HGLs have been developed in recent years [10]. This chapter summarizes current advancements in the creation of HEP-based HGLs, with a focus on tissue engineering and drug delivery applications.

5.2 HEP-based HGLs HGL refers to a molecular hydrophilic polymer-based network that is crosslinked and produced in a three-dimensional space. It has the capacity to take in a considerable volume of water, which may range from 10 to 1000 times its original volume while remaining stable upon expanding [11]. Researchers developed various types of HGLs for specific and targeted deliveries [12]; detailed are shown in Fig. 5.2. HEP is used to make HEP-based HGLs that may be categorized according to the mechanism of HGL formation and release behavior. HEP-based HGLs are categorized according to their physical characteristics into conventional HGLs or smart HGLs. Similarly, they might be classified according to the nature of crosslinking into chemical crosslinked HGLs, physical crosslinked HGLs, and combined interaction (duel crosslinked) HGLs.

5.2.1 Chemically crosslinked HEP HGLs Covalent bonding is the most commonly used method for constructing HEP-based HGLs. In order to create a persistent HGL, chemical crosslinking involves covalent bonding between polymer strands. Chemically crosslinked HEP-based HGLs fabricated using various techniques.

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FIGURE 5.1 Structure of HEP.

FIGURE 5.2 Classification of HGLs.

5.2.1.1 Photo-crosslinked HEP-based HGLs HGLs may be created in situ using photo-induced polymerization. HGLs may be created in situ using photo-induced polymerization. The existence of photo-sensitive functional groups influences the formation of HGLs via photocrosslinking. When a photo-sensitive functional group is linked to a polymer, it becomes capable of forming crosslinks when exposed to light, such as ultraviolet (UV) light [13]. Benoit et al. 2007 synthesized photo-crosslinked HGL of methacrylated HEP with PEG dimethacrylate. The HEPfunctionalized PEG HGLs could sequester bFGF and release it in a controlled manner. In comparison to controls, the (fibroblast growth factor) bFGF-loaded HGLs stimulated the growth of human mesenchymal stem cells (hMSCs) [14]. A photo-induced radical polymerization has been used to incorporate methacrylated HEP into a wide range of polymeric systems such as poly-(vinyl alcohol) (PVA), alginate, and hyaluronan [15e19]. To enhance angiogenesis, Oliviero et al. fabricated porous photo-triggered HGLs as a 3D matrix for sustained vascular endothelial growth factor (VEGF) release [20]. As HGL precursors, the HEP was conjugated with N-(3-amino propyl) methacrylamide hydrochloride (APMMA) (Fig. 5.3) and PEG is modified to poly(ethylene glycol) diacrylate (PEGDA). Further, the precursors were photocrosslinked in conjunction with a foaming process, resulting in the generation of a porous network HGL. Only 34% of the VEGF was released from the HGLs over a 13-day period. In vitro, the released VEGF was shown to stimulate the development of human umbilical vein endothelial cells (HUVECs) along with the formation of capillary networks. Results reveled that, the induction of angiogenesis was more than doubled when VEGF-laden HGLs were compared to PEG HGLs treated with fresh VEGF. Similar strategy was also adopted by Jun et al. for fabrication of photo-induced HGL of APMMA-conjugated HEP with di-acryloyl Pluronic for bFGF delivery [21]. Similarly, Goh et al. used photo-polymerization to develop a HEP-based HGL sheet that was loaded with human epidermal growth factor (hEGF) and thiolated HEP to promote wound healing [22]. Shah et al. 2011 fabricated HGL by photo-patterning of thiolated HEP and diacrylated PEG (PEG-DA) as shown in Fig. 5.4. Results reveled that photo-crosslinked HGLs synthesized faster and were tougher than HGLs generated by the competitive Michael addition process (Fig. 5.4). The amount of albumin generated by cells grown near HEP-based HGLs was fourfold more than that produced by cells grown near inert PEG HGLs at Day 7 [23].

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FIGURE 5.3 Schematic representation of HEP-APMMA conjugates. Reprinted from Oliviero O, Ventre M, Netti PA. Functional porous hydrogels to study angiogenesis under the effect of controlled release of vascular endothelial growth factor. Acta Biomater 2012;8(9):3294e3301, Copyright (2012), with permission from Elsevier.

5.2.2 HEP-based HGLs formed via Michael-type addition reactions One of the most common crosslinking methods for HGLs is the Michael-type addition reaction, which happens in aqueous medium, at ambient temperature, and physiological pH. Michael-type addition processes like the thiol-ene reaction were often used to make HEP-based HGLs by attaching ene [24e31] or thio [22,23,32e44] groups to HEP molecules. Recent research carried out by Hesse et al. resulted in the development of a matrix metalloprotease (MMP)-responsive starPEG/HEP HGL for cartilage regeneration [24]. The EDC/sulfo-NHS activation of HEP resulted in the reaction of the carboxyl groups with maleimide amine, which ultimately led to the development of maleimide group-conjugated HEP, which was subsequently crosslinked by Michael-type addition to generate HGL. Chwalek et al. fabricated an HGL platform based on starPEG-HEP that may exhibit a variety of adhesion and breakdown sites, growth factors, and cells [26]. A biohybrid HGL was prepared by Nowak et al. using glycosaminoglycan HEP, starPEG, and MMP-cleavable crosslinkers to investigate the biophysical and biochemical cues that promote human MEC development (Fig. 5.5) [27]. Their research offers a flexible framework for studying mammary epithelial tissue morphogenesis in a chemically defined and precisely programmable 3D in vitro milieu. These findings allow the researchers to explore the biophysical and biochemical components of the mammary gland’s biology. Aside from ene group-conjugated HEP precursors, thiol group-conjugated HEP precursors were also frequently used to make HEP-based HGLs with diverse architectures by Michael-type addition reaction. For instance, Foster et al. synthesized a HEP-based HGL sandwich by an UV-induced ene-thiol linkage process of diacrylated poly(ethylene glycol) and thiolated precursors, which provided a culture environment for primary rat hepatocyte maintenance [45]. The HEP-PEG HGL was flexible, with adjustable mechanical characteristics, changeable HEP concentration, and photo-induced gel disintegration; it might be used as a possible method for retaining differentiated and polarized primary hepatocytes. In addition to the sandwich construction, they created HGL microstructures for the culture of primary hepatocytes [37,46]. The HEP HGL-based microwells were created using a mix of micromolding and microcontact printing. The HEP gel walls of the microwells boosted the expression of hepatic phenotypes, implying that the gel would be a potential hepatic

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FIGURE 5.4 (A) A diagrammatic illustration of the UV-initiated thiol-ene reaction among thiolated HEP and acrylated polyethylene glycol. Gelation kinetics (B) and gel strengths (C) of synthesized HGLs. Reprinted from Shah SS, Kim M, Cahill-Thompson K, Tae G, Revzin A, Chitosan-based hydrogels for controlled, localized drug delivery. Soft Matter 2011;7(7):3133e3140, Copyright (2011), with permission from Royal Society of Chemistry.

habitat for differentiation of stem cells. Michael-type addition HGLs can be used for a variety of biomedical purposes. HEP was added to an HA-based HGL for matrix-assisted cell transplantation (MACT) in order to assist the trophic activities of CPCs via endothelial cell differentiation and vascular-like tubular network creation of CPCs (Fig. 5.6) [42].

5.2.3 Amide coupling for crosslinking of HGLs HEP has carboxyl groups that can interact with amino groups to produce amide groups by free radical polymerization [15,16,20,47e51]. Furthermore, by directly creating crosslinking networks, the amide bond formation mechanism might be exploited to make HEP-based HGLs using simpler production procedures [52e56]. Seib et al. manufactured a PEG-containing HEP-based HGL containing doxorubicin as a targeted breast cancer therapeutic transporter by an amide bond formation process [53]. According to Borg and colleagues, direct chemical crosslinking (EDC/sulfo-NHS chemistry) was used to create an amino-terminated 4-arm (starPEG) and HEP HGL in star shape (starPEG) [54]. A formulated porous starPEG-HEP HGL not only house pancreatic islets and mesenchymal stromal cells (MSCs) accessory cells but also providing mechanical protection to them. The MSCs might well biofunctionalize the HGL scaffold by secreting ECM proteins, making it a promising vehicle for pancreatic islet housing. The same team created a PEG-HEP multilayer HGL covering that contained silver to improve hemocompatibility while also reducing microbe growth [55]. With human whole blood incubated, the multilayered HGL coatings demonstrated good plasmatic

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FIGURE 5.5 (A) Schematic representation of in vitro mammary epithelial cell (MEC) morphogenesis. (B) The study plans to evaluate MEC development in an HGL matrix. (C) Polarized MEC acini formation. Reprinted from Nowak M, Freudenberg U, Tsurkan M V, Werner C, Levental KR. Modular GAG-matrices to promote mammary epithelial morphogenesis in vitro. Biomaterials 2017;112:20e30, Copyright (2017), with permission from Elsevier.

coagulation, platelet activation, and hemolysis performance, and long-term antiseptic efficiency against Escherichia coli and the Staphylococcus epidermidis strain. An adaptable starPEG-HEP HGL platform was used by Watarai et al. to cultivate human dermal fibroblasts with different mechanical and biochemical capabilities. Fibroblast adhesion, spreading, proliferation, matrix deposition, and remodeling were enhanced in HGLs containing RGD peptides as compared to HGLs without alterations (Fig. 5.7). Myofibroblast differentiation of fibroblasts was induced by the stimulation of collagen type I, ED-A-fibronectin expression as well as alpha smooth muscle actin and palladin incorporation into F-actin stress fibers

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FIGURE 5.6 Representation of fabrication of HyA HGLs including the cell-adhesive RGD peptide and HEP by the Michaeltype addition reaction. Reprinted from Jha AK, Tharp KM, Ye J et al., Enhanced survival and engraftment of transplanted stem cells using growth factor sequestering hydrogels. Biomaterials, 2015;47:1e12, Copyright (2015), with permission from Elsevier.

when TGFb1 was reversibly coupled to starPEG-HEP HGLs for many days. Customized starPEG-HEP HGLs may be useful in promoting wound healing by rising the proliferation andand differentiation of human dermal fibroblasts, as a result [56]. Additionally, HEP may be grafted onto a variety of polymer matrixes via the formation of amide bond. This process was used for the manufacture of a variety of HEP-based HGLs [7,57e59]. Nagai et al. immobilized HEP in collagen and then gelled it at 37
C to create an extracellular matrix-modeled HGL that was used to create an in vivo-like environment and an in situ assessment method for HGL-embedded cell responses [57]. Tran et al. created a quercetin-conjugated HEP HGL for improved blood compatibility [58].

5.2.4 Enzyme-mediated crosslinked HGLs Enzyme-mediated crosslinking of HGLs has been widely investigated and offers great prospects [60]. To recreate biological creatures’ ECMs, many enzymes operating at physiological pH and temperature settings have been found and created [61,62]. Enzymes offer distinct advantages for HGL production in addition to their crosslinking activity. 3D substrate shapes that suit enzymes for binding are recognized by enzymes. Because of the “substrate specificity,” enzyme activity may be tightly controlled without causing undesirable side effects. Kiick K. L. (2008) assembled heparinized HGLs in a noncovalent manner using a PEGeHBP or HEP-binding proteins. In addition to modifying these materials’ viscoelastic properties, such as by increasing or decreasing the LMWH:HBP ratios or polymer concentrations, it should be possible to immobilize other therapeutic proteins using these materials, which could lead to a wider variability of mechanical and biological properties. Sulfated peptides are one such glycosaminoglycan or HEP mimic. The elastic moduli of LMWH-functionalized HGLs might be used in a variety of soft-tissue engineering applications, and it is expected that a modulus might be easily altered using different branched and multiarm polymer structures [63].

5.2.5 Other covalent bonding approaches HGLs derived from HEP may be generated via alternative covalent bonding processes. Click chemistry is one of the alternative covalent crosslinking approaches used for fabrication of HGLs. Click chemistry reaction between dibenzocyclooctyne (DBCO) and azide analogs was utilized by Adil et al. for the development of functionalized hyaluronic acid HGLs with HEP and RGD. Afterward, these stem cells were used for in vitro development and insertion of human pluripotent stem cell-derived neural progenitors into the central nervous system. Afterward, these stem cells were used for in vitro development and transplantation of human pluripotent stem cell-derived neural progenitors into the central nervous system (Fig. 5.8) [64]. HGL particles (HGPs) decorated with HEP were also synthesized using the DVS crosslinking method, which was applied in conjunction with an inverse emulsion polymerization process. HGPs decorated with HEP were also synthesized using the DVS crosslinking method, which was applied in conjunction with an inverse emulsion polymerization process [65]. HGPs created using the covalently immobilized HEP might be used as an appealing prospect for the release of bone morphogenetic protein-2 (BMP-2) and, therefore, utilized for cartilage repair and regeneration. Due to the fact that dopamine (DA), a chemical inspired by mussel adhesive protein can form irreversible covalent connections

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FIGURE 5.7 Schematic representation of cell-instructive starPEG-HEP HGLs to induce myofibroblast differentiation. (A) TGFb is reversibly conjugated to the HGL scaffold. (B) Network structure of the starPEG-HEP HGL. Reprinted from Watarai A, Schirmer L, Thönes S et al. TGFb functionalized starPEG-heparin hydrogels modulate human dermal fibroblast growth and differentiation. Acta Biomaterialia 2015;25:65e75, Copyright (2015), with permission from Elsevier.

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FIGURE 5.8 (A) HGL crosslinked with polyethylene glycol (PEG) diazides containing RGDs and HEPs. (B) Gelation kinetics determine: elastic (blue) and loss (red) moduli of HA HGLs functionalized with RGD and HEP. (C) Schematic representation for making of mDA neurons from hESCs. Reprinted from Adil MM, Vazin T, Ananthanarayanan B et al. Engineered hydrogels increase the post-transplantation survival of encapsulated hESC-derived midbrain dopaminergic neurons. Biomaterials 2017;136:1e11, Copyright (2017), with permission from Elsevier.

to solid surfaces in alkaline medium, DA self-polymerization has attracted considerable interest for use in surface modification [66e68]. Deng and colleagues created a highly flexible HGL tube by attaching a DP-grafted HEP to an alginate/polyacrylamide double-network HGL with a mussel-inspired coating and postfunctionalizing it for prospective use as an artificial blood artery [69]. Other compounds, such as hydrazide conjugated derivates, might be used to create HEP-based HGLs [70]. Tae’s group created a crosslinked HEP-based HGL with HEP-binding domains (Hep-ADH). This HGL might be used as an affinity-based controlled release system for growth factors (VEGF) [52]. HEP can be covalently bonded to HGL matrixes to create HGLs that are stable. In spite of these shortcomings, the chemical method is still a viable option since it is less invasive and less likely to introduce harmful compounds into the human body and the environment. As an ecofriendly solution with a straightforward procedure, mussel-inspired coating has the potential to be a viable alternative to the chemical procedures described above.

5.3 Physically crosslinked HEP HGLs Entanglements among dynamic macromolecular species between polymer chains cause physical crosslinking [71]. Physical HGLs eliminate the inclusion of hazardous crosslinking chemicals, which may possibly compromise the stability and effectiveness of the therapeutic substances contained inside HGLs. In addition, the shear-thinning ability and ease with which noncovalent connections may be easily reestablished make it possible to inject physically crosslinked HGLs [72]. Physical HGLs eliminate the inclusion of hazardous crosslinking chemicals, which may possibly compromise the integrity and bioactivity of encapsulated medicinal medicines. In addition, the shear-thinning ability and ease with which noncovalent connections may be easily reestablished make it possible to inject physically crosslinked HGLs [63].

5.3.1 Electrostatic interaction HGLs HGLs made from HEP may also be produced via electrostatic interactions. Electrostatic interactions between HEP, chitosan (CS), and poly(gamma glutamic acid) resulted in the formation of composite HGLs, which were then loaded with an

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FIGURE 5.9 Diagrammatic representation for fabrication of HELP gel of HEP-protein complex and laponite. Reprinted from Ding X, Gao J, Wang Z, Awada H, Wang Y. A shear-thinning hydrogel that extends in vivo bioactivity of FGF2. Biomaterials 2016;111:80e89, Copyright (2016), with permission from Elsevier.

antioxidant to treat diabetic wounds. After loading with superoxide dismutase, the HGLs demonstrated cytocompatibility with fibroblasts and the potential to speed healing of wounds by increasing the amount of collagen that is deposited and wound sealing [73]. HEP and Laponite-based rapid gelation HGL was fabricated by Ding et al. (Fig. 5.9). They retain HEP’s inherent affinity for numerous proteins in HGLs and shield FGF2 from proteolytic degradation, thereby releasing it with preserved bioactivity to stimulate angiogenesis in vitro [74]. Kumar et al. 2015a used ionic contact and hydrogen bonding to create an HGL for drug administration utilizing multivalent compounds (HEP, trypan, clodronate, phosphate, and suramin) [75].

5.3.2 Hosteguest interaction HGLs HEP-based HGLs can be formed by nesting molecules with functional groups together, which can be accomplished through hosteguest interaction. Diang et al. employed a dual dynamic bond of hosteguest and electrostatic contact to create a biodegradable, injectable, and biodegradable HEP HGL, as shown in Fig. 5.10 [76]. When the PEG-g-PEEAD polymer was combined with the anticoagulant HEP and the alpha-cyclodextrin (alpha-CD), the supramolecular HGL network was formed in a short period of time with shear thinning capabilities. During the synthesis, PEG was converted to

FIGURE 5.10 (A) Schematic representation of b mPEGx-g-PEAD, a-CD, and HEP. (B) Representation of fabricated hydrogel without HEP. (C) Representation of fabrication of hydrogel with HEP utilizing hosteguest chemistry as initial crosslinking mechanism and electrostatic interactions as secondary crosslinking mechanism. Reprinted from Ding X, Gao J, Awada H, Wang Y. Dual physical dynamic bond-based injectable and biodegradable hydrogel for tissue regeneration. J Mater Chem 2016;4(6):1175e1185, Copyright (2016), with permission from Royal Society of Chemistry.

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poly(ethylene aspartate diglyceride) after being linked with arginine and monocarboxylic acid (PEAD). They demonstrated that the HGL technology may be used for ischemic tissue regeneration and wound healing by releasing FGF2 at a steady pace. Ma et al. fabricated supramolecular HGL networks by using the hosteguest interaction among a-cyclodextrin and amino-terminated polyethylene glycol methyl ether-conjugated HEP (Fig. 5.11A) [77]. Because of its high potential as an injectable matrix for the encapsulation and release of the model protein bovine serum albumin as well as its regulated release profile for the conjugated HEP, the HGL produced demonstrated strong anticoagulant and blood-compatible properties (Fig. 5.11B).

5.3.3 Hydrogen bonding (growth factor-crosslinked HGLs) In order to create crosslinking networks, hydrogen bonding among the hydrogen and electronegative atoms has been used in several studies to construct HEP-based HGLs [78,79]. When HEP sodium salt and RAD16-I peptide self-assembled, they created a HEP-based HGL [80]. As a result, the HGL might be exploited as a potential solution for adipose-derived stem cells (ADSCs) survival and chondrogenic commitment, as demonstrated in this study. An intermolecular hydrogen bond between the PVA side chains was used to construct a crosslinked HGL containing HEP embedded in a PVA HGL under high hydrostatic pressure (HHP) [78]. The HEP-PVA complex HGL demonstrated the ability to inhibit clot formation and demonstrated potential as a vascular access device, while the HHP technique demonstrated significant potential for the development of specifically functionalized PVA HGLs. CTT and its peptide counterpart CTTHWGFTLC could also be used to create an injectable, fast gelling thermosensitive and degradable HGL to allow for the controlled release of CTT at low molecular weight by using hydrogen bonding interactions [79].

5.3.4 Hydrophobic interaction HGLs Researchers discovered that hydrophobic domains containing polymers might lead to supramolecular self-assembly caused by hydrophobic interactions between and within polymers., and that this may possibly be used to create HEP-based HGLs, according to the findings [31,81,82]. In order to create crosslinker-free HGLs, GAGs-modified amphiphilic multiblock copolymers (Pluronics, HP, and HA) were introduced to self-assemble HGLs including triblock poly(ethylene oxide)polypropylene oxide copolymers [31]. Ma et al. synthesized Hep-F-127 as an amphiphilic polymer (Fig. 5.12) and were loaded into micelles with HEP, while G-CSF was introduced into hydrophilic shell and CPT into the hydrophobic core. G-CSF and camptothecin were successfully encapsulated and released from supramolecular HGL/micelle composite (CPT). The HGL/micelle composite retained the G-CSF and CPT’s biological action with encapsulation and delayed release characteristic by adding more alpha-CD to the system [82]. Depending on the hydrophobicity of the manufactured substrate, gelation might be affected, which in turn could have an impact on the degradation of HGLs [31].

5.3.5 Other physical interactions HEP-based HGLs might be prepared using a wide range of physical processes, including those outlined in this article. HEP-based HGLs that respond to stimuli have been developed and tested for therapeutic drug delivery. A thermoresponsive injectable HGL network for lysozyme distribution was developed using a HEP-bearing poly(-caprolactoneco-lactide)-b-poly (ethylene glycol)-b-poly(-caprolactone-co-lactide) (Hep-PCLA) [83]. They could form gels under physiological circumstances (37
C) even though the Hep-PCLA conjugates were free-flowing in aqueous solutions at 25
C (Fig. 5.13). Hydrophobic and ionic interactions loaded the model protein lysozyme into the HGL and allowed it to be gently released even during the first time. In the poloxamer triblock copolymer, ethylene oxide (EO) and propylene oxide (PO) are present, and the hydrophobic character of the PO block causes temperature-induced aggregation [84]. Thus, Zhao’s team synthesized a variety of HEP-poloxamer (HP) thermosensitive HGLs for a wide range of applications. They used a HEP-poloxamer HGL to enhance the quality and safety of vascular anastomosis in a rabbit model, which was successful [84]. HP HGL had great potential for enhancing vascular anastomosis quality and safety by combining the advantages of HEP and poloxamer. It was shown that raising the temperature of aFGF bridging HEP-poloxamer HGLs used to treat spinal cord injury (SCI) to 37
C resulted in prolonged release of aFGF while maintaining bioactivity, which might be used to treat SCI in the future.

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FIGURE 5.11 (A) Diagrammatic representation for the fabrication of Hep-MPEG/R-CD supramolecular HGLs; (B) images of in vitro blood clotting protocol. Reprinted from Ma D, Tu K, Zhang LM. Bioactive supramolecular hydrogel with controlled dual drug release characteristics. Biomacromolecules 2010;11(9):2204e2212, Copyright (2010), with permission from American Chemical Society.

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FIGURE 5.12 Schematic representation of method of fabrication of Hep-F-127. Reprinted from Ma D, Zhang H bin, Tu K, Zhang LM. Novel supramolecular hydrogel/micelle composite for co-delivery of anticancer drug and growth factor. Soft Matter 2012;8(13):3665e3672, Copyright (2012), with permission from Royal Society of Chemistry.

5.4 Combined interaction (duel crosslinked) HGLs Chemical bonding and physical conjugation were used to produce HEP-based HGL networks, and the combination of chemical and physical contacts was used to further develop dual-crosslinked networks, for improved HGL formation [21,85]. A biodegradable HEP/Pluronic composite HGL for growth factor administration was developed by Yoon and coworkers using the sol-to-gel transition and photo-initiated ene crosslinking techniques. HEP and diacrylate were ended by a vinyl group that had been conjugated. In the gelation procedure, bFGF was incorporated into the Pluronic F127 HGLs. HUVEC proliferation was shown to be enhanced by the use of HEP-based HGLs, which might be used to promote angiogenesis by loading and releasing bFGF over an extended period of time [21]. It is also possible to make dualcrosslinked HEP-based HGLs by first using a chemical-bonded technique and then using a physical-conjugated technique as shown in Fig. 5.14A. But Ding et al. achieved their goal by using biological HEP as the scaffold material and an in-situ HGL network with ultra-compliant properties in order to accomplish their mission. HEP-based HGL networks were created via chemical and physical conjugation, and a combination of these two methods was employed to further build dual-crosslinked networks. Results showed such duel crosslinked HGLs have electronic/ionic conductivity and good mechanical properties [85]. To construct photo-crosslinkable HEP scaffold, HEP was esterified with methacrylic anhydride (MA). Then, aniline was doped with aniline (electrostatic interaction) using a template with a controlled microstructure. There were no further surface treatments needed to sustain the adhesion, proliferation, and differentiation of murine

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FIGURE 5.13 Diagrammatic representation of sol-to-gel transition of Hep-PCLA HGLs. Reprinted from Sim HJ, Thambi T, Lee DS. Heparin-based temperature-sensitive injectable hydrogels for protein delivery. J Mater Chem 2015;3(45):8892e8901, Copyright (2015), with permission from Royal Society of Chemistry.

myoblasts in their HEP/polyaniline HGL, thus demonstrating its capability to be employed as a cell culture substrate. To construct photo-crosslinkable HEP scaffold, HEP was esterified with MA. Then, aniline was doped with aniline (electrostatic interaction) using a template with a controlled microstructure. There were no further surface treatments needed to sustain the adhesion, proliferation, and differentiation of murine myoblasts in their HEP/polyaniline HGL. Thus the finding concluded that it has the potential to be used as a substrate for cell culture [85]. By using the Michael addition reaction between acrylate groups and thiol groups, it is possible to synthesize covalently crosslinked, CS-based HGLs that are composed of PEGDA and thiolated polymers like CS, gelatin, and HEP under physiological conditions. Microparticles of PECL micelles with double bonds are introduced into the HGL network to create micelle-mediated dual crosslinked HGLs from CS-based HGLs that have been chemically crosslinked further (Fig. 5.14B). The inclusion of bi-acrylated PECL micelles into CS-based HGLs is regulated by a Michael addition reaction, which occurs when acrylate groups evenly distributed on the surface of bi-acrylated PECL micelles and thiol groups present in HGL networks are added in the same proportion. As shown by the results, there are fewer free thiol groups in dry HGL than there are in thiolated CS, suggesting that a Michael addition reaction takes place among the thiol and acrylate groups. The PECL micelles are incorporated into the HGL network through interfacial polymeremicelle interactions, which allows for a development of pseudo- or noncovalent crosslinks in the HGL network. Given the relatively rapid initial release of these dual crosslinked HGLs, they were found to release IMC and/or bFGF at a slower rate than traditional HGLbased drug delivery systems until a steady state was achieved [86].

5.5 Smart HGLs or stimuli-responsive HEP HGLs A variety of physical and chemical changes occur in stimulus-responsive HGLs in response to environmental stimuli. These include swelling and shrinking, degradation of the HGL, and transition from the solegel to the liquid phase. The use of a wide range of chemically labile linkages to govern HGL disintegration in retort to environmental stimuli such as pH, enzyme, and light has become increasingly popular in recent years, and this has the potential to deliver a mechanism for therapeutic molecular release. Although the field of stimuli-responsive HEP HGLs is still in early stages, several strategies have been developed. Some of the first environmentally sensitive HEP-based HGLs were developed through the use of a

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FIGURE 5.14 (A). Diagrammatic representation for fabrication of dual crosslinked Hep-MA/PANI HGL. Reprinted from Ding H, Zhong M, Kim YJ, Pholpabu P, Balasubramanian A, Hui CM, He H, Yang H, Matyjaszewski K, Bettinger CJ. Biologically derived soft conducting hydrogels using heparindoped polymer networks. ACS Nano 2014;8(5):4348e4357, Copyright (2014), with permission from American Chemical Society]; (B) Diagrammatic representation for the dual cross-linked HGL of chitosan PECL micelles. [Reprinted from Wen Y, Li F, Li C, Yin Y, Li J. High mechanical strength chitosan-based hydrogels cross-linked with poly(ethylene glycol)/polycaprolactone micelles for the controlled release of drugs/growth factors. J Mater Chem 2017:5(5):961e971, Copyright (2017), with permission from Royal Society of Chemistry.

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combination of covalent and noncovalent crosslinking techniques. These HGLs exhibited reversibly temperatureresponsive rheological behavior under environmental conditions [87e89]. Multiple other strategies been designed and constructed since then, as will be discussed further below.

5.5.1 Enzymatically responsive HGLs In human platelets, skeletal muscle, and smooth muscles, the heparinase enzyme cleaves the anticoagulant HEP chains at the 3-O-sulfated glucosamine residue, decreasing the anticoagulant properties of HEP and enhancing the signaling activity of FGF-2. It has also been demonstrated that additional enzymes, such as heparinase and heparitinase, are effective in HEP degradation, by cleaving the HEP molecule into di- or tri-saccharide units. In human platelets, skeletal and smooth muscles, and adipose tissue, the enzyme heparinase degrades HEP chains at the 3-O-sulfated glucosamine residue, decreasing HEP’s anticoagulant properties while growing its FGF-2 signaling activity. Moreover, additional enzymes such as heparinase and heparitinase are shown to be proficient of degrading HEP [90]. HGLs containing HEP/PVA have been shown to degrade enzymatically when incubated with platelet extract (PE) [91]. It was discovered that the PE, which includes heparinase, may release covalently crosslinked HEP fragments from HGLs. The release of HEP from PE-treated HGLs supernatant was validated by an increase in proliferation index and plasma clotting time in PE-treated HGLs correlated to non-PE-treated HGLs. HGL scaffolds that are heparinized and applied to an injured region might have beneficial outcomes, according to this HEP/PVA HGL system. Incorporating peptides into HGL networks, which may be broken by cell-produced proteases, is a frequently used alternative strategy for designing enzymatically degradable HGLs. MMPs breakdown a wide range of peptide patterns, which may then be easily integrated into HGLs. Incorporating peptides into HGL networks, which may be broken by cell-produced proteases, is a frequently used alternative strategy for designing enzymatically degradable HGLs. MMPs breakdown a wide range of peptide patterns, which may then be easily integrated into HGLs [46,92e98]. The interaction of a multifunctional polymer (e.g., PEG-acrylates or PEG-vinyl sulfones) with thiols from the cysteine residues of the peptides results in Michael-type addition reactions, which are probably the extensively employed technique of conjugating enzyme-cleavable peptide sequences [99]. Dextran tyramine and HEP tyramine conjugates were crosslinked by horseradish peroxidase (HRP) enzyme in order to produce injectable HGLs containing HEP [100]. These materials demonstrated variable gelation kinetics (ca. 30e350s) produced by modifying the HRP concentration, as well as changeable mechanical characteristics (G0 , 3.6e48 kPa) achieved by changing the HGL composition (HEP content, 100% to 0%, respectively). HEP inclusion into the HGLs significantly improved HGL swelling, allowing nutrient transfer for cell growth. These materials demonstrated variable gelation kinetics (ca. 30e350s) produced by modifying the HRP concentration, as well as changeable mechanical characteristics (G0 , 3.6e48 kPa) achieved by altering the HGL concentration (HEP content, 100 to 0 w/w %, respectively). HEP inclusion into the HGLs significantly improved HGL swelling, allowing nutrient transfer for cell growth. Bovine chondrocyte survival, proliferation, and matrix synthesis were increased by the incorporation of bovine chondrocytes into these HGLs, suggesting the promise of DexeTA/HepeTA HGLs as matrices for cartilage tissue engineering. Crosslinking with the HRP enzyme is also used to create HEP HGL surfaces for coating metallic biomaterials with blood compatibility, which have significantly improved blood compatibility and lower fibrinogen absorption [101].

5.5.2 Glutathione-responsive HGLs Reduction-sensitive bonds like disulfide linkages have been used in the manufacture of HGLs because the tripeptide GSH is found in cells and tumor microenvironments. A significant increase in GSH concentration in intracellular compartments (2e10 mM in cells compared to plasma concentrations of 2e20 M) allows GSH-sensitive HGLs to persist steady external side of cells while quickly degrading in intracellular compartments, allowing for targeted delivery of medications to specific intracellular compartments [64e67,69]. Due to a lack of control over cleavage kinetics (T1/2 ¼ 8e45 min), disulfide bonds are being reduced, which has the potential to restrict the therapeutic benefits of biomolecules embedded in GSH-sensitive HGLs. The stability of the adducts produced by thiol-maleimide Michael-type reactions has long been recognized as a factor in in situ HGL formation; however, Faure and coworkers have taken advantage of the underappreciated capability of these processes to be reversed in the existence of GSH [68]. Maleimide-thiol adducts created via the Michael-type addition were found to experience retro- and exchange process in the existence of GSH. The addition of various thiols (4-mercaptophenylacetic acid (MPA), N-acetylcysteine, or 3-mercaptopropionic acid (MP)) to Nethylmaleimide resulted in the formation of small-molecule adducts in the beginning (NEM). With glutathione treatment, the retro reaction occurred in the adducts for 20e80 h, with conversion ranging from 20% to 90%. The arylthiol adducts exhibited the highest rate and degree of conversion during a retro reaction, moreover MPA-NEM adduct exhibiting

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approximately 85% conversion after 70 h of incubation in the laboratory. It should be noted that the kinetics and conversion of the N-acetylcysteine-NEM conjugation were significantly slower, while the MP-NEM adduct had virtually no activity in the retro reaction. It should be noted that the kinetics and conversion of the N-acetylcysteine-NEM conjugation were significantly slower, while the MP-NEM adduct had virtually does not shown a retro reaction [70]. In comparison to disulfide-crosslinked HGLs, the PEG-LMWH HGLs made via thiol-maleimide Michael reactions had a 10-fold reduced degradation rate, suggesting that the Michael adducts might provide an alternative to disulfide-crosslinked HGLs for longer-term drug administration. Matrix remodeling in tissue engineering applications and targeted/controlled distribution of medicines to diseased locations where GSH is overproduced might be achieved by using these methodologies on a wide range of biomaterials (e.g., up to 10-fold greater levels of GSH in tumors with respect to adjacent normal tissues) [102].

5.6 Conclusion and prospective Aside from their versatility in terms of biocompatibility and therapeutic efficacy, HEP-based HGLs have significant potential in a wide range of biomedical applications. For instance, HEP-based HGLs have demonstrated favorable biocompatibility and therapeutic efficacy, which has the potential to have substantial repercussions in a range of applications including growth factor delivery and cell carriers. This chapter focuses on HEP-based HGLs, which can be produced in a variety of ways. Crosslinking networks in HEP-based HGLs can be created using a variety of methods, including chemical covalent bonding, physical conjugation, or a combination of the two. Methods such as click chemistry, divinyl sulfone crosslinking, and mussel-inspired coating are among those highlighted. In contrast to chemically crosslinked HGLs, physically conjugated HGLs can be formed more easily through hoste guest or electrostatic interactions, as well as hydrogen bonding or hydrophobic interactions, for example. HEP-based HGLs have the potential to be used as bioactive HGLs due to the fact that they belong to a significant family of biomedical HGLs. Because HEP-based HGLs that respond to stimuli (in situ gelling during HGL preparation, swelling or degradation under stimulation for controlled drug delivery) are uncommon, this chapter goes into great detail about how to make stimuliresponsive HEP-based HGLs. HEP-based HGLs must be designed to be stable while being excreted from the body in a manner that prevents the formation of any unwanted byproducts if they are to be used in clinical settings. As these challenges are overcome, it is expected that HEP-based HGLs will play an increasingly important role in the treatment of a wide range of diseases and damaged tissues.

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[78] Negishi J, Nam K, Kimura T, et al. Fabrication of a heparin-PVA complex hydrogel for application as a vascular access. J Biomed Mater Res Part B Appl Biomater 2014;102(7):1426e33. [79] Fan Z, Fu M, Xu Z, et al. Sustained release of a peptide-based matrix metalloproteinase-2 inhibitor to attenuate adverse cardiac remodeling and improve cardiac function following myocardial infarction. Biomacromolecules 2017;18(9):2820e9. [80] Fernández-Muiños T, Recha-Sancho L, López-Chicón P, Castells-Sala C, Mata A, Semino CE. Bimolecular based heparin and self-assembling hydrogel for tissue engineering applications. Acta Biomater 2015;16(1):35e48. [81] Gajewiak J, Cai S, Shu XZ, Prestwich GD. Aminooxy pluronics: synthesis and preparation of glycosaminoglycan adducts. Biomacromolecules 2006;7(6):1781e9. [82] Ma D, Zhang H, Tu K, Zhang LM. Novel supramolecular hydrogel/micelle composite for co-delivery of anticancer drug and growth factor. Soft Matter 2012;8(13):3665e72. [83] Sim HJ, Thambi T, Lee DS. Heparin-based temperature-sensitive injectable hydrogels for protein delivery. J Mater Chem B 2015;3(45):8892e901. [84] Zhao YZ, Lv HF, Lu CT, Chen LJ, Lin M, Zhang M, et al. Evaluation of a novel thermosensitive heparin-poloxamer hydrogel for improving vascular anastomosis quality and safety in a rabbit model. PLoS One 2013;8(8):e73178. [85] Ding H, Zhong M, Kim YJ, Pholpabu P, Balasubramanian A, Hui CM, et al. Biologically derived soft conducting hydrogels using heparin-doped polymer networks. ACS Nano 2014;8(5):4348e57. [86] Wen Y, Li F, Li C, Yin Y, Li J. High mechanical strength chitosan-based hydrogels cross-linked with poly(ethylene glycol)/polycaprolactone micelles for the controlled release of drugs/growth factors. J Mater Chem B 2017;5(5):961e71. [87] Seal BL, Panitch A. Viscoelastic behavior of environmentally sensitive biomimetic polymer matrices. Macromolecules 2006;39(6):2268e74. [88] Seal BL, Panitch A. Physical matrices stabilized by enzymatically sensitive covalent crosslinks. Acta Biomater 2006;2(3):241e51. [89] Kyung JJ, Panitch A. Interplay between covalent and physical interactions within environment sensitive hydrogels. Biomacromolecs 2009;10(5):1090e9. [90] Ernst S, Langer R, Cooney CL, Sasisekharan R. Enzymatic degradation of GlycosaminogIycans. Crit Rev Biochem Mol Biol 1995;30(5):387e444. [91] Nilasaroya A, Martens PJ, Whitelock JM. Enzymatic degradation of heparin-modified hydrogels and its effect on bioactivity. Biomaterials 2012;33(22):5534e40. [92] Lutolf MP, Lauer-Fields JL, Schmoekel HG, et al. Synthetic matrix metalloproteinase-sensitive hydrogels for the conduction of tissue regeneration: engineering cell-invasion characteristics. PNAS 2003;29. www.pnas.orgcgidoi10.1073pnas.0737381100. [93] Kraehenbuehl TP, Ferreira LS, Zammaretti P, Hubbell JA, Langer R. Cell-responsive hydrogel for encapsulation of vascular cells. Biomaterials 2009;30(26):4318e24. [94] Patterson J, Hubbell JA. Enhanced proteolytic degradation of molecularly engineered PEG hydrogels in response to MMP-1 and MMP-2. Biomaterials 2010;31(30):7836e45. [95] Fairbanks BD, Schwartz MP, Halevi AE, Nuttelman CR, Bowman CN, Anseth KS. A versatile synthetic extracellular matrix mimic via thiolnorbornene photopolymerization. Adv Mater 2009;21(48):5005e10. [96] Anderson SB, Lin CC, Kuntzler DV, Anseth KS. The performance of human mesenchymal stem cells encapsulated in cell-degradable polymerpeptide hydrogels. Biomaterials 2011;32(14):3564e74. [97] Tsurkan Mv, Levental KR, Freudenberg U, Werner C. Enzymatically degradable heparin-polyethylene glycol gels with controlled mechanical properties. Chem Commun 2010;46(7):1141e3. [98] Tsurkan Mv, Chwalek K, Levental KR, Freudenberg U, Werner C. Modular StarPEG-heparin gels with bifunctional peptide linkers. Macromol Rapid Commun 2010;31(17):1529e33. [99] Burdick JA, Murphy WL. Moving from static to dynamic complexity in hydrogel design. Nat Commun 2012;3. [100] Jin R, Moreira Teixeira LS, Dijkstra PJ, van Blitterswijk CA, Karperien M, Feijen J. Chondrogenesis in injectable enzymatically crosslinked heparin/dextran hydrogels. J Contr Release 2011;152(1):186e95. [101] Joung YK, You SS, Park KM, Go DH, Park KD. In situ forming, metal-adhesive heparin hydrogel surfaces for blood-compatible coating. Colloids Surf B Biointerfaces 2012;99:102e7. [102] Yeh CC, Hou MF, Wu SH, et al. A study of glutathione status in the blood and tissues of patients with breast cancer. Cell Biochem Funct 2006;24:555e9.

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Chapter 6

Xanthan gum and its composite-based hydrogels Kaushik Mukherjee1, Pallobi Dutta1, Hemant Ramachandra Badwaik2 and Tapan Kumar Giri1 1

Department of Pharmaceutical Technology, Jadavpur University, Kolkata, West Bengal, India; 2Shri Shankaracharya Institute of Pharmaceutical

Sciences and Research, Bhilai, Chhattisgarh, India

6.1 Introduction Biomaterials have been widely exploited for pharmaceutical and biomedical applications including regenerative medicine, drug delivery, biomedical devices, etc. [1]. Biopolymers are an essential class of materials such as polysaccharides, polynucleotides, and polypeptides [2]. Polysaccharide-based biopolymers are of special interest to the pharmaceutical industry because of their structural and functional diversity. Polysaccharide biopolymers include starch, cellulose, hyaluronic acid (HA), dextran, and hydrocolloids like xanthan, chitosan, alginate, etc. [3]. These polysaccharides have the advantage of being readily available, biocompatible, biodegradable, and widely accepted by regulatory authorities [4]. Xanthan gum (XG), one of such polysaccharides obtained from the bacterium Xanthomonas campestris, was discovered in the 1960s and used commercially since the 1970s [5e7]. Worldwide production of xanthan gum is 50,000 MT annually, and the bulk portion of it is utilized in the food and biomedical industry [8]. It is approved by the Food and Drug Administration (FDA) in 1968 for applications in food and pharmaceutical products. Since then the gum has become a part and parcel of the food and pharmaceutical industry, and industry experts are of the opinion that there will be a 15% growth in the annual production of xanthan gum, which will result in a business size of US$ 455.9 million by 2027 [9]. The unique properties of xanthan gum like its water solubility, rheological characteristics, stability in a wide range of temperatures, and pH make the gum a versatile candidate for application in the food industry [10,11]. It is mainly used as a stabilizer and thickener in the food industry. The gum is widely used as a suspending, thickening, emulsifying, and filmforming agent in the pharmaceutical industry [12,13]. The gum is also used in drug delivery fields [14]. Due to its excellent synergism with various other polymers, it is used in conjunction with them in diverse drug delivery applications. Xanthan gum has also been derivatized, and the derivatized form is used for various pharmaceutical and biomedical applications [15]. Xanthan gum is also used as an emulsifying and stabilizing agent in cosmetic products like lotions, facial creams, moisturizers, etc. Xanthan gum due to its shear thinning and gelling behaviors makes it an outstanding candidate for tissue engineering applications [16]. Xanthan gum also finds application in the petroleum industry and is an essential constituent of water-based drilling fluid compositions [17,18]. Herein this chapter, we will present a detailed discussion on the various aspects of xanthan gum which will include its source, chemical structure, and properties, and its particular applications in biomedical fields comprising drug delivery and regenerative medicine.

6.2 Xanthan gum 6.2.1 Source Xanthan gum is obtained as a fermented product from a gram-negative bacterium X. campestris [19,20]. Xanthomonas is a genus belonging to the family Pseudomonaceae. The bacteria occur as single straight rods, 0.7e1.8 mm long, and 0.4e0.7 mm wide. They are mobile, gram-negative, having a single polar flagellum (1.7e3 mm long) [20] (Fig. 6.1).

Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine. https://doi.org/10.1016/B978-0-323-95351-1.00017-X Copyright © 2024 Elsevier Inc. All rights reserved.

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FIGURE 6.1 TEM of Xanthomonas campestris (at 12,000 magnification). Reprinted from García-Ochoa F, Santos VE, Casas JA, Gómez E. Xanthan gum: production, recovery, and properties. Biotechnol Adv 2000;18:549e579, Copyright (2000), with permission from Elsevier.

6.2.2 Chemical structure and composition Xanthan gum is a heteropolysaccharide. The primary structure is made of repetitive pentasaccharide units made by one glucuronic acid unit, two mannose units, and two glucose units in a molar ratio of 2:2:2.8 [20]. The backbone comprises b-D-glucose units linked at 1 and 4 positions (Fig. 6.2). The chemical structure of the backbone is quite similar to that of

FIGURE 6.2 Chemical structure of Xanthomonas campestris. Reprinted from García-Ochoa F, Santos VE, Casas JA, Gómez E. Xanthan gum: production, recovery, and properties. Biotechnol Adv 2000;18:549e579, Copyright (2000), with permission from Elsevier.

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cellulose. The trisaccharide side chains consist of D-glucuronic acid units sandwiched between two mannose units which are linked at O-3 positions of every other glucose residue in the main chain. About half of the concluding D-mannose units carry a pyruvic acid residue linked by keto groups at 4 and 6 positions. The D-mannose unit connected to the main chain carries an acetyl group at the O-6 position. The presence of pyruvic and acetic acid units imparts an anionic character to the polysaccharide [21]. The pyruvate and O-acetyl units deprotonate at pH > 4.5, accentuate charge density along the xanthan chains, thus making xanthan gum amenable to divalent ion-mediated physical crosslinking [22,23]. The trisaccharide side chains are closely oriented with the polymer backbone. The ensuing rigid chain exists as a single, double, or triple helix, which forms complexes upon interaction with other polymer molecules [24,25]. The molecular weight of the gum is between 2
106 and 20
106 Da. The molecular weight of the gum is dependent upon its association with the chains, which in turn is dependent on the fermentation conditions used in the production of the gum.

6.2.3 Production A brief outline of the commercial xanthan gum production procedure is shown in Fig. 6.3. At first, a microbial strain of the bacterium is selected, which is then cultured by growing them in a liquid media or solid surfaces to obtain inoculums for large commercial bioreactors. The bacterium and production of the gum are dictated by various factors like the composition of the FIGURE 6.3 The xanthan gum production process. Reprinted from García-Ochoa F, Santos VE, Casas JA, Gómez E. Xanthan gum: production, recovery, and properties. Biotechnol Adv 2000;18:549e579, Copyright (2000), with permission from Elsevier.

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medium, the process of operation (continuous or batch), the type of bioreactor used, and the culture conditions (pH, temperature, etc.). The bioreactor should be well agitated and well aerated. The culture media should contain sources of carbohydrates (like glucose), nitrogen, and nutrient salts. Upon completion of the fermentation process, the broth is heated to destroy the bacteria and the gum is obtained by precipitating with isopropyl alcohol. It is then dried, milled, and packed for further use [26].

6.2.4 Physicochemical and rheological properties Xanthan gum is completely biodegradable and soluble in both hot and cold water. Like many other hydrocolloids, if not heated when dispersed in water, it will form lumps [27]. Unlike many other water-soluble polysaccharides, xanthan gum’s thermal stability against hydrolysis is better. The reason behind this is the oriented helical structure of xanthan gum which protects the molecule from depolymerization [28]. The oriented conformation of the gum is stabilized by salts, and the presence of salts is essential for the ideal functionality of the gum. Xanthan is stable over a wide range of pH, which makes the gum suitable for food applications (like fruit systems and salad dressings) as well as for cleaners. The gum is shielded against enzymatic attack due to the barrier function of the trisaccharide side chains. Fungal celluloses can cleave the main chain of the gum when it is in disoriented form; however, it is stable when the gum is in oriented form. This makes the gum a very versatile stabilizer and thickener even in high enzymatic loads. Xanthan gum is nondigestible and is known to lower the calorific value of foods [29]. Xanthan gum’s calorific value is 0.6 kcal/gm. The rheological properties of gum are studied in detail. The aqueous solution of xanthan gum shows high intrinsic viscosity even at low concentrations. The apparent viscosity of the gum changes considerably at different shear stress; the higher the shear, the lower the viscosity. Xanthan gum shows non-Newtonian pseudoplastic behavior. The shear-thinning nature of the gum can be justified by the conformational status of the gum. The pseudoplastic behavior of the gum makes it ideal in food products (optimum mouth feel and flavor release) and confirms exceptional pourability, pumpability, and mixability. To get the desired flow behavior from xanthan gum, it is often combined with other hydrocolloids. Xanthan is known to have synergistic interactions with galactomannans like guar gum, locust bean gum, and glucomannans like konjacmannan. Optimum synergism between galactomannan and xanthan is obtained when heated above orienteddisoriented transition temperature of xanthan gum [30]. The synergism between locust bean-xanthan gum is of particular interest because of the formation of a firm and thermoreversible gel between the two nongelling components [31]. The viscosities of xanthan-guar mixtures are significantly higher as compared to the viscosities of the single solutions, although they do not gel.

6.3 Xanthan gum-based hydrogel in drug delivery applications 6.3.1 Oral controlled release drug delivery Xanthan gum, in combination with chitosan, has been explored as a potential anionic polymer to form polyelectrolyte complex matrix tablets for oral prolonged release of highly water-soluble drugs [32]. Shao et al. prepared matrix tablets loaded with sodium valproate and valproic acid (model highly water-soluble drugs) using chitosan and a combination of different anionic polymers (xanthan gum, sodium carboxymethyl cellulose (CMC), sodium alginate, and carrageenan) by wet granulation method [32]. At first, the authors prepared matrix tablets of single anionic polymers composed of xanthan gum, sodium CMC, and carrageenan and studied their in vitro drug release behavior. Sodium CMC and sodium alginate matrix system could control the drug release for 8e10 h only. However, the xanthan gum-based matrix system could control the release of drugs up to 14 h. The authors reported that drug release from a single polymeric-based matrix system is primarily dependent upon the swelling and erosion characteristics of the polymer used and also on the drug solubility. The poor gel-forming capacity and substantial erosion of the polymeric system coupled with high drug solubility were primarily responsible for fast drug release from a single polymeric-based matrix system [33,34]. In a quest to further prolong the release of the drug by 24 h, the authors prepared a polyelectrolyte complex matrix by combining chitosan with the anionic polymers xanthan gum, sodium CMC, carrageenan, and sodium alginate. Release from chitosan-sodium alginate system was the fastest (94% drug released in 6 h), followed by chitosan-sodium CMC/carrageenan system (complete release by 12 h), and slowest for chitosan-xanthan system (extended up to 24 h). Since chitosan-xanthan system exhibited the slowest drug release, such a system was selected for further investigation. Keeping the total weight of the matrix constant, the authors varied the ratio between chitosan and xanthan (1:1, 3:1, and 1:3). However, no significant change in the drug release pattern was observed. At chitosan:xanthan ratio of 1:1, the effect of the molecular weight of chitosan on drug release behavior was investigated by preparing a matrix with chitosan of

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different molecular weights of 50, 100, and 400 kDa. Though there was not much substantial difference in drug release pattern in the first 6 h of the dissolution study, drug release seemed to slow down after 6 h of dissolution study with the increase in molecular weight of chitosan used in the matrix. The authors opined that the gel strength of the matrix is directly related to the molecular weight of the polymer used. An increase in the molecular weight of the polymer will increase the strength of the gel matrix, which in turn will decrease the erosion of the matrix [35,36]. This resulted in slower drug release from the matrix with the increase in polymer molecular weight. The observations were supported by the swelling and erosion behavior of the matrix. Erosion of the matrix (% weight of the matrix remaining) was 82% at 1 h, decreased to 68% at 6 h, and thereafter remained almost constant till 24 h. Similarly, the matrix exhibited initial rapid swelling (58%/h for the first 6 h), thereafter decreased to 17%/h for 6e24 h time period. Taken together, both the swelling and erosion of the matrix decreased after 6 h of study, which explains the in vitro drug release characteristics of the matrix. The investigators further opined that in situ polyelectrolyte complex formation was progressive and gradual and almost completed at 6 h, which resulted in a decrease in swelling and erosion of the matrix, which in turn accounted for the decrease in drug release rate after 6 h of dissolution study [37]. Kundu et al. prepared aluminum ions crosslinked interpenetrating network hydrogel beads containing sodium CMC and sodium carboxymethyl xanthan for sustained delivery of aceclofenac, a model antiinflammatory drug [5]. The hydrogels were investigated for their surface morphologies, swelling, and in vitro drug release studies. Scanning electron microscopic study of the hydrogels indicated that sodium carboxymethyl xanthan beads were spherical and discrete and the surface was smooth with few depressions. As sodium carboxymethyl xanthan was replaced with increasing amounts of sodium CMC, the hydrogels became elongated; the surface became dense and folded and depressions disappeared. The authors attributed such changes in surface morphologies of the hydrogels were due to the viscosity of the polymeric solutions and the extent of ionic crosslinking. Low viscosity of sodium carboxymethyl xanthan (113.6cp; 1% w/v solution) solutions provided sufficient fluidity to the polymeric solution to attain a spherical shape upon extrusion from the needle. Low viscosity coupled with a lesser degree of crosslinking (degree of O-carboxymethyl substitution 0.7) with aluminum ions produced depressions on the surface of the beads upon drying. As sodium carboxymethyl xanthan was replaced with increasing amounts of sodium CMC, the viscosity of the polymeric solutions increased (283.47cp; 1% w/v sodium CMC solution), which made the beads look elongated when extruded from the needle. High viscosity coupled with greater degree of crosslinking (degree of O-carboxymethyl substitution 1.1) made the surface denser, folded, and devoid of any depression. The swelling study indicated that the degree of swelling was less in acid media pH 1.2 when compared with swelling in phosphate buffer pH 6.8, the swelling of sodium carboxymethyl xanthan beads were highest, decreased as sodium carboxymethyl xanthan was substituted with increasing amounts of sodium CMC, and lowest for sodium CMC beads. The authors attributed such differences in swelling behavior of hydrogels were due to the difference in the extent of crosslinking. As the degree of O-carboxymethyl substitution of sodium CMC was more than that of sodium carboxymethyl xanthan, the extent of crosslinking in sodium CMC beads was more than that of sodium carboxymethyl xanthan beads. As crosslinking increased, the mobility of the polymeric chains decreased, which decreased the swelling behavior [38]. Similar to the swelling behavior, drug release in acid media was less as compared to drug release in phosphate buffer pH 6.8, drug release from sodium carboxymethyl xanthan was highest, decreased as sodium carboxymethyl xanthan was substituted with increasing amounts of sodium CMC, and lowest for sodium CMC beads. The viscosity of the polymeric solutions, the extent of crosslinking, and differences in the swelling ratio of the hydrogels were mainly responsible for such in vitro drug release patterns from the hydrogels. Xanthan gum was also evaluated as matrix material in a triple-layered matrix tablet [39]. Gohel et al. prepared triplelayered matrix tablets of venlafaxine hydrochloride, where the drug was dispersed in the middle layer [39]. Drug-free barrier layers composed of xanthan gum were compressed on either side of the middle layers. The authors used xanthan gum both intra- and extragranularly. The layered tablet was prepared by a hybrid method of wet granulation and direct compression. Xanthan gum exhibited optimum compressibility characteristics as defined by its Carr’s index and Hausner ratio. The concept behind using xanthan gum intra- and extragranularly was to overcome the problem of stickiness of xanthan gum when granulated and also to modulate the drug release rate. The total amount of xanthan gum used in the matrix tablets was distributed among the layers and also between intragranular and extragranular compositions. Several batches were prepared by altering the compositions intragranularly and extragranularly and also between the layers. The optimized formulation which contained drug:xanthan in the ratio 1:1 and xanthan gum distributed in the core and barrier layers in the ratio 1:2.5 produced sustained release of the drug up to 24 h. The optimized formulation was also compared with a marketed commercial product, Effexor XR capsules. The optimized formulation produced slower drug release as compared to the drug release from the marketed product. Finally, the authors concluded that drug release was dependent on

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the amount of xanthan gum used in the middle and the barrier layers and xanthan gum can be used as a potential barrier material for sustained drug release. El-Gazayerly et al. prepared xanthan gum-based matrix tablets for the controlled release of pentoxifylline [40]. The tablets were prepared by direct compression method, and the effect of polymer concentration, rotational speed, and salt concentration on in vitro drug release rate was determined. It was observed that with the augmentation in the concentration of polymer, the release of pentoxifylline was slowed down. The results were verified by an increase in the mean dissolution time (MDT) of the tablet matrix with an increase in the polymer concentration. At a polymer concentration of 3.4%, the MDT was 2.36 h, which increased to 6.44 h at a polymer concentration of 29.3%. The authors opined that xanthan gum upon contact with the dissolution medium forms a thick and viscoelastic gel structure, which retards the influx of the dissolution medium into the core of the matrix and efflux of the dissolved drug from the matrix core. An increase in polymer concentration made the gel structure thicker and more viscoelastic which slowed down the penetration of dissolution media and diffusion of dissolved drug from the matrix. The rotational speed of the dissolution apparatus also had a significant effect on the drug release rate. An increase in rotational speed from 25 to 100 rpm increased the drug release rate. The above observation is justified because an increase in the stirring speed will decrease the thickness of the gel layer, which in turn will result in faster drug release. Ionic concentration in the matrix also had a significant effect on the drug release rate. An increase in salt concentration from 0.05 M sodium chloride to 0.1 M sodium chloride produced faster drug release. Higher drug release rate at higher salt concentration was due to enhanced rate of diffusion of the dissolved drug from the matrix. The release mechanism of the drug was Fickian diffusion as indicated by n values. Maity et al. prepared calcium-crosslinked carboxymethyl xanthan mini-matrices and investigated the swelling, erosion, and in vitro prednisolone release behavior from the mini-matrices [15]. The swelling of carboxymethyl xanthan minimatrices increased rapidly in the first hour, decreased considerably in the second hour, and then gradually decreased with time in both acid media pH 1.2 and buffer solution of pH 7.4. Incorporation of calcium chloride in the carboxymethyl xanthan mini-matrices decreased the swelling of the mini-matrices considerably when compared with carboxymethyl xanthan mini-matrices, but the pattern of swelling was the same as that of carboxymethyl xanthan mini-matrices. The carboxymethyl xanthan mini-matrices demonstrated initial burst erosion in both test media, following which the erosion progressively increased with time. The incorporation of calcium chloride in the carboxymethyl xanthan mini-matrices increased the extent of erosion of the mini-matrices. Prednisolone release from carboxymethyl xanthan mini-matrices was slow, liberating 15% drug in acid media pH 1.2 at 2 h, and 50% in buffer media pH 6.8 in the next 8 h. The authors attributed the slow release of the drug from the mini-matrices due to the formation of a thick viscous gel layer around the surface of the matrix and also due to the poor aqueous solubility of the drug. The incorporation of calcium chloride to the carboxymethyl xanthan mini-matrices altered the drug release behavior from the mini-matrices considerably. Calcium chloride concentration up to 40% w/w in the mini-matrices decreased the drug release rate from the mini-matrices. Further increase in calcium chloride concentration to 50% w/w in the mini-matrices increased the drug release rate from the minimatrices. Mini-matrix containing 40% w/w of calcium ions released 5%, 12%, and 20% of prednisolone in 2, 5, and 10 h, respectively. Calcium chloride, which liberates Ca2þ ions during the wet massing stage of tablet preparation and also during dissolution, binds with the free carboxylic groups of carboxymethyl xanthan gum and crosslinks the free carboxylic groups of carboxymethyl xanthan [41]. This crosslink hinders the movement of the polymer chains and results in the formation of a thick gel layer around the surface of the matrix [42]. An increase in calcium ions concentration increases the crosslinking density which in turn increases the gel strength of the matrix [43]. The influx of the dissolution media and efflux of the dissolved drug through the gel layer is decreased, which results in a decrease in drug release rate with an increase in the concentration of calcium ions in the mini-matrices. However, a further increase in calcium ions concentrations to 50% w/w in the mini-matrices increases the drug release rate from the mini-matrices as the excess unreacted calcium ions act as a channeling agent in the matrix. Since the drug release rate from the matrix was slow in the first 5 h of the dissolution study, the authors concluded that further investigation on calcium crosslinked carboxymethyl xanthan minimatrices would make such matrices suitable for colon-specific delivery of drugs. With this background, Maity et al. prepared compression-coated tablets of prednisolone for colonic drug delivery [44]. Prednisolone-loaded core tablets were prepared by direct compression of the drugs with excipients microcrystalline cellulose, crospovidone, magnesium stearate, and trisodium citrate. The core tablet was compression coated with a calcium crosslinked blend of carboxymethyl xanthan and sodium alginate in various ratios. The objective of the investigation was to achieve minimum drug release in the first 6 h of the dissolution study (long Tlag) and then immediate pulse release of the drug (short Trap). The optimized core tablet disintegrated within 30 s to expose the drug for absorption. The authors in their previous study (calcium crosslinked carboxymethyl xanthan mini-matrices) reported long Tlag (12% drug released after 5 h of dissolution study), however failed to achieve complete drug release (short Trap) even after 10 h. To achieve long Tlag and short Trap, carboxymethyl xanthan was replaced with sodium alginate. Substitution of carboxymethyl xanthan with various amounts of sodium alginate altered

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the drug release from the tablets. The drug release became faster as indicated by a decrease in both Tlag and Trap. The optimized formulation where the ratio between carboxymethyl xanthan and sodium alginate was 1.5:3.5, provided Tlag and Trap of 5 and 7 h, respectively. Further optimization of the tablets by incorporating osmogen in the core tablets and increasing the coat weight provided reasonably long Tlag and short Trap of 6 and 4 h, respectively. The authors concluded that optimization of the composition of the core tablets and coat layer would help achieve colon-specific delivery of the loaded drug.

6.3.2 Ophthalmic drug delivery Ophthalmic drug delivery is a challenging task for various reasons. Rapid elimination of instilled drugs due to increased reflex lachrymal secretion reduces the residence time of the drug in the ocular cavity, which in turn reduces the bioavailability of the drug [45,46]. In order to improve ocular bioavailability, frequent drug instillation is required, but is associated with several drawbacks [47]. The use of ointments, viscous solutions, gel, etc., solves the above problem, but they are associated with blurred vision (ointments) and patient noncompliance [48]. An ideal ophthalmic delivery system would be a liquid that can be easily administered, remain in the ocular cavity for a prolonged period of time, and control the drug release rate for extended time periods. Polymeric solutions which can undergo solegel phase transitions in response to external stimuli like temperature, pH, etc., are the most suitable for ophthalmic drug delivery [49e51]. Various polymers have been studied for such purposes. Poloxomer is one such polymer, which is widely investigated in ocular drug delivery. Poloxomer solution is associated with reverse thermal gelling properties, remains sol at lower temperatures, and begins to gel as temperature increases [52,53]. This makes them suitable for ophthalmic drug delivery. However, poloxomer solutions have low mucoadhesive properties and are effective at high concentrations [54]. Incorporation of a second polymer to poloxamer solutions improves the mucoadhesive property of the solution and enhances the residence time and thus drug bioavailability to the ocular region. In this context, poloxamer combined with carbopol, chitosan, and methyl cellulose has been investigated [55,56]. Bhowmik et al. prepared a novel in situ gelling ophthalmic delivery system of poloxamer and investigated the effect of xanthan gum on the gelling, residence, and drug release time of the poloxamer in situ gelling system [57]. In vitro gelation, thermogelling properties, gel dissolution, and drug release studies were conducted. In vivo residence time of the delivery system was investigated on albino rabbits. Poloxomer at a concentration of 20% w/v and room temperature (25 C) was a free-flowing liquid (sol state) and quickly gelled at physiological temperature (37 C). However, when xanthan gum was incorporated into poloxamer solutions, the solegel transition of poloxamer occurred at 16% w/v concentration of poloxamer. Thus incorporation of xanthan gum into poloxamer solutions makes poloxamer effective at a much lower concentration. It was also observed that the addition of xanthan gum also reduces the in situ gelling time of the poloxamer solutions. An increase in poloxomer concentration reduces gel dissolution time due to an increase in the viscosity of the polymeric solutions. The addition of xanthan gum to poloxamer solutions further decreases gel dissolution time due to the enhanced viscosity of the poloxamer-xanthan system. In vitro drug release study of the developed formulations was conducted and compared with a marketed formulation, Topin. While the marketed formulation released its complete drug load at 4 h, the developed formulation containing 18% w/v poloxamer released 88% of the loaded drug at the end of 8 h. The addition of xanthan gum at a concentration of 0.06% w/v to 18% w/v poloxomer solution further reduced drug release to 55% at 8 h. This observation was attributed to an increase in the viscosity of the polymeric solutions upon gelling at the physiological temperature of the ocular cavity. This study indicated xanthanpoloxomer system has better drug retaining characteristics as compared to the poloxomer system alone. In vivo resident time in respect of percentage increase in pupil diameter with respect to time profiles of the developed formulation (16% w/v poloxomer þ 0.225% w/v xanthan gum) was studied and compared with Topin. It was found that there was a 10-fold increase in pharmacological response and 4.5 times duration of response from developed formulation compared to Topin (Fig. 6.4). The above study indicated the usefulness of xanthan gum in poloxomer-based in situ gelling formulations.

6.3.3 Vaginal drug delivery The vaginal cavity has been utilized for the treatment of local infections and systemic absorption of drugs as well. Drugs that are administered to the vaginal cavity include prostaglandins, steroids, antibiotics, antifungal agents, microbicides acting as contraceptives and protection against sexually transmitted diseases (including AIDS), and drugs to maintain vaginal hygiene and to treat dry vagina conditions in postmenopausal conditions [58,59]. Vaginal dosage forms include suppositories, soft capsules, films, gels, ointments, etc. These dosage forms have poor patient compliance and lack aesthetic qualities due to discomfort caused by leakage and messiness of the product. All these factors lead to failure in attaining desired therapeutic response. A well-formulated vaginal delivery system should have high patient compliance,

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FIGURE 6.4 Mydriatic response of developed in situ gel formulation 16% MPS and Topin (n ¼ 3). Reprinted from Bhowmik M, Kumari P, Sarkar G, Bain MK, Bhowmick B, Rahaman Mollic MM, et al. Effect of xanthan gum and guar gum on in situ gelling ophthalmicdrug delivery system based on poloxamer-407. Int J Biol Macromol 2013;62:117e123, Copyright (2013), with permission from Elsevier.

great aesthetic qualities, should be able to deliver the drug optimally to achieve proper therapeutic effects, and above all should be retained in the vaginal cavity for an extended period of time to get therapeutic effects for longer time periods [60,61]. All these requisites can be achieved by bioadhesive drug delivery systems. Traditionally, natural polymers have been used for bioadhesive drug delivery systems. Among them, carbopol, polycarbophil, chitosan, sodium CMC, and sodium alginate have been extensively used for their bioadhesion and bioretention characteristics. Specifically for vaginal bioadhesion, carbopol and polycarbophil have been used [62,63]. In vitro methods to evaluate bioadhesion include measurement of tensile strength, shear stress, fluorescent probe method, adhesion weight method, etc. [64e67]. However, there are very few in vitro methods to evaluate vaginal bioadhesion [68]. Since natural polymers with vaginal bioadhesion properties are very few in number and in vitro methods to evaluate vaginal bioadhesion have not been explored properly, Vermani et al. embarked on a research activity to find the right kind of natural polymers with vaginal bioadhesion properties and also develop in vitro methods for the evaluation of their bioadhesion strength [69]. Accordingly, the authors prepared simulated vaginal fluid containing mucin mimicking the pH, viscosity, color, etc., of vaginal secretions. Model membranes have been prepared with cellophane to study bioadhesion mimicking the biological membranes. Sheep vaginal has been widely used as biological substrates for the measurement of bioadhesion in vaginal mucosa. The model membranes were soaked in the simulated vaginal fluids for complete hydration and thereafter used for bioadhesion measurements, and the results were compared with the bioadhesion from sheep vaginal membranes. Several polymers either alone or their combinations have been prepared in such concentrations that their viscosities were 30,000 cP  10%. Polymers studied included xanthan, sodium alginate, carbopol, polycarbophil, and chitosan. Assemblies for in vitro measurement of bioadhesion strength were prepared by the modification of buccal bioadhesion test assemblies. Of the various polymeric gels prepared, xanthan gum was found to possess the highest bioadhesion strength in model cellophane membranes in terms of tensile strength (160 N/m2) and shear strength (60 N/m2) followed by sodium alginate (tensile strength 120 N/m2; shear strength 60 N/m2). Similarly, in sheep vaginal mucosa, xanthan gum was found to possess greater and statistically significant bioadhesion strength in terms of shear strength (80 N/m2), followed by sodium alginate shear strength 60 N/m2. However, when bioadhesion strength was measured in terms of tensile strength applied to separate the polymeric gels from the sheep vaginal mucosa, it was found that sodium alginate possessed higher bioadhesion strength (tensile strength 120 N/ m2; followed by xanthan gum tensile strength 110 N/m2). In the case of a combination of polymeric gels, the bioadhesion strength of xanthan-alginate gels was found to be greater than xanthan-polycarbophil polymeric gels. However, an interesting observation was that the bioadhesion strength of individual polymeric gels was greater than the combination polymer gels. Retention of gels in the vaginal mucosa was higher for xanthan and alginate gels. Based on the above results, the authors concluded that xanthan gum can be a potential natural polymer bioadhesion-based vaginal drug delivery.

6.4 Xanthan gum-based hydrogel in regenerative medicine Tissue regeneration is the method of regenerating injured tissue or reconstructing organs by creating a suitable threedimensional (3D) milieu for optimal cell attachment and growth, followed by the development of new tissues [17]. Such a suggested 3D milieu may be produced by creating an artificial biomimetic structure that is physiologically similar to the natural extracellular matrix (ECM) as a supporting framework [70,71]. Based on the tissue regeneration application,

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several natural polysaccharides are being widely utilized to create artificial biomimetic assemblies such as hydrogels, microparticles, microbeads, foams, microspheres, and nanofibrous films. Natural polysaccharides are abundant, simple to produce, and biocompatible, and they imitate the biological ECM of tissues. XG has recently received much interest in the production of tissue scaffolds for tissue regeneration. XG-based tissue grafts are widely recognized, and it has a 3D milieu with appropriate structural configuration and nanotopology and possesses suitable strength for the successful regeneration of injured tissues [72]. On interaction with a certain bivalent cation, XG produces a weak substance whose consistency is like gel due to its double helical structure. Other biomaterials or reinforcements can be used to control the mechanical stability of XG and various other features. XG offers remarkable biomimicking potential in tissue regeneration applications such as bone scaffolds, cartilage, and skin regeneration [73]. Fabrication and characterization of hydrogels made up of XG are done for tissue regeneration applications and cellular absorption experiments [74e76]. Materials for biological purposes should have physical and chemical qualities that allow cells to adhere, multiply, and differentiate [77]. Scaffolds should replicate the elastic moduli of soft neural to rigid nonmineralized bone tissues, which are 0.1 and 40 kPa, correspondingly [78,79]. Hydrogels which are made up of XG are biocompatible and biodegradable; therefore, they are considered intriguing materials to study as tissue scaffolds [80].

6.4.1 Xanthan gum hydrogel for bone tissue regeneration Bone abnormalities have reached the peak of the list of reasons for sickness and impairment among the aged worldwide [81]. Even though autografting is regarded as the gold standard for bone defect restoration, it is limited by donor-site sickness and unidentified adverse effects. As a result, regenerative medicine has caught the attention of researchers as a potential strategy for repairing bone lesions without the limitations and constraints of xenografts, bone allografts, or autografts [82]. Injectable hydrogels have an added benefit over other biomaterials since they can conform to the edges of the defect and can be introduced into deep lesion areas with minimal invasion. As a result, injectable hydrogels have the potential to greatly reduce operation time, scar formation, postoperative pain (leading to reduced muscles being impacted during treatment), and recovery period [83]. Furthermore, several hydrogels can mimic the original ECM, allowing cells to grow and mature in an optimal environment [84]. Integrating growth factors, as well as other osteoinductive elements, can aid to enhance therapy outcomes [85]. Although XG-based injected hydrogels have indeed been effectively used for bioink, biomolecule delivery systems, and tissue engineering of bone and cartilage, there is still little research being done on these or other biomedical applications. Dyondi et al. created tissue-engineered polymeric nanoparticles injectable hydrogel for the administration of various growth for bone tissue regeneration [86]. A double growth factor-loaded injectable hydrogel technique for osteoblast formation in bone tissue regeneration was devised and evaluated. The growth factors used were basic fibroblast growth factor (bFGF) and bone morphogenetic protein 7 (BMP7), and these were incorporated into a polymeric blend and the whole setup was used as an injectable matrix for bone repair. The hydrogels possessed a highly permeable network model that offered an optimal ecosystem for osteoblast growth. In several tissue engineering applications, it was observed that gels prepared using gellan and xanthan blend proved to be a potential transporter of growth factor. Chitosan-based nanoparticles were made and mixed with gels made up of a gellan and xanthan blend which led to the formation of in situ gelling scaffolds. Gellan and xantham gum exhibited viscoelasticity, whereas chitosan showed potent antibacterial activity. A double growth factor delivering method was utilized to induce the development of human lethal osteoblasts using the in situ gelling systems made up of gellan and xanthan gum having chitosan nanoparticles. They had a particle size of 297  61 nm which was ideal to deliver BMP7 and bFGF. The nanoparticle-loaded hydrogel displayed muchimproved cell differentiation and prolonged cell growth due to the release of growth factors for a longer time. In the prepared in situ gelling systems, BMP7 helped the formation of human embryonic osteoblasts as it is an osteogenic factor. In the intro, drug release study of the growth factors without incorporation in nanoparticles, it was seen that on Day 17, there occurred an 80% liberation of BMP7 loaded in hydrogels without a major immediate release. Whereas, BMP7 enclosed in the nanoparticles within the hydrogels demonstrated a release of 27% by Day 17 following an initial release of around 10%. In the case of bFGF, the cumulative release of bFGF from chitosan nanoparticles had approached 80% by Day 20. The release of bFGF encased inside hydrogel-entrapped chitosan nanoparticles had reduced even more by Day 21, with a total percentage release of around 49%. The growth factors were entrapped into the nanoparticles by the process of soaking. This process was responsible for the deterioration of the nanoparticle’s structure and the immediate release of growth factors. According to the observations, the stability of growth factors within the in situ gelling system seemed to be promising for bone tissue repair. The prepared gels had also proven their potency against the pathogens involved in implant failures like Staphylococcus aureus, Pseudomonas aeruginosa, and Staphylococcus epidermidis. A noninvasive injectable

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tissue scaffold for tissue growth can be provided by the in situ gelling system for the delivery of growth factors for a prolonged period of time. Bueno et al. synthesized hybrid biomaterials by electropolymerizing polypyrrole (PPy) (conductive polymer) with xanthan hydrogels that are obtained by crosslinking with citric acid (XCA) [87]. This resulted in the formation of a very hydrophobic hydrogel having great elasticity. Circular dichroism studies demonstrated that xanthan chains acquire a disorganized structure (coils) in swelled XCA hydrogels due to the carboxylate groups imparting negative charges. The hydrophilic nature of neat XCA promotes adherence to the tin-doped indium oxide (ITO) interface, whereas the positively charged nature of pyrrole tends to polymerize near the ITO interface as well as to the negatively charged XCA, which functions as PPy dopants. PPy’s planar nature encourages the development of porous stratified layers. The growth of fibroblasts and their adherence property on XCA and XCAPPy were examined for 21 days keeping consequently seven days gap in the presence of 0.4 T external magnetic field (EMF), as depicted in Fig. 6.5. After one day, cell attachment was assessed. Cell attachment on XCA had been very poor in the absence of EMF. Since XCA possesses a highly negatively charged density, this might induce repulsive forces to proteins involved in cell attachment. Although XCAPPy scaffolds are stiffer and more hydrophobic compared to XCA scaffolds; therefore, cells attached better to XCAPPy. Whereas in the presence of EMF, the proliferation and attachment were similar in both cases. The reason might be due to the interference of the EMF with the electrical field of XCA which leads to cell attachments. Although when EMF was absent, it was seen that cell proliferation upon XCA or XCAPPy hydrogels augmented significantly after seven days. However, after 14 days or more on XCA or XCAPPy, it did not rise substantially any longer, most likely because of confluence. ANOVA was used to compare the fibroblast proliferation onto pure XCA and electropolymerized hydrogels. Since XCAPPy hydrogels are harder and much more hydrophobic than XCA hydrogels, cell proliferation was always substantially higher (P < .05). After seven days under EMF, there was no noticeable difference between cell growth onto XCAPPy and neat XCA (P > .05). After 14 or 21 days of cell proliferation, the variations were noticeable. Thus, it was seen that the cell proliferation onto pure XCA was double in number, whereas in the case of XCAPPy, it was only 42%. This might be due to the magnetic stimulation which was more prominent in the case of scaffolds with greater charge density. These findings suggested that tissue scaffolds built of a mixture of natural and conductive polymers when combined with some magnetic field provide novel techniques for cell growth and attachment. Aguiar et al. developed films by blending two natural polysaccharides xanthan gum and chitosan in order to utilize their opposite charges to produce films with improved stability and was mineralized using hydroxyapatite [88]. In order to produce stable films, the layer-by-layer method was used as it enabled the study of interactions among each constituent. The dip-coating technique using gels having opposite charges increased the entanglement of polysaccharide chains, and pH alterations could affect the entanglement of xanthan and chitosan chains, producing films that are more resistant to varying pH. The films prepared were very much stable in all the buffer solutions specifically in the Tris-HCl buffer. The addition of two oppositely charged ions like calcium (Ca2þ) and phosphate (PO4)3 ions rendered more stability to the films, and the pH change caused them to precipitate as hydroxyapatite is present in bones which is a vital inorganic material. The controllable swelling behavior of the prepared films in a varied pH range, which is unusual for films manufactured with a single polysaccharide, or otherwise, could permit their application in physiological pH as well as locations with a pH drop,

FIGURE 6.5 Mean values and standard deviations were calculated for cell growth (MTT test) upon XCA and XCAPPy hybrid substance in the absence and presence of EMF. All values were deducted from the control measurements. Reprinted from Bueno VB, Takahashi SH, Catalani LH, Torresi SIC, Petri DFS. Biocompatible xanthan/polypyrrole scaffolds for tissue engineering. Mater Sci Eng C 2015;52:121e128, Copyright (2015), with permission from Elsevier.

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such as in the case of infected wounds. Understanding the polyelectrolyte characteristics of polysaccharides necessitates a knowledge of these interactions and the circumstances used to generate the films using the layer-by-layer technique. The prepared mineralized films of combined natural polysaccharides are extremely essential because they contain (Ca2þ) and (PO4)3 ions, mimicking bone tissue, and this can aid in the formation of fresh hydroxyapatite minerals. These ions must be available in order to regulate the metabolic pathway in bone tissue scaffolds. Furthermore, films made up of combining xanthan gum and chitosan containing these ions (Ca2þ and PO3 4 ) might be promising options for bone tissue engineering, as they permitted cell attachment and proliferation.

6.4.2 Xanthan gum hydrogel for cartilage tissue regeneration Osteoarthritis (OA), often known as a degenerative joint disorder, is a chronic condition that causes both structural and functional joint degeneration [89]. Articular cartilage is the specialized connective tissue that provides frictionless bone movement within the body’s synovial joints. Its finely organized ECM contains specific collagens and proteoglycans, which contribute to the tissue’s desirable viscoelastic and swelling characteristics. In contrast to the bone, articular cartilage lacks vascularity, lymph vessels, and nerves [90]. Articular cartilage has a restricted endogenous potential for regeneration owing to its avascularity and the minimal mitotic activity of its residing chondrocytes. Due to the lack of long-term clinical remedies, cartilage-related injury problems continue to be a major economic cost to society and a renowned source of global disability [91e93]. To overcome these deficiencies in cartilage defect treatment, recent efforts are centered on the application of tissue engineering concepts to produce better solutions at certain phases of skeletogenesis with the objective of developing functioning articular cartilage tissues [94e96]. Chondrocyte cells are present in the articular cartilage and have a vital role in the degradation process of ECM. Apoptosis, often known as automatic cell death, takes part in the pathophysiology of many disorders. Eventually, apoptosis of chondrocyte cells is responsible for the initiation and progression of various diseases related to cartilage tissue. The overexpression of inflammatory agents such as prostaglandin E2 (PGE2), nitric oxide (NO), tumor necrosis factor (TNF), and interleukin-1 (IL-1) in blood or synovial fluid promotes chondrocyte death. The inability of chondrocytes to proliferate after injuries or throughout the healing process affects cartilage regeneration. In the case of OA, cartilage tissue regeneration is crucial in overcoming these difficulties [97]. To repair articular cartilage abnormalities, autologous transplantation of articular chondrocytes utilizing tissue engineering has been widely researched. To allow cartilage tissue to rejuvenate, a cell-carrier material that closely resembles the normal environment of the cartilage-specific ECM must be produced. Artificial biodegradable polymers have recently been widely used to create 3D scaffolds for cell culture. Temporary scaffolds must be biocompatible and have mechanical characteristics that are extremely close to those found in the target tissue. The scaffolds must be exceptionally porous for seeding cells, repairing ECMs, and encouraging the admittance of supplied nutrients, waste elimination, and tissue in-growth [98]. Current OA therapy focuses on preserving joint mobility, alleviating clinical symptoms, and slowing cartilage deterioration [99]. Nowadays, treatments using intraarticular (IA) injection have gained much attention in the treatment of OA, as it provides greater joint mobility and leads to lower joint pain [100]. Han et al. prepared an IA injection using XG in order to discover its potential to repair joint cartilage and slow the development of OA induced by papain. It had a comparable rheological characteristic resembling HA [101]. Pure XG and IA injections made up of XG were tested on rabbit knees in order to evaluate their potential for the treatment of OA. They produced the rabbit OA model by injecting a solution via the IA route into the right knees. The solution was comprised 2% (w/v) papain and 0.03 mol/L L-cysteine which simultaneously triggered various degeneration alterations in the cartilage that closely resembled OA. The effectiveness of XG therapy in knee joint cartilage was determined by monitoring sulfated glycosaminoglycans (GAGs) and evaluating the breadth of rabbit right knee joints. The experiment lasted only a few minutes and was easily repeatable. Moreover, the rabbits survived the trial with no symptoms of a widespread toxic response. The prepared IA injection made up of XG possessed high stability against heat, had high transmittance properties, was free of endotoxin, and had lower dosing frequency. In the case of IA injection of HA, the dosing regimen was once every seven days for five consecutive weeks, whereas in the case of XG, it was once every 14 days for five weeks, but the chondroprotective impacts were similar in both cases. It was evident from macroscopic images (Fig. 6.6) that the exterior surface of articular cartilage tissue of the control group was intact, having a smooth and shiny light white-colored texture. In the case of the group treated with only saline, the surface was yellowish-white, irregular, and destroyed (Fig. 6.6B and F). Furthermore, the surfaces of the two groups receiving treatments appeared shiny with a white texture, and there was a mild injury in the femoral condylar tissue of the cartilage (Fig. 6.6C and D). Whereas the tissue of the tibial plateau cartilages had numerous fractures and flaws (Fig. 6.6G and H). It was observed the degree of breakdown of the cartilage tissue in the case of the tibial plateau was greater than in the femoral condyle. The reason might be due to the regular body posture of the rabbit.

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FIGURE 6.6 Comprehensive morphological analysis of the tibial plateau and femoral condyle. Femoral condyle: (A) control group; (B) saline group; (C) group treated with HA; (D) group treated with XG. Tibial plateau: (E) control group; (F) saltwater group; (G) group treated with HA; (H) group treated with XG. The surface of the cartilage was normal in (A). The cartilage’s surface was yellowish-white and severely degraded in (B). In (C) and (D), there were minor lesions on the cartilage surface. The surface of the cartilage was normal in (E). The cartilage surface was irregular, yellowish-white, and deteriorated in (F). (G) and (H) both had damaged cartilage surfaces; however, the severity was less severe than in (F). Reprinted from Han G, Wang G, Zhu X, Shao H, Liu F, Yang P, et al. Preparation of xanthan gum injection and its protective effect on articular cartilage in the development of osteoarthritis. Carbohydr Polym 2012;87:1837e1842, Copyright (2012), with permission from Elsevier.

From Fig. 6.7, it was observed from the graph that, inflammation took place which was explored by the increasing thickness of the knee joint. The inflammation was much more severe during the second week but reduced considerably thereafter. The sham group did not show any of these alterations. The swelling of the XG-treated group was much lower when compared to the saline group. However, there were no prominent differences between the XG and HA-treated groups. The dosing regimen of IA injection made up of XG reduced the severity of knee swelling, diminished cartilage deformations, reduced cell variations, and structural changes. In contrast to the saline group, researchers discovered that the XG-treated group had greater levels of sulfated GAGs. As a result, IA injection of XG might prevent ongoing cartilage damage and slow down the development of OA. This discovery has made a significant contribution to the establishment of a novel OA treatment strategy. Han et al. studied the impacts of xanthan gum on chondrocyte cell death and assessed them using immunohistochemistry and Western blot assay. The expression of metalloproteinase-1, 3 (MMPs) and tissue inhibitors of metalloproteinase-1 (TIMP-1) in the cartilage of an OA-induced rabbit by papain was also determined [102]. One of the

FIGURE 6.7 The widths of the knee joints in each group. When P < .05, comparisons are considered important. The width of the right knee joints was narrower in the XG-treated group than in the saline group, with an aP value of 60
C). The double-helical structures were formed when the solution was cooled. In the influence of cations, gellan gum has the capability to form gels much more quickly even at extremely low concentrations of gum. The occurrence of aggregation and double helix production at the transition temperature is impacted by the presence of cations and the concentration of polymers [37]. Other properties include nontoxicity, good temperature stability, mucoadhesiveness, biodegradability, biocompatibility, large water-holding capacity, the potential to pull back substantially upon mechanical disruption as well as the ejection of water by progressive compression, and resistance to acidic environments of the GIT.

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7.1.5 Gelling properties Both acetylated and deacetylated forms of gellans are water-soluble; however, HA gellans get entirely soluble in only hot water (at 85e95
C), whereas LA gellans are miscible in cold water [38]. Gellan gum molecules acquire coil shape at suitably high temperatures when it gets entirely dissolved in water. A thermally reversible transition reaction takes place; once this mixture is cooled, the random coil strands reform into highly complex double helices. These changes take place regardless of the presence of ions [39]. However, when cations are added to the solution, the electrostatic interaction has a considerable effect on the coilehelix transition. Temperatures, where these coil helix structural transitions occur, are also affected by the concentration of polymer and ion content [40]. The viscosity of the basic gellan solution increases in this ordered form, and at a subsequent high concentration, the polymer solution potentially acts as a “weak gel” (It could really hold particulates at zero shear and release them when the yield stress is surpassed). However, the addition of metal ions in the polymeric solution leads to the formation of “true hydrogel” [41]. If the generated double-helical structures are further cooled following the coilehelix transformation, they consolidate and form a 3D matrix by the formation of complexes with existing cations as well as hydrogen bonding with water. This is known as the solegel transformation, and the temperature required for this transformation can range around 30 and 50
C. Thus, the characteristics of the developed gel are affected by a variety of factors, the most crucial being the concentration and type of acyl groups present in it, the nature and amount of the cations added, pH, and the appearance of any water-soluble ingredients, and cooling constraints [42]. Gellan gum develops gels at smaller concentrations when gel-promoting cations are added in order to cool heated solutions. It can be found in two forms: substituted and unsubstituted types [43]. The degree of substitution affects gel characteristics, with the substituted type creating soft, elastic gels and the unsubstituted type producing hard, brittle gels. Due to their electrostatic interactions with the carboxylate groups present in the polymeric chains, gellan gum can produce gels in the presence of cations. Divalent cations enhance aggregation via site binding within pairs of carboxylate groups on nearby helixes, making gellan gum gelation more difficult than with monovalent cations. Ionotropic gelation requires monovalent or divalent cations, and the texture of gels may be altered by changing the type and concentration of ions [36]. Divalent ions have a greater impact on gel properties in comparison to monovalent ions. Among all the cations, calcium has the greatest impact on gel strength. The double helices of the gel generate a cation-induced junction zone, resulting in gel channels of larger strength.

7.2 Gellan gumebased hydrogel in drug delivery Gellan gum has wide applicability in the pharmaceutical industry in the form of release retarding and coating agents. It is also an FDA-approved stabilizer and thickener. The structure of gellan gum has free carboxylate groups which can be utilized to form ionotrophically crosslinked gels (hydrogels) in the presence of select metal cations like Al3þ and Ca2þ.

7.2.1 Oral drug delivery 7.2.1.1 Drug delivery to the oral cavity Periodontal diseases are a group of localized inflammatory diseases affecting the gum sulcus and are mainly due to the infestation of bacteria [44]. Periodontal diseases are primarily treated with antimicrobial agents. Systemic delivery of antimicrobials suffers from several side effects like nausea, vomiting, abdominal pain and also poor localized drug concentrations, and the emergence of bacterial resistance [45e47]. On the other hand, localized delivery seems to be promising as it provides an optimum concentration of drug at the infection site and reduction of side effects associated with systemic antimicrobial therapy [48]. The main challenge of localized drug delivery is the ability of the delivery system to penetrate the subgingival region and be retained there over a period of time. Classical agents for localized drug delivery to the periodontal region like dentifrices, chewing gums, irrigation solutions, and slow-release agents are associated with the above problems [49,50]. Polymeric smart gel formulations bypass all such problems and additionally provide sustained release of the loaded drug into the infection site for a prolonged period of time. These smart gel formulations undergo solto-gel transformations at the infection site which increases their bioavailability and enhances patient acceptability. Gellan gum, effective at low concentrations, has been utilized for the development of these smart gel formulations. Smart gels can be formulated under simple and mild conditions. It requires the preparation of dispersion of the gum at various concentrations in distilled water under constant stirring. Following this, the drug is added to the dispersion with continuous stirring to dissolve the drug. Dabhi et al. prepared a smart gel to be used in a periodontal drug delivery system for localized delivery of an antimicrobial agent, ornidazole [51]. Parameters for evaluation of the prepared gels include determination of pH of the formulation, in vitro gelling capacity, viscosity and rheological studies, in vitro drug release,

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and stability studies. The viscosity of the gellan gum solution is influenced by the pH of the medium [52]. Again, the aqueous solution of ornidazole is most stable in the pH range 5e6 [53]. The pH of aqueous dispersions of gellan gum was observed to be between 5.5e and 5.9, thus no pH adjustment agent was included in the formulation. In all the batches produced of varying concentrations of the gum, it was observed that sol to gel transformation was instantaneous. However, it was observed that the gelation strength and the integrity of the gel were markedly influenced by the concentration of the gum. An increase in the concentration of the gum increased the gelation strength, and the gels were stiffer in nature and eroded slowly. An increase in the gelation strength thus increased the residence time of the gels in the periodontal cavity. The viscosities of the solutions were also dependent on the concentrations of the gum. An increase in the gum concentration leads to an increase in the solution’s viscosity. The viscosity of the solution prepared with 1% w/v gellan gum was much higher compared to the viscosity of the solution prepared with 0.1% w/v gum. Rheological characterization of the polymer solutions indicated the shear-thinning pseudoplastic nature of the gum. In vitro release studies of ornidazole indicated initial burst release in the first 1 h of the study, after which there was polymer concentration-dependent release retardation during the rest of the study period. The initial burst release of the drug was attributable to hydration and the resulting water permeation to the gel matrix. An increase in polymer concentration leads to an increase in the viscosity of the gels and thereby resulting in a slower drug release. As the gel structure was more closely packed, it provided an increased barrier to drug release. Stability studies indicated that the formulations were stable as there was little effect on the pH, clarity, viscosity change, and drug content variation. The investigators concluded that gellan gum could be a potential candidate in smart gel formulations and drug delivery to the oral cavity.

7.2.1.2 Drug delivery to the stomach Helicobacter pylori is a gram-negative bacterium that resides on the human gastric mucosa and is known to produce serious gastric diseases like peptic ulcers, chronic gastritis, etc. [54]. Eradication protocols include treatment with antimicrobials (like clarithromycin) combined with an antisecretory agent [55]. Eradication is better if high drug concentration is attained and maintained in the gastric mucosa for an extended period of time, which can be achieved by increasing the residence time of the delivery device in the gastric mucosa. With this background, Rajinikanth and Mishra formulated gellan gumebased gastro floating hydrogel beads for localized delivery of clarithromycin, to eradicate H. pylori [56]. The beads were prepared by ionotropic gelation technique where calcium carbonate, dispersed in the gum matrix, was used as a gas-producing agent, and calcium chloride was used as the crosslinking agent. Aqueous dispersions of the gum (0.25e1% w/v) were prepared by sprinkling the gum onto deionized water. The drug and calcium carbonate were then dispersed in the gum solution with constant stirring. The resultant homogeneous dispersion was then passed through 21 gauge hypodermic needles and dropped into calcium chloride solutions to effect ionotropic gelation reactions. The hydrogel beads thus produced were kept for some time in the ionic solution (known as curing time) under continuous stirring for strengthening the beads and preventing aggregation. Morphology and particle size analysis of the beads, drug encapsulation efficiency, in vitro floating capacity, and drug release behavior are the important parameters evaluated by the authors. Growth inhibitions of the bacterium were also determined. SEM photomicrographs revealed a spherical-shaped bead with a rough outer surface. With an increase in gum content, the beads’ size increased. However, ionic concentration had no such effect on the bead size. The floating ability of the beads was dependent on the amount of carbon dioxide evolved, which in turn is dependent on the concentration of calcium carbonate in the formulations. The minimum concentration of calcium carbonate was found to be 0.5% w/v to affect the floating of the beads. An increase in the concentration of calcium carbonate reduced the floating lag time (time taken for the beads to come to the medium surface) and increased the floating time. A similar effect was observed with an increase in gellan concentration. High encapsulation efficiency (73%e89%) of the loaded drug was reported by the authors. The authors opined the low solubility of clarithromycin in calcium chloride solution is the reason for such high encapsulation efficiency. Similar encapsulation efficiency of other low solubility drugs was also reported by the authors [57]. In vitro release of clarithromycin from the beads was dependent on many factors like polymer concentration, calcium carbonate, calcium chloride concentration, and drug loading into the beads. An increase in polymer concentration increased the matrix density and also the diffusion coefficient of the drug in the beads, all of which accounted for slower drug release. An increase in calcium carbonate and calcium chloride concentration furnished more Ca2þ to the gellan matrix, thus increasing the internal ionotropic gelation and crosslinking density. This resulted in the retardation of clarithromycin release from the hydrogels. An increase in drug loading to the beads had a positive effect on the drug release pattern. The mechanism of drug release is mainly diffusion controlled. The in vitro growth inhibition of the bacterium was determined in terms of percentage growth inhibition by calculating the ratio of optimal density of the formulation mixture tubes against those of H. pylori tubes alone. It was observed that as the incubation time was increased, growth inhibition also progressively increased, and complete inhibition of growth was achieved after 12 h of incubation.

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7.2.1.3 Drug delivery into the intestine Tranexamic acid-loaded gellan microbeads were prepared for intestine-specific controlled release of the loaded drug [58]. The microbeads were prepared by ionotropic gelation technique using Al3þ ions as the crosslinking agent. Al3þ ions were chosen as the crosslinking agent as in comparison to divalent Ca2þ ions, the crosslinking rate was quicker [59]. The microbeads were evaluated for swelling, in vitro drug release studies, and in vivo pharmacokinetic studies. A swelling study of the prepared microbeads was conducted by measuring the amount of medium uptake by the beads. The authors observed that the results of the swelling study could provide insight onto the in vitro drug release pattern. The swelling of the beads was less at pH 1.2 and increased considerably at pH 7.4. It was because in the acid medium, the carboxylic acid group remains protonated and provided negligible electrostatic repulsive force, which resulted in the minimum swelling of the microbeads. However, in the medium of higher pH, the COOH was ionized to yield COO ions which crosslinks with the Al3þ and exert electrostatic repulsive force to the Hþ ions, resulting in elevated swelling. The in vitro drug release behavior of the drug was in complete agreement with that of the swelling study. As the polymer concentration was increased, the release became faster (Fig. 7.2). The authors explained that more polymer concentration would mean more swelling, and as swelling increased, the release also increased. The results of in vivo pharmacokinetic study are shown in Fig. 7.3. While Cmax and Tmax for oral solution were 70  1.9 ng/mL and 2  0.72 h, respectively, the same values for the prepared microbeads were 63  1.5 ng/mL and 4  0.87 h. The results were indicative of the slow release behavior of tranexamic acid from the microbeads.

7.2.1.4 Drug delivery to the colon Gellan gum has also been explored for colon-specific controlled release of therapeutic agents [60]. Resveratrol, found abundantly in grapes, peanuts, and other medicinal plants, has immense beneficial values to humans and is used to treat several illnesses like neurodegeneration, cardiovascular and metabolic diseases, and even in the cure of cancer [61e63]. However, this useful agent suffers from several drawbacks, the main being rapid and extensive metabolism in the gastric environment [64]. Targeting the release of resveratrol to the colonic region would bypass the problem. Accordingly, gellan gumebased mucoadhesive microsphere beads for colon-targeted controlled release of resveratrol were developed [60]. They evaluated the in vitro release pattern of the drug. It was seen that resveratrol release was significantly less from the mucoadhesive beads in the upper gastric environment and was sustained for about 48 h in the colonic region. Free resveratrol was dissolved to approx. 85% in 120 min and completely dissolved in 150 min, the same from beads were only 17% in 120 min and for complete dissolution it was around 48 h. This release pattern protected the useful therapeutic agent from metabolism in the gastric pH and improved its bioavailability. Gellan beads were also investigated for their safety in the human system. They used HT29-MTX and Caco-2 cell lines and colorimetric MTT assay to evaluate the cytotoxicity of gellan. After 24 h incubation in both cell lines, gellan beads didn’t decrease the cell viability. Subsequent incubation with gellan, cell viabilities of Caco-2, and HT29-MTX were found to be higher than 90.2% and 89.4%, respectively. The above findings indicated the nontoxicity and safety of the gellan beads. FIGURE 7.2 Release profile of TA in acidic and alkaline medium from TA-loaded GG microbeads prepared by variation in concentration of polymer at a constant AlCl3 concentration of 3% (w/v). (Mean  SD, n ¼ 3). [Reprinted from Bhattacharyaa SS, Banerjee S, Chowdhury P, Ghosh A, Hegde RR, Mondal R. Tranexamic acid loaded gellan gum-based polymeric microbeads for controlled release: in vitro and in vivo assessment. Colloids Surf B 2013;112:483e491, Copyright (2013), with permission from Elsevier].

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FIGURE 7.3 Plasma drug concentration versus time profile of TA in rabbits after single oral administration of TA solution and TA-loaded GG microbeads (Mean  SD, n ¼ 3). Reprinted from Bhattacharyaa SS, Banerjee S, Chowdhury P, Ghosh A, Hegde RR, Mondal R. Tranexamic acid loaded gellan gum-based polymeric microbeads for controlled release: in vitro and in vivo assessment. Colloids Surf B 2013;112:483e491, Copyright (2013), with permission from Elsevier.

7.2.2 Ocular drug delivery Drug delivery to the ocular region presents many opportunities and is also associated with many challenges. The physiological protective mechanisms of the eye are responsible for the rapid washout of the medication from the ocular cul-desac and thus account for the low absorption profiles of the drug [65]. To maintain a steady plasma concentration, frequent dosing is required and it adds to patient noncompliance and rising medication costs [66]. Since ocular medications are mainly in solution form, it aggravates the situation. Increasing the retention time of the formulation could be a potential solution to the problem. In situ gel systems have proved to be a potential ocular formulation option for scientists. The formulation is prepared in solution form, but when instilled into the ocular cul-de-sac, will readily undergo sol to gel transformation to increase the ocular residence time. These systems would thus offer the dual advantage of ease of administration and enhanced residence time [67]. Gellan-based ion-activated in situ gels have been found to increase the ocular residence time of the drug in animals and humans and are thus an area of research interest. Kotreka et al. prepared estradiol-loaded gellan-based in situ gels and evaluated the essential ocular formulation necessities like pH, clarity, osmolality, and rheological and drug release behavior [67]. The pH of the formulations (6.35e6.36) was very close to the physiological pH of tear fluid (7.4) which would be well tolerated by the eye. Additionally, the formulation was also isotonic to tear fluid, which would not be irritating to the eye upon installation. There were insignificant changes in the pH and osmolality of the solution upon storage, an indication of the stability of the gels. The in situ gels were shear thinning in nature, and there was a drop in apparent viscosity upon an increase in shear rate. This nature is advantageous and also desirable as it will provide easy handling and administration [68]. Upon instillation to the ocular cul-de-sac, the apparent viscosity will increase which will aid in retention of the formulation. The apparent viscosity of the formulation was found to be 12.5e23.2 cps which was dependent upon the storage condition. The apparent viscosity also didn’t change significantly during storage (P > .05), indicating stable formulation. The release of estradiol was 20% at 0.5 h, and the rest was released at 7e8 h. The above findings point out that gellan-based in situ gelling systems may possibly be a potential candidate for ocular drug delivery.

7.2.3 Nasal drug delivery The nasal route could be used to deliver drugs directly to the brain and offers several advantages like bypassing the bloodebrain barrier, ease of administration (noninvasiveness), and protection of the drug from the harsh environment of the GIT region [69,70]. The challenge is to enhance the residence time of the formulation in the nasal cavity because of rapid mucociliary clearance. As with ocular targeting, in situ nasal gels could bypass the problem of rapid mucociliary clearance. Hao et al. prepared in situ gellan gumebased resveratrol nanosuspensions nasal gels for brain drug targeting [71]. First, the nanosuspensions were produced by the antisolvent precipitation method. Resveratrol and surfactant were dissolved in ethanol to produce the organic phase. The stabilizer was dispersed in water to prepare the antisolvent stage.

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FIGURE 7.4 The mean sample concentrations of resveratrol versus time after i.v. and i.n. administration of resveratrol nanosuspensions and resveratrol suspensions-based DGG in situ gel, respectively (error shown as one standard deviation of the mean, n ¼ 5 per point). Reprinted from Hao J, Zhao J, Zhang S, Tong T, Zhuang Q, Jin K, et al. Fabrication of an ionic-sensitive in situ gel loaded with resveratrol nanosuspensions intended for direct noseto-brain delivery. Colloids Surf B 2016;147:376e386, Copyright (2016), with permission from Elsevier.

Then the organic stage was quickly added into the antisolvent phase with rapid stirring at 9000 rpm to produce the bluish opalescent nanosuspensions, which were then lyophilized for further use. Gellan gumebased resveratrol nanosuspensions were then prepared. Different concentrations of gellan gum dispersions were prepared and to which the lyophilized nanosuspensions were added with constant stirring to get the final product. The product was evaluated for isotonicity and gelling capacity and was found optimum. The rheological characterization revealed shear thinning pseudoplastic behavior of the gel. The gels were also evaluated for in vivo pharmacokinetic and brain-targeting efficiencies. The observations of pharmacokinetic studies showing plasma concentration time profile in the brain and plasma following intranasal administration of gellan gumebased resveratrol-loaded nanosuspensions and intravenous administration of resveratrol nanosuspensions are depicted in Fig. 7.4. As can be seen, the resveratrol plasma concentration in case of resveratrol IV injections, there is rapid drug elimination with very short biological half-life. Intranasal administration of gellan in situ gel showed significantly higher brain drug concentration compared to intravenous administration of resveratrol nanosuspension. The above findings indicated the potential of gellan gumebased formulation for drug-brain targeting via the intranasal route.

7.3 Gellan gumebased hydrogel in regenerative medicine Tissue engineering is becoming increasingly popular as a method to treat tissue and organ failure as well as their deterioration [72]. Tissue transplantation or implantation is required in many medical conditions where normal physiological parameters or homeostasis have been compromised. Part of the rising demand for such approaches might be attributed to the increased life expectancy due to medical and technological advancements. Donations of cells/tissues are commonly used in the treatment of patients with diseased or degraded tissue, although donor availability is much outstripped by patient demand [73]. The ideal tissue engineering paradigm involves introducing cells or tissue grafts that are local to the wounded area to aid in the healing process. Introducing tissue grafts to the damaged location for regeneration is the appropriate strategy in tissue engineering (Fig. 7.5). Hydrogels are ideal candidates for mediating cell transport and accommodating cells in their 3-D milieu due to their innate resemblance to the extra cellular matrix (ECM) [74]. Hydrogels are soft, porous structures that are made largely of water that mimic the original ECM and provide a perfect 3-D milieu for cell culture. Therefore, hydrogels have become a useful tool for studying the implications of ECM characteristics on cell activity [75e78]. Physicochemical changes of the gel-forming polymers and/or crosslinking chemicals can also be used to increase mechanical and biochemical capabilities, allowing them to better imitate native ECMs [79e86]. In situ-generated hydrogels have been extensively researched as cell transporters for in vivo tissue engineering and offer the additional advantage of injectability [87]. Natural or synthetic polymers can be used to make injectable hydrogels [88,89]. The use of natural hydrogels is advantageous over synthetic ones because they either contain ECM components or have an analogous molecular structure to glycosaminoglycans of natural origin [90]. Many of them possess cellular binding domains since they are obtained from natural sources, which

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FIGURE 7.5 Various strategies of cell delivery. Reprinted from Das M, Giri TK. Hydrogels based on gellan gum in cell delivery and drug delivery. J Drug Deliv Sci Technol 2020;56:101586, Copyright (2020), with permission from Elsevier.

permit cellular uptake and provide soluble signaling components toward becoming functional and effective in regulating cellular functions. Furthermore, natural hydrogels may be immunogenic inherently or as a result of the presence of contaminants such as proteins and endotoxins. Synthetic polymers may have a slow rate of deterioration in physiological circumstances, and their manufacture may include the usage of dangerous chemicals in some cases [91]. Gellan gumebased polymers are commonly utilized in the form of injectable hydrogels in cell transport, and a significant area has been focused on improving their properties, such as the addition of cell adhesion ligands or biodegradable moieties.

7.3.1 Gellan gum hydrogel for bone tissue regeneration Bone tissue regeneration is always a clinical problem in the field of oral, orthopedics, and maxillofacial surgery, despite the invention of many different bone augmentation techniques and bone transplant materials. Implementing bone graft materials, which come in two varieties: natural and synthetic bone grafts, is the present treatment standard [92]. Autografts, allografts (from human donors), and xenografts are examples of natural bone grafts (other species). Many surgeons regard autologous bone graft as the “gold standard” for bone regeneration and repair, owing to its absence of immunogenicity and superior biological action in terms of osteoconductivity, osteogenicity, and osteoinductivity [93]. Artificial bones are an appealing option that has the potential to overcome problems associated with allografts and autografts, for example, donor site toxicity, loss of bone-stimulating factors, postoperative discomfort, and erosion during regeneration [94]. Bone tissue engineering has been considered a leading approach, with the ability to transfer both osteoconductive and osteoinductive materials to the defect site (such as growth factors and cytokines) [95]. A novel method for bone regeneration is based on bone tissue engineering and exclusively designed implants with functionally graded properties, with the intent of developing a morphologically and physiologically implantation which is acceptable to the surrounding tissue [96,97]. Over the

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last few decades, a vast quantity of synthetic grafts has been developed to address the limits of natural grafting. Though, their biological effect in terms of induction and supporting bone formation is inferior to that of natural grafts. Adding cells and growth factors is one strategy to circumvent this and create more physiologically efficient systems capable of healing, maintaining, or boosting bone function [98]. Biodegradable polymers (such as hydrogels) used in bone tissue engineering provide benefits of flexibility in terms of controllable mechanical behavior and natural metabolic degradation inside the body. Aside from modifying polymeric structural features to promote cellematerial interactions, the polymer can be designed especially for a gel, and solidify in situ, leading to the development of bone regeneration treatments with minimal invasion. For regional effects, a variety of chemicals, such as growth factors, pharmacological compounds, and recombinant proteins, can be administered to the bone defect site [99,100]. Dyondi et al. designed a nanoparticulate dual growth factor-loaded injectable hydrogel system as a tissue-engineered substrate for the delivery of various growth factors namely bone morphogenetic protein 7 (BMP7) and basic fibroblast growth factor (bFGF) for bone regeneration [101]. A combination of two biopolymers was used namely gellan gum and xanthan gum in a ratio of 9:1 for the production of nanoparticulate in situ gel dosage forms for delivering cells and various growth factors at the defect area. Growth factors are rapidly eliminated from the injured site due to the shorter half-life; consequently, incorporating these growth factors inside nanosystems is preferable to parenteral administration. Growth factors also have relatively low tissue-penetrating ability, which may result in ineffective distribution at the site of injury. Nanocarriers made up of chitosan were synthesized and analyzed for the potential transport of growth factors. Since nanoparticles are tiny, they allow for easier diffusion and greater management of overdosage and biomolecule delivery time along with improving the in vivo effects of growth factors, this being essential for tissue regeneration. This nanocarrier was combined with gellan-xanthan gels to create an in situ gelling system that combines the viscoelastic behavior of biopolymers with chitosan’s strong antibacterial characteristics. The hydrogels had a porous network structure which provided an ideal microenvironment for the proliferation of the osteoblast. The prepared gellan-xanthan gels have been proven to be excellent growth factor transporters and other carriers in a variety of applications in tissue engineering. The nanoparticle-loaded gels showed significantly higher cell proliferation and differentiation because the growth factors were released over a longer period of time. To investigate the effects of single versus dual growth factors and free versus encapsulated growth factors, a differentiation marker test was carried out, examined, and compared. In comparison to single growth factor-loaded gels, dual growth factor-loaded gels revealed more calcium and alkaline phosphatase deposition. In the prepared gellan-xanthan hydrogels, BMP7 was employed as the osteogenic agent for the formation of human embryonic osteoblasts. On Day 17, hydrogels released BMP7 at an average rate of 80% without a significant initial burst. BMP7 included in hydrogels made up of nanoparticles showed a cumulative release of 27% by Day 17 following an initial release of about 10%. Encapsulation and stabilization of growth factors within the prepared nanoparticles and gels appear to be promising for bone regeneration, according to the findings. Gellan-xanthan gels exhibited antibacterial effects against Staphylococcus aureus, Pseudomonas aeruginosa, and Staphylococcus epidermidis which are the most prevalent pathogens in implant failure. The prepared nanoparticulate gellan-xanthan gel with chitosan nanoparticles offers a platform technology for cellular differentiation, because it is a noninvasive, injectable framework with optimum growth factor supply over weeks. Bellini et al. developed a three-component in situ gelling hydrogel for osteochondral defect repair, consisting of gellan gum (GG), HLA, and calcium chloride (Ca) [102]. These biocompatible polymers were combined in a brilliant way for the reparation associated with the osteochondral defect, resulting in the formation of an extremely stable system that prevents any material from leaking out. The polymeric substances employing various combinations of polymers/calcium ratios were tested for rheological and mechanical features, as well as the rate of degradation and adherence to bone defects. Gellan has the ability to produce gels via the following mechanism: the polymer is in its coil configuration at elevated temperatures; when the temperature drops, a thermally reversible phenomenon occurs. Conversion from the coil to double-helical structure eventuates, accompanied by the establishment (solegel transition) of a systematized structure made up of antiparallel double-helical strands. The junction zones are linked by untwined polysaccharide chains in the structure of expanded helical chains, resulting in the development of a 3D network and a hydrogel matrix. Monovalent or divalent cations can form a cluster around the helices to prevent electrostatic repulsion, which lowers the helices’ overall negative charge and strengthens the hydrogel (Fig. 7.6A). Furthermore, divalent cations may increase aggregation by forming site-specific bonds among pairs of carboxylate groups present on adjacent helices, resulting in ionotropic gelation. When strands of separate polymers, such as HLA, interact with GG strands during the gelling process, the G strands’ association can be disrupted, resulting in weaker junction zones, as shown schematically in Fig. 7.6B. The resulting hydrogels, HLAeCaeGG, exhibit mechanical and rheological properties distinctive from the CaeGG system. According to the approach shown in Fig. 7.7, these new features can be used to fill and permanently close minor bone defects.

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FIGURE 7.6 The development of the HA eCaeG hydrogel system is depicted schematically. Reprinted from Bellini D, Cencetti C, Meraner J, Stoppoloni D, D’Abusco AS, Matricardi P. An in-situ gelling system for bone regeneration of osteochondral defects. Eur Polym J 2015;72:642e650, Copyright (2015), with permission from Elsevier.

FIGURE 7.7 Schematic representation of the HAeCaeG hydrogel and its use as bone defect filler. Reprinted from Bellini D, Cencetti C, Meraner J, Stoppoloni D, D’Abusco AS, Matricardi P. An in-situ gelling system for bone regeneration of osteochondral defects. Eur Polym J 2015;72:642e650, Copyright (2015), with permission from Elsevier.

In vitro adhesion studies employing pig bones and human primary osteoblasts were also performed, demonstrating that this technique is suitable for bone tissue regeneration in osteochondral deformities. In vitro biocompatibility experiments with human primary osteoblasts revealed that this innovative system could be an excellent candidate for bone regeneration, potentially boosting surgical approaches in this area. Primary human osteoblasts were cultured in the specimen with excellent adherence to the swine bone regeneration: analysis revealed that the prepared hydrogel supported cell viability and osteoblastic advancement; cells, getting tightly entrapped inside the hydrogel matrix, osteoblast cells became translucent, a distinguishing trait of osteoblast cells, and started generating new tissue, as evidenced by the existence of mineralized growths. Various findings proposed that the semi-IPN HLAeCaeGG system produced could be an effective solution for the reparation of the osteochondral bone defect. Gantar et al. synthesized bioactive glass (BAG)-reinforced GG spongy-like hydrogels (GG-BAG) as unique hydrogels which could be used as scaffolds in bone tissue engineering [103]. The objective of establishing GG-BAG composites was to verify if there was a way to improve the mechanical characteristics and bioactivity of the GG by incorporating nanoparticulate BAG, extending the GG’s application range to include bone regeneration. As a result of the BAG particle reinforcement, both the microstructural and mechanical features of the material improved. These mechanical properties were

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observed to be composition-dependent and enhanced with the addition of BAG; yet, the study did not achieve the values required to accept biomechanical loading. When submerged in simulated body fluid, the composite material developed a tendency to form an apatite layer due to the inclusion of biologically active particles. In vitro tests were also performed using cultured human adipose-derived stem cells on the composite materials. Interestingly, stem cells derived from human adipose were capable of binding and proliferating while remaining alive within the gellan gum which was augmented with biologically active glass materials, which is a significant outcome when contemplating their usage in bone tissue regeneration. Despite being agglomerated within the GG matrix, the nanoparticulate, globular BAG particles considerably improved the mechanical properties and microstructure of the native GG hydrogels. The composition and quantity of BAG can be used to tailor mechanical properties. The composite structure including 2-component has Young’s modulus of 1.9 MPa (dry form) and 1.2 MPa (hydrogel form). These values were nevertheless significantly far below the tolerance level for load-bearing purposes, and they were lower than the largest achievable values for fused BAG having average porosity and BAGcontaining hydrophobic natural polymers. Careful handling led to a better distribution of nanoparticles inside the hydrogel matrices, which improved mechanical parameters even further. Furthermore, the inclusion of BAG particles into the GG hydrogel matrix ended up giving the developed system, the potential to mineralize in vitro, which could have been improved by further merging them with adipocytes, as the prepared structure assisted their adherence and distribution inside the structure without compromising the integrity of the structure. The prepared biologically active glass-reinforced gellan gum hydrogels showed tremendous potential for their usage in bone tissue regeneration.

7.3.2 Gellan gum hydrogel for skin tissue regeneration Surface injury or skin damage is one of the oldest complications in the operative field. Numerous skin substitutes and therapeutic approaches for skin wounds have been proposed in the previous 25 years. Skin tissue engineering is among the most superior skin therapy technologies for covering skin defects produced by medical procedures as well as other nonintentional injuries such as burns, soft tissue damage, painful and debilitating scars, or tumor removal [104,105]. Nanobiomaterials are increasingly being used in tissue regeneration because they resemble the arrangement of the ECM and provide a location for cell adhesion, differentiation, and growth [106,107]. GG has long been utilized for the healing process and tissue regeneration due to its biodegradability, biocompatibility, and high oxygen and water vapor permeability [108]. Pacelli et al. created and analyzed a novel nanocomposite (NC) network comprising gellan gum methacrylate (GG-MA) and laponite to investigate the impact of the nanoclay on the final features of the prepared biomaterials [109]. A biocompatible polymer, GG-MA, was mixed with laponite to produce a novel NC hydrogel matrix as a revolutionary wound-dressing substance. The concentration of clay was adjusted from 0.1% to 1% w/v to investigate probable claypolymer reaction before and after irradiation of ultraviolet rays, resulting in the development of poor gels with storage modulus (G0 ) ranging from 1 to 100 Pa. The NC hydrogel matrix containing 1% w/v laponite had the highest G0 value (above 1000 Pa), as expected. A similar system showed no changes in mechanical characteristics after steam sterilization when compared to a system containing clay of 0.5%, which showed a significant reduction in stress failure after steam management. The addition of laponite clay to the GG-MA network increased the viscosity of polymeric solutions. Furthermore, the addition of laponite clay allowed the formation of stronger hydrogels in comparison to the single matrix of GG-MA that could not be sterilized without deterioration. In contrast, based on the quantity of nanoclay added to the polymeric matrix, the introduction of laponite facilitated the production of stable hydrogels with nearly unaffected or unaltered mechanical behavior following heat treatments. This feature is crucial in the development of biomaterials that could be used as wound-dressing films in any other biomedical field that necessitates the use of a sterile environment. The incorporation of laponite nanoclay could also change the swelling behavior of the polymer system, which is severely impacted by the pH of the medium and the strength of ions present in it. The influence of laponite nanoclay on the swelling behavior of GG-MA hydrogels, as well as its extraordinary capability to interface with diverse types of molecules, presents an intriguing technique for regulating drug release by modifying the quantity of laponite added to the polymeric networks. Rheological investigations on the solution before and after gel production were used to explore the effect of clay/polymer molecular interactions. Prior to and after sterilization, frequency and strain sweeps were evaluated to measure the influence of the heat treatments on the mechanical characteristics of the prepared gels. It was also crucial to identify whether the introduction of laponite clay into the polymeric structure influenced the drug delivery potential of GG-MA by modulating and controlling swelling and release characteristics. As a result, the potential of the novel NC hydrogels to swell in various mediums was tested utilizing the swelling study and in vitro release study by taking a model drug called ofloxacin. Lastly, a neutral red analysis was conducted on a human fibroblast cell line (WI-38) to examine the in vitro bioactivity and physiological safety of NC hydrogels. It was proven to be biocompatible, as per the international regulations on biological analysis of cytotoxicity (ISO 10993-5) study. A human fibroblast cell line was treated with neutral red last (WI-38) to

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determine the in vitro biocompatibility and biological safety of NC hydrogels. Overall, these findings expanded the possibilities for using GG-MA as a beginning material for developing a smart sterilized device system that can be utilized as a transporter to regulate the release of drugs, as well as wound-dressing objects for the management of chronically infected burn wounds. Ismail et al. developed a unique gellan gumeincorporated titanium dioxide nanotubes (GG þ TiO2-NT) film for tissue regeneration that was successfully produced using the solvent casting process [110]. FTIR, XRD, and SEM were used to analyze the film’s physicochemical characteristics. The TEM micrograph exhibited hollow elongated structures, indicating that the nanotubes were successfully formed. Nanotubes were reported to have a diameter of 8e10 nm and a length of several hundred nm. A bundle of nanotubes was formed by connecting the nanotubes together. Both films had nearly identical diameters and thicknesses of 9 cm and 30 mm, respectively. At 600 nm, the prepared thin films were transparent in nature, with transmittances of roughly 98% and 95% for GG and GG þ TiO2-NT, respectively. The transparent films were ideal for wound dressing since the healing progression could be simply observed with the naked eye. The presence of interactions between the GG polymer matrix and TiO2 nanotubes has been demonstrated by FTIR investigations. The neat GG film had a smooth and uniform surface shape. When TiO2 nanotubes were introduced into the film, the surface became rough and less uniform. This was mostly due to TiO2 nanotube agglomeration after drying, as strands of extended nanostructured TiO2 were detected on the exterior layer of the GG þ TiO2-NT film, which could encourage cell proliferation. When compared to pure GG film, the young modulus, the tensile strength, and the robustness of the prepared GG þ TiO2-NT film were all improved. This is related to the TiO2 nanotubes’ reinforcing effects on biopolymers. The consistently distributed hard nanotubes in the GG matrix could explain the impacts of TiO2 nanotubes on GG film. The excellent mechanical stability of GG þ TiO2-NT films in soaking culture was due to their mechanical characteristics. However, after three days of testing for cell growth, the GG þ TiO2 film persisted in culture media. The fluorescent images of the cells at different time intervals showed that the cell propagation rate increased with time, especially for the GG þ TiO2-NT film. However, the cell proliferation for the GG film and the control trial were nearly identical, implying that pure GG was not significantly assisting cell proliferation. Cell growth for the GG þ TiO2-NT film, on the other hand, was exceptionally healthy, with the cells fully spreading after three days of incubation, demonstrating that the TiO2 nanotubes might speed up cell proliferation. Because of the higher surface density of extremely curved nanotube ends, nanotubes with thinner diameters of 30 nm had previously been demonstrated to enhance cell attachment and spreading. Furthermore, the TiO2 nanotube surface’s small diameter offered the optimal length scale for integrin clumping and focal contact generation, resulting in the maximum rate of cell growth, migration, and differentiation. The viability and proliferation of the cells further confirmed this. Until Day 3, the number of cells in the GG þ TiO2-NT film gradually increased. Meanwhile, after three days, the cell number of the GG film was practically identical to that of the control but significantly lower than that of the GG þ TiO2-NT film. These results showed that the prepared GG þ TiO2-NT film was friendly toward cell adhesion, nontoxic, and biocompatible. As a result, it could be utilized to treat superficial wounds, chronic ulcers, and second-degree injuries temporarily. They served as stimulators for cell growth and blood vessel formation as well as permanent dermis replacement. Cell proliferation experiments revealed that there was no toxicity and that the number of cells grew, making it a suitable candidate for skin tissue regeneration.

7.3.3 Gellan gum hydrogel for cartilage tissue regeneration Articular cartilage in adults has a relatively low reparative potential; hence, cartilage abnormalities do not generally regenerate in adulthood. The difficulty of activating the existing cells to produce new cartilage has hampered efforts to repair these defects. The implantation of stem cells and autologous chondrocytes has been researched in the desire of cartilage repair among the many diverse clinical techniques to stimulate the production of cartilaginous tissue [111,112]. The advancement of hydrogel technology is a primary emphasis in articular cartilage regeneration. These materials have desirable properties, such as an excellent level of hydration and diffusive capability, which influence chondrogenesis positively [113e115]. Due to its limited degree of self-repair, injured articular cartilage tissue typically results in a considerable decline in quality of life [116]. Tissue engineering has indeed been offered as a novel way to deal with issues like organ failure and tissue deterioration [117e119]. In the case of cartilage tissue regeneration, the cells are deposited on the hydrogel matrix and then transported to the defected area of the host, wherein novel functional tissue is produced and modified [120]. Hydrogels have received considerable attention as potential carriers for cartilage tissue regeneration since they can be shaped into products with identical 3D shapes and fundamental mechanical strength as biological extracellular matrices (ECMs) since they have the ability to retain large amounts of water. Shin et al. developed robust hydrogels that could encapsulate cells using the double network (DN) technique as loadbearing tissue scaffolds [121]. The DN hydrogels were prepared by photocrosslinking the hard and brittle first network

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FIGURE 7.8 Representation of the synthesis of (A) gellan gum methacrylate (GGMA) and (B) gelatin methacrylamide (GelMA). (C) A two-step photocrosslinking method for the development of DN hydrogels. Reprinted from Shin H, Olsen BD, Khademhosseini A. The mechanical properties and cytotoxicity of cell-laden double-network hydrogels based on photocrosslinkablegelatin and gellan gum biomacromolecules. Biomaterials 2012;33:3143e3152, Copyright (2012), with permission from Elsevier.

with GGMA and the soft and flexible second network with gelatin methacrylamide (GelMA). GG is a polysaccharide made up of hard repeating units with a six-membered ring formation and numerous hydroxyl (OH) groups that could be modified with photoreactive parts of methacrylates, the amount of which controls the crosslinking density of the final network. As a result, GG can be a good option for the rigid initial network of a DN hydrogel if it is heavily methacrylated. Gelatin, on the other hand, has a much more elastic chain, and the amine groups present in the remnants of the lysine or hydroxylysine are evenly distributed all through the polymer. As a result, methacrylated gelatin could be employed to create a second network, which is soft and ductile. A two-step crosslinking process was used to develop DN hydrogels (Fig. 7.8C). The first system, GGMA hydrogels, was developed by photocrosslinking the various concentrations of GGMA solutions. After that, the photocrosslinked GGMA hydrogels were further placed in GelMA solution, enabling the entry of the GelMA molecules into the prepared GGMA hydrogels. Following that, the gels were removed and subjected to light again for the second crosslinking process. The molecular weight of the substance was also taken into account. Therefore, to achieve considerably enhanced mechanical characteristics, the second component’s concentration must be significantly greater compared to the first component for the DN formation. The first polymer is required to be exceedingly high in molecular weight in order to develop rigid hydrogels at minimal polymer concentrations. Additionally, the rigidity of a hydrogel rises as the polymer’s molecular mass rises due to a rise in the number of crosslinked chains. Moreover, as soon as the polymer’s molecular weight increases, the effective number of crosslinked chains gets consequently increased, thereby increasing the rigidity of the hydrogel. Following all of these considerations, GG and gelatin were selected as the first and second polymers for this study. Methacrylation of both polymers was achieved by reacting them with methacrylic anhydride (Fig. 7.8A and B). Using 1H NMR spectroscopy, the degree of methacrylation (DM) of GGMA was found to be 24.5%, and that for gelatin ranged from 5.7% to 76.0%. Given that a gelatin molecule has a molar mass of almost 100 kDa, the average molar mass among

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two metharylate units in GelMA was estimated to be somewhere between 4000 and 50,000 g/mol. As a result, the primary and secondary components of DN hydrogels, extremely methacrylated gellan gum and methacrylated gelatin were effectively produced. The development of the DN was examined using diffusion experiments of bigger GelMA particles into the GGMA matrix, their resulting improvement in mechanical properties, and the variation in mechanical behavior between GGMA/GelMA single matrix and DNs. The compressive failure stress of the DN hydrogels produced was 6.9 MPa, which approached the strength of the cartilage. The mechanical force of DN hydrogels with a greater GelMA to GGMA mass ratio showed potential for the future development of even stronger DN hydrogels. The cell compatibility of the DN production process was demonstrated by encapsulation of NIH-3T3 fibroblasts in three dimensions (3D) and subsequent viability testing. The DN hydrogels generated from photocrosslinkable macromolecules could be advantageous for the regeneration of tissues that bear load due to their high strength and capacity to enclose cells. Tang et al. mixed oxidized gellan gum (O-GG) and carboxymethyl chitosan (CM-chitosan) to generate a dual hydrogel (C-GG) with considerably improved gelation temperature (Tgelation) and mechanical properties [122]. To reduce the gelation temperature of gellan gum, sodium periodate (NaIO4) was used to oxidize the neighboring dihydroxyl groups present in gellan gum. Furthermore, cells may be hazardous if there are too many aldehyde groups. Calcium ion (Ca2þ) crosslinking and the Schiff reaction were used to form the DN structure. The outcomes were very promising with respect to Tgelation temperature and mechanical strength. Following oxidation, the Tgelation temperature range was lowered from 42
C to 40 kDa). It has a wide range of qualities since the generating strain determines the molecular weight and branching of the polymer. Due to its solubility, viscosity, heat, and rheological attributes, it has excellent commercial interest in the food, pharmaceutical, and research industries. The low- and high-molecular-weight DEX can be synthesized by employing LAB. Studies have shown that the DEX can be obtained with specific chemical properties (branches and molecular weights) using specialized enzymes, which reveals their DEX-producing abilities and explores the possibility of using them to emulate a commercial strain. LAB obtained from unconventional sources is given special consideration (Leuconostoc mesenteroides NRRL B512) [11]. Lactic acid is produced by LAB as the primary or only byproduct of fermentation of carbohydrates. Acid-resistant gram (
ve), nonsporulating, catalase (
ve), cocci or rod-shaped, and having low guanidine cytosine content are typically immotile and commonly utilized. While, Lactococcus, Carnobacterium, and Enterococcus, many LAB species are nonpathogenic, have a widespread consensus, and are generally regarded as safe (GRAS) [12]. The dietary requirements are composite since they depend on elements like carbohydrates, amino acids, fatty acids, minerals, vitamins, and peptides frequently found in their natural environments [13]. Numerous bacteria and sugars exist in nature, and fermentation occurs in an anaerobic environment [11]. Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine. https://doi.org/10.1016/B978-0-323-95351-1.00016-8 Copyright © 2024 Elsevier Inc. All rights reserved.

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FIGURE 8.1 Chemical structure of DEX.

8.2.2 Structure of DEX DEX is made up of the leading chemical chain of D-glucose connected along a-(1/6) binds with potential branches of Dglucose connected through a link of a-(1/2), a-(1/3), and a-(1/4). Fig. 8.1 shows the backbone made up of D-glucose building blocks on a-(1/6) binds and various branchial arches such as a-(1/2), a-(1/3), and a-(1/4). DEX can be divided into two groups based on the molecular weight range of 40 kDae440 MDa and the distance of their chemical chain. Molecular weight 40 kDa is known as DEX, and less than 40 kDa is referred to as oligo-dextran, respectively [11]. DEX is available in several molecular forms; sucrose concentrations between 10% and 20% produce the maximum amounts of DEX. This is due to the inhibitory effect of sucrose, which reduces EPS generation. However, each glucan is complex and unique because of changes in the proportions, branch, and molecular weight that depend on the generating enzyme or fermentation, for example, yeast extract and milk permeate.

8.2.3 Source of DEX L. mesenteroides of the UICT/L18 strain, however, produced a forked DEX on a-(1/4) binds, contrary to this speciesdtypical propensity to have DEX on a primary chemical chain associated through a-(1/6) chemical bond and branches on a-(1/3) chemical bonds [13]. L. mesenteroides of the KIBGE-IB22 developed a DEX with (1/3) and (2/6) linkages. However, Leuconostoc, Lactobacillus, and Weissella (LAB) solely produce DEX that contains -(16) and -(13) bonds, with respective percentages ranging from 52% to 97% and 3% to 48%, respectively [14]. A solid culture medium comprising sucrose allowed DEX-producing bacteria to develop two distinct morphologies unrelated to the species or genus. After 72 h, the colonies of the Leuconostoc strains started to become dispersed. L. mesenteroides have subspecies like SF2, SF3, SD1, and SD23 of salmiana (agave), and they have 10% sucrose which shows linkages at a-(1/3) and a-(1/6) [12]. DEX-producing subspecies such as CUPV27 of Lactobacillus mali and CUPV 411 of Leuconostoc carnosum were separated from slimy ham and ropy apple must, respectively [12]. EPS synthesized subspecies BD1710 of L. mesenteroides from juice (tomato) with sucrose as culture medium bearing a relative molecular mass of 635 kDa, a one-dimensional backbone compiled from sequential a-(1/6) connected D-glucopyranosyl building block and around 6% a-(1/3) branches [15]. Chinese traditional fermented pickles were used to isolate Leuconostoc citreum SK24002, whose bioengineered d-glucan made up of distinct and composed of a-(1/3) and a-(1/6) connected D-glucopyranose building block on a relative molecular mass of 4.62 107 Da. A newly discovered LAB called Weissella cibaria 27 (W27) isolated significantly from dairy fermentation and has a relative mass of 1.2  107 Da [14]. W. cibaria subspecies such as 11GM-2 separated from the milk and W. confusa isolated from cassava fermentation produces DEX of both low and high molecular weight. Further, enzymes that can synthesize low molecular weight DEX directly or even in conjunction with other enzymes have been developed [16]. Most enzymes have been modified to synthesize large molecular weight DEX (up to 23 MDa). The type of enzyme (its source or acquisition method) determines the most of the characteristics of DEX produced by enzymatic means. Still, the molecular weights are strongly related to the substrate concentration [17]. The DEX obtained through enzymatic process showed slight variation at glucosidic linkage of -(1e6) and -(1e3) bonds, when compared to DEX made through fermentation. Designing enzymes to alter or transform DEX usually results in the production of DEX with the required characteristics [14].

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8.2.4 Properties of DEX The properties of DEX can vary depending on its molecular mass and branching. The main chain of it assumes a helical shape on a-(1/6) bonds at branches of a-(1/2), a-(1/3), or a-(1/4), which alters how the one-dimensional build of glucan is folded up. The molecular weight and branching have impact on its solubility and rheological characteristics. The interaction of the molecule with water via interactions with hydrogen bridges is referred to as polymer solubility because the hydroxy groups (-OH) as it is accessible to interact with water molecules if it were entirely linear (without branches) [18]. A DEX molecule has more amorphous regions, which facilitate water absorption and retention. Long-chain polysaccharides are less soluble than any low-molecular-weight polysaccharides [19]. There is no direct connection between the variations in qualities and the properties of the molecule [20]. DEX is categorized as soluble EPS because of its capacity to absorb significant amounts of water and produce hydrogels, regardless of their level of solubility. These characteristics are ascribed to eOH groups, which are known to join forces with other molecules through weak hydrogen bonds that can break under pressure [21]. DEX typically behaves Newtonian at low concentrations (regardless of shear rate) and non-Newtonian (or pseudoplastic) at high concentrations since its viscosity is inversely proportional to its shear rate and concentration [22]. Other research demonstrates that the viscosity and molecular weight are directly correlated, with the one increasing as the other does [23]. However, the temperature has influence on the polymer’s size, crystallinity, and intermolecular forces. On the other side, temperature determines how flexible polymers are. Because the polymer chains are immobile at low temperatures, linear amorphous polymers exhibit glass-like characteristics [17]. They often go through the stages of leathery (at the glass transition temperature, Tg), rubbery (Tm), and melting when the temperature increases (Tg). During this transition, polymers show their maximum degree of flexibility. Furthermore, the intermolecular interaction between the polymers causes rise in Tg of crystalline polymer chains [17].

8.3 Gelation Gelation is forming a gel-like rigid structure that acts as a flow resistance substance under pressure. At the beginning of 1982, Glicksman defined it as a 3D structure of an extended long polymeric chain-like network produced through physical or chemical cross-linking, which can entrap and immobilize liquid [24]. The polymers generally undergo a solidification process termed gelation [25]. After gelation, the polymer possesses infinite viscosity, that is, converted into the gel from the viscous polymeric solution. Phase separation does not occur during gel formation; it can occur in a homogenous solution made up of a solvent and a polymer. Several polymers tend gelation, such as cellulose acetate, poly(vinyl alcohol) (PVA), poly(vinyl chloride), and poly(phenylene oxide) [26]. In a recent study, Yagar and Balkan [27] described that laurel seed lipase was trapped in chitosan beads using an ionotropic gelation method that used tripolyphosphate as a multivalent covalent counterion. Gelation refers to forming a larger branched network-like structure of hydrogels by the cross-linker polymeric chains. The cross-linking of branched polymers to form a larger polymeric chain with reduced solubility is known as an infinite polymer gel or network-like structure. Gelation or “sol-gel transition” refers to conversion from finite branched polymer to infinite branched chain molecules. The critical point where the beginning stage of gel formation occurs is called the “gel point” [28]. Gelation time can be regulated by altering the solution’s pH and the concentration of the involved polymers. Both types of gelation, that is, physical and chemical gelation, involve the physical and chemical cross-linking of polymer chains [1].

8.4 Methods of preparations 8.4.1 Physical methods Hydrogel preparation by physically cross-linked polymers is an increased interest in the current research area. The different methods of physically cross-linked gel are summarized below.

8.4.1.1 Cross-linked hydrogel by hydrogen bonding A complex is formed in polyacrylic acid and polymethacrylic acid using polyethylene glycol (PEG). Herein, the physically prepared cross-linked hydrogel is produced by establishing hydrogen bonds between the carboxylic group of polyacrylic acid/polymethacrylic acid and the oxygen of PEG [29]. Mathur et al. described that hydrogen-bonded complex is also found in poly(methacrylic acid-g-ethylene glycol) and PEG [30]. Physically cross-linked hydrogel is produced by an association of hydrogen bonding which depends on pH-dependent swelling.

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8.4.1.2 Ionic interaction Electrostatic or ionic interaction can be found in various species, such as polypeptides, oppositely charged polyelectrolytes, and multivalent cations such as Ca2þ, Fe3þ, etc. Strong ionic interaction between oppositely charged ions can come close to forming a hydrogel. For example, to build a polypeptide chain of amino acids, positively charged species of amino acids such as negatively charged glutamic acid (E) and aspartic acid (D) associated with positively charged lysine (K) and arginine (R) [1]. Another example of ionic interactions is chitosan is cross-linked with glycerol-phosphate disodium salt to produce hydrogels. The chitosan solution remains liquid in this salt below room temperature and transforms into gel on heating. Such a gel can deliver the drug in protein-induced bone and cartilage formation [9,29,31]. Some polymers or polysaccharides undergo the gel formation with or without potassium ions or salt-free ions, which was synthesized by 1, 4linked-a-D- galactose and 1, 3-linked-b-D-galactose associated with an irregular part of sulfate groups. Stronger hydrogels can be produced when metallic ions are present [32]. DEX, in the presence of potassium ions, can undergo formation of most robust gel [9,31]. Thus, di- or trivalent counterions are utilized to form cross-linked ionic polymers. Some examples include chitosan-glycerol phosphate salt, and chitosan-polylysine, and chitosan-dextran hydrogels [33].

8.4.1.3 Freeze-thawing The freeze-thaw cycles are most commonly used to cross-link polymer to generate its hydrogel physically. The mechanism entails the structure forming microcrystals due to freezing and thawing. For instance, gel formation of PVA and xanthan using freeze-thawed method [33]. When compared to chemical cross-linking, freeze-thawing hydrogels are more biocompatible and biodegradable [9,32]. The formation of PVA cross-linked hydrogel is an example of a freeze-thawing cycle. The PVA has an ability to form intra- and intermolecular hydrogen bonds through freezing-thawing processes. PVA chains physically cross-link at the crystallite, created by hydrogen bonding, resulting in the production of PVA hydrogels [34].

8.4.1.4 Crystallization Crystallization is a method of hydrogel preparation involving the freezing-thawing process repeatedly. The first reported hydrogel preparation method was the preparation of PVA hydrogels using a freeze-thaw cycle. PVA-chitosan, PVA-starch, and PVA-gelatin hydrogels are some examples of hydrogel preparation by crystallization and have often been employed in tissue engineering [9]. DEX hydrogel and PVA hydrogel can also be prepared by crystallization. During the freeze-thaw cycle, the water content of the solution freezes and ejects the PVA chain and concentrates it. Crystallizations occur after forming hydrogen bond between PVA chains when it comes together. After the thawing procedure, the crystalline region remains unbroken, which may be called as physical cross-linking of PVA hydrogel [1]. The concentration of the polymeric solution, its molecular weight, the freezing cycle, and the freezing period all have a significant impact on hydrogel formation [32].

8.4.1.5 Heating-cooling a polymer A polymeric chain can be physically cross-linked by the cooling or heating process. Cooling hot gelatin or carrageenan solution leads to the formation of a gel. This phenomenon occurs because of the formation of helix-like structures and helix association due to the closure of polymeric chain and junction zone formation [35]. The physically cross-linked hydrogel can be produced due to the warming of polymeric solution, which results in the block copolymerization. Hoffman et al. described that the PEG-polylactic acid hydrogel was used to form block copolymer [31,33].

8.4.2 Chemical methods The chemical method is a popular hydrogel preparation which generally requires water during their preparation. The hydrogel prepared using chemical reactions possesses significant mechanical strength. The following methods of hydrogel preparation using chemical reaction are Schiff base reaction, addition reaction, condensation reaction, cross-linking by enzymes, and click reaction.

8.4.2.1 Schiff base reaction The preparation of hydrogel is performed by the Schiff base reaction method. In this reaction, imine bonds are formed when two polymers with amine and aldehyde groups come into contact. Imine bonds are created due to an interaction between two hydrophilic polymers with amine and aldehyde groups that have swelling properties. Hydrophilic polymers with hydrophilic moieties such as NH2, COOH, and OH are generally used for hydrogel synthesis [29]. In Schiff base reaction, cross-linking between hydrophilic polymer containing eOH groups and an aldehyde group may occur, for

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example, cross-linking of PVA using glutaraldehyde. In other instance, Schiff base reaction includes a cross-linking agent, for example, glutaraldehyde, PEG-dialdehyde, and vanillin. Glutaraldehyde is effective, readily available, and has high water solubility, most frequently used in cross-linking to produce hydrogel [36]. The cross-linking hydrogel is made from the aldehyde and chitosan’s amine groups.

8.4.2.2 Cross-linking by addition reactions Another hydrogel preparation method is an addition reaction using cross-linking of a functional group of participant hydrophilic. Akhtar et al. described examples of cross-linking addition reactions such as cross-linking of polysaccharides using divinyl sulfone or 1,6-hexane dibromide, 1,6-hexamethylene diisocyanate [29]. The Michael addition reaction is the addition of nucleophiles to conjugated (unsaturated) molecules such as unsaturated aldehydes/ketones, vinyl esters, or vinyl sulfones [37]. This reaction is highly selective and does not produce any toxic intermediate during the reaction. In addition reaction, a nucleophilic group such as amine and thiol of polymer reacts with an electrophilic group of methacrylate, acrylate, and vinyl sulfone [38].

8.4.2.3 Cross-linking by condensation reaction Condensation reactions are commonly used for the preparation of polyamides and polyesters. By utilizing these compounds, further syntheses of hydrogels can be possible among eNH2 with eCOOH or eOH group, respectively. Highly effective N, N-(3-dimethylaminopropyl)-N-ethyl carbodiimide (EDC) is typically utilized to prepare natural polymer-based hydrogel. Polysaccharide hydrogels are formed by Passerini and Ugi condensation processes [39,40]. This is a threeelement chemical condensation reaction involving isocyanide, carbonyl, and carboxylic acid. On condensation, the formation of a-(acyloxy)-amide occurs from these three components. Ugi condensation is a four-element chemical reaction in which protonated imine is formed from a condensation reaction between amine and carbonyl. After that, protonated imine, carboxylate, and isocyanide react together to produce a-(acylamine) amide [40].

8.4.2.4 Cross-linking using enzymes Hydrogel preparation based on PEG using an enzyme as an efficient and attractive method was described by Sperinde et al. [41]. In this cross-linking method, glutaminyl groups are associated with tetrahydroxy PEG (PEG-Qa). PEG network formation occurs after the transglutaminase is added to the blend of poly(lysine-co-phenylalanine) and PEG-Qa aqueous solutions. An amide bond is formed by the association of the g-carboxamide group of PEG-Qa and the ε-amine group of lysine [29,41].

8.4.2.5 Click chemistry Click chemistry is one of the methods of preparation of hydrogel, which have biomedical and pharmaceutical importance. Some typical reactions include click chemistry for forming in situ hydrogels, thiol-Michael addition, azide-alkyne cycloaddition, DielseAlder cyclo-addition, etc. In each reaction, the polymer chains must be joined to the proper functional groups to prevent the production of harmful byproducts during the synthesis of hydrogels [1]. Moreover, it has high-yield chemical products with fast reactions reported by Kolb et al. [42]. In 2001, Sharpless and coworkers prepared functionalized hydrogel cross-linker using copper as the catalyst [42]. This reaction is highly selective and occurs under mild conditions. The polymeric chain has plenty of room for incorporating different functional groups, leading to the development of responsive hydrogel matrix biomaterials. In this reaction, triazole-based cross-links were formed by cycloaddition of the azide and alkyne functionalized polymers catalyzed by copper [37]. In another example, Piluso et al. produced copper-catalyzed alkyne-functionalized hyaluronic acid (HA). In this reaction, a diazide-functionalized cross-linker was added to the alkyne-functionalized HA to cross-link it for the production of a hydrogel. Since this reaction is facile, the modification of mechanical properties of hydrogel could be possible by alkyne/azide ratio [43].

8.5 Biomedical application 8.5.1 Tissue engineering applications 8.5.1.1 Tissue engineering Tissue engineering is one of the experimental fields that found application of many biological and engineering principles. This principle explores in developing a functional replacement for the damaged tissue [44]. Tissue engineering technique

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needs cell signaling protein molecules for cell activity and self-gathered cells into functional tissue. In DEX-based hydrogels, highly hydrated cross-linked polymers are dependent upon the native tissue and used in 3D tissue engineering. After a week of therapy, DEX-based hydrogels appeared to hasten the requirement of endothelial cells to the area of the wound, facilitating quick neovascularization. DEX-based hydrogels greatly decreased intraabdominal adhesions without impairing wound healing [44e46].

8.5.1.2 Therapeutic vascularization Therapeutic vascularization is an opportunity but also presents a challenge for regenerative medicine. There are three broad methods through which vascularization of the tissue-engineered scaffold. The first step is the administration of regulatory constituents that will increase the growth of vasculature [47]. The second step is in vitro vascularization through cell engineering [48,49]. The last step is prevascularization in vivo before the transplantation of injured tissue. Dextranallyl isocyanate-ethylamine and PEG diacrylate were copolymerized in an 80/20 ratio to create the hydrogel. Compared to the control hydrogel and the injury covered just with dressing, the hydrogel when applied to a skin lesion increased blood flow to the burn wound area.

8.5.1.3 Tissue regeneration The proper regeneration of cells in the tissue needs a functional and within time vascularization. DEX hydrogel can easily give its function in wound-healing property. The hydrogel is a 3D-shaped structure that is like an extracellular matrix to the cell, when this penetrates within the cells leads to assist for tissue regeneration. Skim regeneration therapy is a common and ideal treatment for wound injuries [47].

8.5.1.4 Drug carrier The main aim of this field is to deliver the drug in specific and a reproducible manner. The specific pharmacokinetic effect of hydrogel is the key point for exploration of hydrogel in the drug delivery. Some biomedical applications of drug delivery system include: delivery of antimicrobial agent and proangiogenic growth factor. The drug can be loaded, while hydrogel preparation through cross-linking or hydrogel can be soaked into a solution containing a drug. The first method is more trustworthy since it ensures a known amount of drug, whereas the second one necessitates an indirect assessment of the drug that is encapsulated [31,50].

8.5.1.5 Advanced in vivo imagining There are various imagining techniques such as photodynamic therapy (PDT), magnetic resonance imaging (MRI), diagnostic magnetic resonance (DMR), and positron emission tomography (PET) imaging. This technique is used in diagnostic and mitigation purposes. This are some of the techniques which are used in advanced in vivo imagining technique [51].

8.5.1.6 Bleeding control and wound healing DEX-based bioadhesive hydrogels have specific biomedical applications such as wound-healing and bleeding control. For many wound injuries, skin regeneration therapy is the best treatment option. All types of wound injuries can have normal skin structures regenerated with the use of stem cells, tissue engineering frameworks, and the release of growth factors. This hydrogel has an ability to gather and form a fibrin bridge that can easily permit fibroblast migration and secrete collagen for tissue injury healing [52].

8.5.2 Drug delivery applications 8.5.2.1 Wound-healing properties The widest organ in the body system is the skin, and it is highly challenging to repair damaged skin in a whole. For this complex process, there is a contribution of many components in the body like matrix synthesis, cell lineages and various surrounding tissue, and the extracellular and intracellular signals. Neovascularization and angiogenesis were showed critical contribution in wound healing. The vascularization provides nutrients and oxygen with the removal of waste. In this way, DEX-based hydrogel provides wound-healing properties [53,54]. DEX-based hydrogel is recently employed in thirddegree burn wounds as educational scaffold for tissue formation and neovascularization. Skin regeneration treatment is the best course of action for many wound injuries. All types of wound injuries can have normal skin structures regenerated

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with the use of stem cells, tissue engineering frameworks, and the release of growth factors. Traditional treatments have less potential than stem cell therapy for facilitating tissue regeneration [47].

8.5.2.2 Antifungal properties DEX-based hydrogels can also stop or prevent fungal growth if there is addition of an antimicrobial agent within the hydrogel. In this condition, the DEX-based hydrogels act as a drug delivery vehicle. If certain drugs are loaded with the hydrogels, their antifungal activity is very potent, for example, Amphotericin. The antifungal substance did not cause hemolytic anemia in human blood and was biocompatible in vivo. Mice were implanted with amphogel and Candida albicans to prevent fungal infection and reduce the growth of fungal biofilms [55,56].

8.5.2.3 Antiadhesive property Amino dextran containing amino acid with an oxidized DEX consisting of aldehyde group from tissue adhesive. It shows applications such as ophthalmic procedures, tissue repair, and antiadhesive properties. Glaucoma causes an increase in intraocular pressure (IOP), and lowering this pressure has proven to be quite difficult. This issue may be solved by manufacturing soft contact lenses made of networks of polymers utilizing hydrogels. When used for glaucoma treatment, which mostly uses hydrophilic drugs, hydrogels’ highly hydrophilic polymer networks cause the drug to elute very fast. Soft contact lenses made of polymers of N, N-dimethyl acrylamide and methacrylic acid, which were modified appropriately, were able to deliver the hydrophilic medication timolol for roughly 24 h. This provided a path toward enabling sustained hydrophilic drug administration utilizing hydrogels [57,58].

8.5.2.4 Dextran-hemoglobin conjugates as a blood substitute People with hemoglobin and the asthmatic problem can be supplemented by this. DEX performs as a blood-human hemoglobin substitute in human. It didn’t provide damage to the kidneys and prevent hemoglobin excretion. It was first found that the amine-incorporating hydrogel disseminated the self-assembled tubules over the hydrogel network, whereas they accumulated in the purified PEG diacrylate hydrogel [58].

8.6 Conclusion This chapter systematically summarized the DEX, its chemistry, structure, and general attributes. DEX has good biodegradability, biocompatibility, and molecular flexibility for chemical modification owing to the presence of the hydroxyl group. Further, the application of DEX in the preparation of hydrogel has widened its scope owing to the chemical modification in the polymeric structure owing to hydroxyl group, which serves as an active site for modification. Further, biomedical and pharmaceutical applications, particularly in drug delivery, also increase its demand in medical sciences.

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Int J Artif Organs 2011;34:192e7. [44] Langer R, Vacanti JP. Tissue engineering. Science (80- ) 1993;260:920e6. [45] Sun G, Chu C-C. Synthesis, characterization of biodegradable dextraneallyl isocyanateeethylamine/polyethylene glycolediacrylate hydrogels and their in vitro release of albumin. Carbohydr Polym 2006;65:273e87. [46] Liu Y, Chan-Park MB. A biomimetic hydrogel based on methacrylated dextran-graft-lysine and gelatin for 3D smooth muscle cell culture. Biomaterials 2010;31:1158e70. [47] Cao Y, Hong A, Schulten H, et al. Update on therapeutic neovascularization. Cardiovasc Res 2005;65:639e48. [48] Tanaka Y, Sugihara M, Kawakami A, et al. In vivo analysis of supersaturation/precipitation/absorption behavior after oral administration of pioglitazone hydrochloride salt; determinant site of oral absorption. Eur J Pharm Sci 2017;106:431e8. [49] Mikos AG, Sarakinos G, Lyman MD, et al. Prevascularization of porous biodegradable polymers. Biotechnol Bioeng 1993;42:716e23.

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_ [50] Imren D, Gümüs¸derelioglu M, Güner A. In vitro release kinetics of bovine serum albumin from highly swellable dextran hydrogels. J Appl Polym Sci 2010;115:740e7. [51] Tassa C, Shaw SY, Weissleder R. Dextran-coated iron oxide nanoparticles: a versatile platform for targeted molecular imaging, molecular diagnostics, and therapy. Acc Chem Res 2011;44:842e52. [52] Heng MCY. Wound healing in adult skin: aiming for perfect regeneration. Int J Dermatol 2011;50:1058e66. [53] Weng L, Romanov A, Rooney J, et al. Non-cytotoxic, in situ gelable hydrogels composed of N-carboxyethyl chitosan and oxidized dextran. Biomaterials 2008;29:3905e13. [54] Draye J-P, Delaey B, Van de Voorde A, et al. In vitro and in vivo biocompatibility of dextran dialdehyde cross-linked gelatin hydrogel films. Biomaterials 1998;19:1677e87. [55] Wisplinghoff H, Bischoff T, Tallent SM, et al. Nosocomial bloodstream infections in US Hospitals: analysis of 24,179 cases from a prospective nationwide surveillance study. Clin Infect Dis 2004;39:309e17. [56] Ferreira L, Zumbuehl A. Non-leaching surfaces capable of killing microorganisms on contact. J Mater Chem 2009;19:7796. [57] Lauder CIW, Garcea G, Strickland A, et al. Use of a modified chitosanedextran gel to prevent peritoneal adhesions in a rat model. J Surg Res 2011;171:877e82. [58] Hoang TTT, Lee Y, Ryu SB, et al. Oxidized cyclodextrin-functionalized injectable gelatin hydrogels as a new platform for tissue-adhesive hydrophobic drug delivery. RSC Adv 2017;7:34053e62.

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Chapter 9

Hydrogels based on scleroglucan Tapan Kumar Giri Department of Pharmaceutical Technology, Jadavpur University, Kolkata, West Bengal, India

9.1 Introduction Polysaccharide is an extensively used natural biopolymer and has numerous advantages in comparison to synthetic polymers since it is obtained from natural sources [1e3]. It is widely used in drug delivery and regenerative medicine owing to its biocompatible, nontoxic, and biodegradable nature [4e6]. In addition, it contains a huge amount of amine (NH2), hydroxyl (OH), and carboxyl (COOH) groups, thereby providing functionalization. Polysaccharides are crosslinked through physical or chemical means to produce a three-dimensional hydrogel structure that absorbs large volumes of water or biological fluids [7e9]. The physical cross-link hydrogel is prepared by noncovalent interactions like hydrophobic interaction, hydrogen bonding, hosteguest interaction, and metal coordination. The chemical cross-link hydrogel is prepared by covalent interactions using radiation, thermal, redox, and photo-initiated free radical polymerization. Hydrogels are proficiently used in pharmaceutical and biomedical fields. It incorporated drugs and cells, which are protected from hostile environments. The different polysaccharides are used from various origins including seaweed (alginate), plants (guar gum), microorganisms (scleroglucan), and animals (chitosan) for the preparation of hydrogel. Scleroglucan is a polysaccharide obtained from fungi with numerous structural properties that are useful for the preparation of physical hydrogels [10]. It produces a triple helix conformation in an aqueous solution that can be converted into random coil chains in DMSO or at pH 12 [11,12]. Scleroglucan has numerous industrial applications, particularly in the oil industry for enhanced oil recovery due to its rheological properties and resistance to hydrolysis. It has other industrial applications such as adhesives, animal feed compositions, printing inks, and water colors. It is used as a carrier for the modified release dosage form [13e15]. Scleroglucan exhibited antimicrobial, anticancer, antiviral, and immune stimulatory activity [16e19]. Moreover, scleroglucan is chemically modified to introduce pH-sensitive groups. The carboxymethyl group is introduced into the polymer chain by reacting with chloroacetic acid in a basic medium. This carboxymethyl scleroglucan is capable of producing physical hydrogels without salt addition. The hydrogel was developed using a carboxymethyl derivative of scleroglucanentrapped antiinflammatory drugs which protect the gastrointestinal mucosa while modulating the drug release rate [20]. The source, structure, and physicochemical properties of scleroglucan have been discussed in this chapter. Additionally, the application of scleroglucan-based hydrogel for drug delivery and tissue engineering has also been highlighted.

9.2 Scleroglucan 9.2.1 Source Scleroglucan is a microbial glucan obtained from numerous filamentous fungi, particularly the Sclerotium genus. For industrial production, the Sclerotium genus as well as other genera are used. It is industrially produced from mainly Sclerotium glucanicum and Sclerotium rolfsii. Other industrial production sources of scleroglucans are Epicoccum nigrum, Schizophyllum commune, and Botrytis cinerea [21e24]. Exopolysaccharides produced by these fungi show analogous chemical structures but different molecular weights, branching frequencies, and side chain lengths [25]. The concept of scleroglucan was first explained by Halleck in 1967. In the early 1960s, the Pillsbury Company recognized the industrial productivity of scleroglucan. Scleroglucan exhibited outstanding rheological properties and was used in numerous industrial applications including the oil industry, food industry, cosmetics industry, and medical applications. Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine. https://doi.org/10.1016/B978-0-323-95351-1.00004-1 Copyright © 2024 Elsevier Inc. All rights reserved.

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FIGURE 9.1 Chemical structure of scleroglucan. Reprinted from Coviello T, Coluzzi G, Palleschi A, Grassi M, Santucci E, Alhaique F. Structural and rheological characterization of scleroglucan/borax hydrogel for drug delivery. Int J Biol Macromol 2003;32(3e5):83e92, Copyright (2003), with permission from Elsevier.

9.2.2 Composition and chemical structure Scleroglucan is a branched, nonionic glucan with a high molecular weight. It contains (1,3)-b-linked D-glucopyranosyl units of the main chain with linking of single (1,6)-b-linked D-glucopyranosyl units in every third unit (Fig. 9.1) [26]. Rinaudo and Vincendon first time established the four glucose repeating units scleroglucan by nuclear magnetic resonance and confirmed by many groups [27e29].

9.2.3 Physicochemical properties Scleroglucan dissolves in water at room temperature owing to the existence of 1,6-b-D-glucopyranosyl units [30,31]. The chains of scleroglucan in solution are present as a triple helical structure [32,33]. This helical structure is not aggregating due to the outside existence of D-glucosidic side groups in the structure. Moreover, macromolecules can be stabilized through interstrand hydrogen bonding [34,35]. Conversely, extreme viscosity changes are generally observed at higher concentrations of NaOH. At higher NaOH concentrations, the hydroxyl groups of polysaccharides ionize which destroys hydrogen bonds and consequently causes denaturation of the polysaccharide [36,37]. The variations in temperature are slightly affecting the viscosity of scleroglucan solutions. The scleroglucan solutions form thermoreversible gels at nearly 7
C owing to the cross-linking of the triple helix structure. It produces a stable gel in the existence of borax and salts of chromium at 10e11 pH values and precipitates out when quaternary ammonium salts are added into the solution at an alkaline pH [38,39]. It dissolves in DMSO or alkaline pH (greater than 12) solution where a random coiling state has been observed. It exhibited higher thermostability for 20 h at 120
C compared to xanthan gum owing to its nonionic nature. Moreover, extreme pH values and high concentrations of salt slightly affect its viscosity [40,41]. Scleroglucan solutions showed pseudoplastic character with a high yield value resulting in good solution pouring properties. Additionally, this characteristic reduces the risk of sedimentation [36]. Its molecular weight depends on the strains and cultivation media used during the commercial production of scleroglucan [42]. Scleroglucan administration via various routes in rats and dogs resulted in no blood abnormalities or toxicity. Skin and eye irritation were not detected in rabbits, pigs, or humans. Scleroglucan was used as a nondigestible dietary fiber and immune stimulant for humans [43e45]. It shows hypoglycemic, antiobesity, hypocholesterolemic, and antioxidant activity.

9.3 Drug delivery applications of scleroglucan 9.3.1 Oral sustained drug delivery Sustained drug delivery offers many advantages compared to conventional drug delivery systems, including enhanced efficiency, better patient compliance, and reduced toxicity [46e50]. Recently, hydrogel-based systems for controlled drug

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delivery have gained significant attention due to their resemblance to living tissue and their ability to retain large quantities of water [51e55]. Recently, biopolymeric hydrogels have received considerable interest in sustained drug delivery due to their biocompatibility, bioadhesivity, biodegradability, good mechanical properties, and thermal and chemical resistance. Among biopolymers, scleroglucan seems to be well suited for the controlled release of drugs [56e59]. A very low concentration of scleroglucan in solution (1% w/w) produces a gel which is a prerequisite for controlled release formulations. A hydrogel matrix of scleroglucan containing theophylline was prepared, and factors affecting the kinetics of drug release were studied [60]. The drug release kinetics follow non-Fickian behavior under different stirring speeds, drug concentrations, and temperatures. Cross-linking scleroglucan with 1-dicarboxylic acids containing four to eight methylene groups resulted in chemical hydrogels [61]. The synthesized polymer exhibited a fast penetration of water into the network system and was completed within 30 min. The weight of the gel, on the other hand, remains constant for 24 h. This indicates the greater compactness and stability of the network system. The water-absorption study was conducted in water, acidic, and basic solutions (Fig. 9.2). It was observed that the absorption increased with an increase in the methylene group from two to four (AeC) and decreased when it was five to six (DeE). The mesh size of the network was increased when the number of methylene groups was increased in the polymer network from two to four resulting in enhanced water absorption. A further increase in the methylene group from five to six decreased water absorption due to diminished total polymer polarity. It was observed that the swelling of polymers is minimal in acidic media. Several carboxylic groups are present in the polymer chain and are present in a unionized state resulting in low swelling. However, in neutral and basic media, carboxylic groups were ionized which created electrostatic repulsion and resulted in more swelling. Finally, theophylline-loaded tablets were prepared using synthesized polymers for controlled release of the drug. Dissolution of the drug from the tablet was studied in pH 7.4 buffer solutions, and tablet formulation C showed the slowest drug release. Release profiles of the drug indicate that hydrogel tablets could be suitable material for sustained release of a drug. Polysaccharides form a gel in the existence of borate ions [62e64], and similarly, scleroglucan reacts with borax to produce a three-dimensional structure [30,65]. The scleroglucan gel is produced through both physical and chemical linkages and is of a peculiar type. A semi-interpenetrating polymer network hydrogel based on scleroglucan, alginate, and

FIGURE 9.2 Water content (WC%) of the cross-linked polymers AeE as a function of n (number of the methylene groups of the cross-linker chain). Reprinted from Casadei MA, Pitarresi G, Benvenuti F, Giannuzzo M. Chemical gels of scleroglucan obtained by cross-linking with 1,u-dicarboxylic acids: synthesis and characterization. J Drug Del Sci Tech 2005;15:145e150, Copyright (2005), with permission from Elsevier.

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FIGURE 9.3 Release of MGB from tablets of Sclg (,) Alg (○), Sclg/Alg (D), Sclg/borax (-), Alg/ borax (C), and Sclg/Alg/borax (:) in SIF at 37
C. Open symbols refer to the sample without borax, whereas filled symbols refer to hydrogels, that is, polymer samples with the added borax. Reprinted from Matricardi P, Onorati I, Masci G, Coviello T, Alhaique F. Semi-IPN hydrogel based on scleroglucan and alginate: drug delivery behavior and mechanical characterisation. J Drug Del Sci Tech 2007;17:193e197, Copyright (2007), with permission from Elsevier.

borax was prepared to characterize mechanical properties and drug release behavior [66]. Theophylline and myoglobin (MGB) were used as model drugs. Release profiles of MGB from different formulations in SIF are represented in Fig. 9.3. The addition of Alg to the tablet formulations considerably influences the release of MGB from the tablet dosage form. The addition of Alg enhances the release rate owing to the hydrophilic nature of the polymer. MGB was completely released from the Alg tablet after 24 h, but an incomplete release occurred from the Sclg tablet. In addition, the release of MGB exhibited different behavior from that of Sclg/borax and Sclg/Alg/borax tablets. Sclg/Alg/borax tablets release the complete amount of the drug after 24 h, but Sclg/borax tablets did not release the complete amount of the drug. Alginates containing carboxylic groups accommodate a large quantity of water compared to Scgl resulting in extensive hydration and enhanced drug release. Pictures of different tablet formulations were taken after 24 h of dissolution studies. Sclg tablets had a compact, jelly like structure after 24 h of dissolution, but Sclg/Alg tablets were completely dissolved. Moreover, Alg and Alg/borax tablets showed complete dissolution. The Sclg/borax and Sclg/Alg/borax tablets showed different dissolution behaviors. The different colors of the two tablets were observed after 24 h of dissolution due to the different amounts of MGB still entrapped in the matrix (Fig. 9.4). It was observed that 40% of the drug was present after 24 h dissolution in Sclg/borax tablets (Fig. 9.4A), but only 10% of the drug was present in Sclg/Alg/borax (Fig. 9.4B). Similarly, releases of TPH from different formulations were carried out in SIF (Fig. 9.5). TPH is completely released from Alg/borax and Sclg/borax matrix tablets after 2 and 8 h, respectively. TPH release is faster compared to MGB release from all the formulations since TPH is a smaller molecule compared to MGB. Semi-IPN hydrogel was prepared using sodium alginate, scleroglucan, and borax for protein delivery into the gastrointestinal environment [67]. Finally, the tablets were prepared by compressing the hydrogel. The developed hydrogel tablet exhibited pH-dependent release behavior. Moreover, the mechanical properties of the tablet are strongly dependent on the pH value of the environment. Sodium alginate modulates the drug release rate from the matrix, and the rate increases

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FIGURE 9.4 Pictures of tablets, loaded with MGB, recovered at the end of dissolution test (24 h) carried out in SIF at 37
C: (A) Sclg/borax and (B) Sclg/Alg/borax. The different “transparency” of the samples is related to the different amount of MGB still present within the swollen tablet at the end of the release experiment. Reprinted from Matricardi P, Onorati I, Masci G, Coviello T, Alhaique F. Semi-IPN hydrogel based on scleroglucan and alginate: drug delivery behavior and mechanical characterisation. J Drug Del Sci Tech 2007;17:193e197, Copyright (2007), with permission from Elsevier. FIGURE 9.5 Release of TPH from tablets of Sclg/ Alg (D), Sclg/borax (-), Alg/borax (C), and Sclg/ Alg/borax (:) in SIF at 37
C. Open symbols refer to the sample prepared without borax, whereas filled symbols refer to polymer samples with added borax. Reprinted from Matricardi P, Onorati I, Masci G, Coviello T, Alhaique F. Semi-IPN hydrogel based on scleroglucan and alginate: drug delivery behavior and mechanical characterisation. J Drug Del Sci Tech 2007;17:193e197, Copyright (2007), with permission from Elsevier.

as the concentration of alginate increases. SA3 tablet (1% Sclg/borax and 3% Alg) disintegrated completely after 24 h in SIF and water resulting in the complete release of myoglobin. However, no MGB was released when the SA3 tablet was immersed in SGF. This may be due to poor alginate dissolution in acidic media and interactions between protonated proteins and the carboxyl group of alginates. The developed matrix tablet releases the protein (myoglobin) in the intestine and subsequently protecting it from the gastrointestinal tract (GIT). Novel hydrogels were prepared by cross-linking a scleroglucan derivative with 1, 6-hexanedibromide [68]. The absorption of water by the low cross-link hydrogel was affected by the ionic strength. Tablets prepared by using the cross-linked polymer acted as a monolithic system and were suitable for controlled drug delivery. Hydrogels were prepared using scleroglucan and borax [38]. The different drugs like myoglobin, theophylline, and vitamin B12 were loaded into the hydrogel which was then freeze-dried and finally compressed into a tablet. The three

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FIGURE 9.6 Release profiles of TPH, Vit. B12 and MGB from tablets in distilled water (pH ¼ 5.4) and 37
C. Reprinted from Coviello T, Coluzzi G, Palleschi A, Grassi M, Santucci E, Alhaique F. Structural and rheological characterization of scleroglucan/borax hydrogel for drug delivery. Int J Biol Macromol 2003;32(3e5):83e92, Copyright (2003), with permission from Elsevier.

different molecules were chosen since they have different dimensions. The release of drugs from a matrix tablet in water (pH ¼ 5.4) is represented in Fig. 9.6. The release of three molecules differs appreciably due to differences in molecular size. The release of theophylline (small molecule) is complete after 8 h. The complete release of vit B12 is observed after 24 h, and about 50% is released after 8 h. The release of myoglobin (molecular size more than the other two molecules) after 24 h is only about 45%. This result may be due to the molecules being trapped within the matrix and unable to diffuse through the meshes of the matrix. Swelling of the tablets was carried out in water for 24 h and is represented in Fig. 9.7A. The experimental data were obtained as a straight line indicating solvent entry into the matrix followed a Fickian process. The tablet absorbed a large amount of water (about 20 times) in 24 h. Moreover, matrix tablets continuously absorb water since after one day the equilibrium plateau is not reached. The dimensions of the tablet increase in the axial direction but remain unchanged in the radial direction which is uncommon. This unusual behavior may be due to the molecular structure of the scleroglucan/

FIGURE 9.7 (A) Water uptake data, relative to tablets of Sclg/borax, as a function of square root of time in distilled water (pH ¼ 5.4) at 37
C. (B) Relative increase of height, for tablets of Sclg/borax, during swelling experiments as a function of square root of time in distilled water (pH ¼ 5.4) at 37
C. Reprinted from Coviello T, Coluzzi G, Palleschi A, Grassi M, Santucci E, Alhaique F. Structural and rheological characterization of scleroglucan/borax hydrogel for drug delivery. Int J Biol Macromol 2003;32(3e5):83e92, Copyright (2003), with permission from Elsevier.

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borax combination and the compression effect which lead to an anisotropic system. Subsequently, the relative height of the tablet at different time intervals was represented in Fig. 9.7B, and a linear relation between tablet elongation and square root time was observed. Theophylline-loaded hydrogels were prepared using scleroglucan and borax [69]. The experimental result indicated that the developed hydrogel is suitable for sustained delivery of the drug. The effects of enhanced pH on the triple helix of scleroglucan and scleroglucan/borax hydrogels were studied [70]. Up to pH 13 and shear stress 9 Pa, the scleroglucan triple helix showed stability but breaks the viscosity at 9 Pa shear stress. In the existence of borax, up to pH 13 and shear stress 20e30 Pa system showed stability but breaks the viscosity at 20e30 Pa shear stress. At pH 14, viscosity dramatically decreased with both scleroglucan and scleroglucan/borax systems owing to the degradation of the triple helix structure and the breakdown of glycosidic linkages. The effect of pH on scleroglucan and scleroglucan/borax structure was examined through a spectrophotometer based on the interaction between the scleroglucan triple helix and Congo red. In the presence of borax, the maximum absorbance of Congo red is shifted as a function of pH. Scleroglucan was cross-linked with borax, aluminum, and iron to prepare a matrix for modified drug delivery [71]. The scleroglucan/borax matrix exhibited the highest strength compared to scleroglucan/aluminum and scleroglucan/iron systems. Tablets were prepared using these gels. A water-absorption study of these tablets was performed and revealed differences in their water uptake.

9.3.2 Colon-targeted drug delivery Drug delivery to the colon is intended to release the drug in the colon in response to the environment of the colon without causing early drug release in the upper GIT [72e76]. The microenvironment and physiological properties of the colon should be considered during the development of colon-targeted drug delivery. Generally, the GIT (stomach to colon) has dynamic changes in pH, motility, enzyme activity, and fluid contents [77e79]. In comparison to conventional drug delivery systems, colon-targeted drug delivery systems have high therapeutic efficacy and lower undesirable side effects [80,81]. To improve bioavailability, drugs that are unstable in the upper GIT due to acidic and/or enzymatic activity are developed in a colon-targeted drug delivery system [82e84]. pH-sensitive hydrogels of carboxymethyl scleroglucan were prepared for the oral delivery of nonsteroidal antiinflammatory drugs [20]. The physicochemical and biological properties of the hydrogel exhibited suitability for colon targeting. The developed hydrogel showed pH-responsive behavior and had a good affinity for aqueous media. The developed hydrogel releases the drug for a longer period of time at the site of action due to the mucoadhesive nature of the hydrogel. Ulcerogenic doses of diclofenac-loaded hydrogel administered to rats showed the absence of any gastric damage. This novel hydrogel could be used for the administration of drugs that damage the gastric mucosa or for colon targeting.

9.3.3 Topical drug delivery Topical dosage forms are commonly used to treat different disorders including pain, inflammations, and infections [85]. The skin is mainly attractive for the entry of drugs because it overcomes the problems associated with other administration routes [86]. However, the administration of drugs through this route is repeatedly hampered owing to its barrier properties. Hydrogel based topical dosage forms can efficiently deliver the drugs through this route [87]. Hydrogel-based drug delivery systems are a good candidate for topical administration owing to their soft consistency and high swelling ability. In addition, this dosage form provides an agreeable consistency and does not produce any greasy deposits on the skin in comparison to ointments and creams [88]. Carboxymethyl scleroglucan was synthesized for the development of hydrogels [10]. Physical hydrogels were formed due to ion-dipole and hydrogen bond interactions of the carboxymethyl group present in the synthesized scleroglucan (Fig. 9.8A). A hydrogel with a uniform consistency and suitable for topical application was produced using 2% w/v synthesized gum. The flow curve of hydrogel (2% w/v) without drugs showed shear thinning behavior (Fig. 9.8B). This shear thinning behavior is suitable for application as a topical formulation. Frequency sweep tests of the hydrogel were carried out in the viscoelastic region. The developed hydrogel exhibited a higher elastic modulus compared to its viscous modulus. Then, carboxymethyl scleroglucan hydrogels were prepared for the topical delivery of betamethasone, fluconazole, and diclofenac. Rheological and in vitro release studies were carried out to investigate the rheological properties of hydrogel and the dissolution of drugs. The developed hydrogel produces a negligible skin irritation index. The entrapment of these drugs did not affect the shear thinning behavior of the hydrogel system. Betamethasone-loaded hydrogel exhibited low mechanical strength. However, diclofenac-loaded hydrogel showed strong mechanical strength owing to the interaction between the diclofenac carboxyl group and carboxymethyl scleroglucan. Faster drug release was observed from hydrogel

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FIGURE 9.8 Rheological characterization of Scl-CM 300 physical hydrogels. (A) Schematic representation of hydrogen bond formation among carboxyl and hydroxyl groups. (B) Flow curve showing the pseudoplastic behavior of the polymer at 37.0  0.1
C. (C) Representative picture of SclCM300 physical hydrogel and the corresponding frequency sweep test carried out in the range of 0.01e10 Hz. Full circles represent the elastic modulus G0 and empty circles the viscous modulus G00 . Reprinted from Paolicelli P, Varani G, Pacelli S, Ogliani E, Nardoni M, Petralito S, et al. Design and characterization of a biocompatible physical hydrogel based on scleroglucan for topical drug delivery. Carbohydr Polym 2017;174:960e969, Copyright (2017), with permission from Elsevier.

loaded with betamethasone and fluconazole and followed the Fickian transport mechanism. However, slower drug release was observed from hydrogel loaded with diclofenac and followed a non-Fickian mechanism. An anionic derivative of scleroglucan (carboxymethyl scleroglucan) was synthesized for the development of topical drug delivery systems [14]. The mucoadhesive properties were studied through the interaction of scleroglucan and carboxymethyl scleroglucan with mucin by the turbidimetric method. The hyaluronic acid was used for comparison purposes. The transmittance value of the sample is low when the interaction between polymer and mucin is more owing to macroaggregates formation [89,90]. The mucoadhesive properties of scleroglucan showed similar properties to those of hyaluronic acid. However, carboxymethyl scleroglucan showed increased mucoadhesive properties. The carboxyl group of the polymer enhanced the hydrogen bonds responsible for interactions between mucin and the polymer. The synthesized polymer is able to form a hydrogel in the existence of calcium ions. Finally, the nonsteroidal antiinflammatory drugs were incorporated into the hydrogel for topical application. The release of the drug can be modulated by changing the concentration of calcium ions. The rheological properties and lack of skin irritation of the developed hydrogel make it suitable for topical application. Carboxymethyl scleroglucan was developed for topical delivery of acyclovir [91]. Hydrogels were prepared using calcium chloride. The mechanical strength of the hydrogel depends on salt and polymer concentration. Moreover, the hydrogel preparation exhibited thermosensitive behavior.

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9.4 Tissue engineering applications of scleroglucan hydrogel Tissue engineering is a useful method for the treatment of damaged tissue caused by trauma and diseases [92e96]. This approach has several advantages in comparison to conventional methods of treatment. It is a multidisciplinary field including material science, genetics, chemistry, engineering, medicine, and pharmacy [97,98]. Four important components of tissue engineering are (i) cells, (ii) scaffolds for cell function, adhesion, and transplantation, (iii) signaling molecules (growth factors and proteins), and (iv) bioreactors. The development of tissues or organs can be obtained through different approaches. One of the common approaches involves the isolation of the patient’s tissue and its harvesting in vitro. Then the cells are seeded into the extracellular matrix (scaffolds). The scaffolds containing cells are transplanted into the patients through injection or implantation in the desired site by surgery. Hydrogels have attracted more attention in tissue engineering for scaffold manufacturing due to their tunable physicochemical properties and similar in vivo environments [99e103]. Natural polymers are generally preferred for the preparation of hydrogel due to their biodegradable, biocompatible, and nontoxic nature [104e106]. Hydrogels are prepared by physical or chemical cross-linking methods to produce a three-dimensional structure. Cells are encapsulated into the hydrogel during the manufacturing of the scaffold and get the environment similar to the extracellular matrix. In situ and injectable hydrogels were prepared using derivatives of scleroglucan and dextran methacrylate [107]. Scleroglucan and its derivatives enhanced the mechanical properties of dextran methacrylate hydrogels. The prepared hydrogels are harder, more elastic, and easier to work. Hydrogels prepared with dextran methacrylate and carboxymethyl scleroglucan have a higher viscosity in comparison to polymers alone without salt addition. In the presence of biological fluids, all the developed hydrogels swelled and exhibited similar consistencies to natural tissue. Overall, hydrogels prepared from these polymers showed sufficient mechanical properties and were suitable for biomedical applications, particularly tissue engineering applications. Thermosensitive nanocomposite hydrogels were prepared using scleroglucan, chitosan, montmorillonite, and carboxymethyl cellulose [108]. The altered ratios of polysaccharides and the addition of montmorillonite change the gel-forming temperature of the hydrogel. The addition of montmorillonite significantly reduced the gel-forming temperature and enhanced the thermal stability of the hydrogel. Moreover, a significant decrease in swelling of the hydrogel was observed with an increase in the amount of montmorillonite in the formulation. The abovementioned properties of the thermosensitive are suitable for application in drug delivery, wound dressing, and tissue engineering.

9.5 Conclusion Scleroglucan has received a lot of interest since it was initially described in the early 1960s. Numerous studies on this polymer have revealed that it has a wide range of potential applications in pharmaceutics, both in its natural and modified forms. The scleroglucan hydrogels were created by using various cross-linking agents to modulate the drug release. The tissue engineering application of scleroglucan hydrogel is limited. The polymer has already demonstrated intriguing and, in some cases, very peculiar properties, indicating the wide potentiality of this polysaccharide.

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Chapter 10

Pullulan-based hydrogels Anca Giorgiana Grigoras "Petru Poni" Institute of Macromolecular Chemistry, Iasi, Romania

10.1 Introduction A polymeric hydrogel is a hydrophilic crosslinked material produced in the form of beads, sheets, or films that does not dissolve in water. The absorption and retention of water in hydrogels is done in stages and involves the followings: the hydration of hydrophilic groups; the expansion of the tridimensional network; interaction of the exposed hydrophobic moieties with water; the absorption of additional water into pores; and the reaching of the equilibrium swelling degree [1]. The polymers possessing hydrophilic pendant groups such as eOH, eSO3H, eCONHe, and eCONH2e are ideal candidates for hydrogel design. The polymeric structures physically crosslinked by hydrogen bonds or electrostatic interactions, along with the chemically crosslinked ones using a crosslinking agent or an ultraviolet (UV) photoinitiator, acquire the ability to swell without dissolving when immersed in water [2,3]. In the last decade, the nanostructured hydrogels have found applications in the fields of drug delivery systems and scaffolds for tissue engineering. To be an effective scaffold, a hydrogel should be able to mimic the extracellular matrix topography, to induce minimum to zero toxicity and inflammatory response, and to achieve the best possible compatibility with the targeted tissue. Beside the biocompatibility, cytotoxicity, and immunological criteria, in order to carry and release a drug, the hydrogels should have some special physicochemical and biological characteristics: to swell and retain a large volume of water; to protect the drug from the hostile variations of pH and temperature of the reaction environment; to realize volume phase transitions in accord with the environmental conditions; and to control the release of drug [4,5]. Pullulan is a microbial polysaccharide produced by Aureobasidium pullulans yeast-like fungus. Depending on the composition of the culture medium of the microorganism, various types of pullulans with variable molecular masses in range of 1
104 Dae2
105 Da can be synthesized and purified. Its chemical structure reveals a main chain composed by repeated units of a-D-glucopyranosyl-(1 / 6)-a-D-glucopyranosyl-(1 / 4)-a-D-glucopyranosyl-(1 / 4) with pendant hydroxyl groups susceptible for chemical substitution, mainly by grafting polymeration. Usually, the new substituted groups are introduced in the pullulan structure to obtain pullulan derivatives able to crosslink [6]. Pullulan-based hydrogels have some physicochemical properties like odorless, tasteless, colorless, flexibility proper for biomedical applications, but a limited biological activity. To overcome this disadvantage, the grafting of monomers on the pulullan chain is often resorted to Ref. [2]. Besides the grafting to a support, others technologies were adopted in hydrogels preparation by specialists: bulk polymerization, solution polymerization, or crosslinking, suspension polymerization or inverse-suspension polymerization, and polymerization by irradiation [7]. The crosslinking approaches included the click chemistry, the Schiff-base reactions, the amide crosslinking, the thiol-disulfide exchange, the enzyme-mediated crosslinking, or the photo-induced crosslinking [3]. A variety of physicochemical methods for characterizing materials have been exploited by researchers to subsequently certify or explain the biological properties of pullulan-based hydrogels. Structural analysis of hydrogels and its components was realized by different spectroscopic and spectrometric methods. In this way, the functional groups and interactions present in a system are investigated. Usually, certain bands characteristic of the IR/UV-Vis/NMR/absorption spectrum of the pullulan and the crosslinking agents undergo a widening which attests to the crosslinking taking place through a specific chemical reaction, but with the preservation of the pullulan structure intact. For example, from the peaks of OeH stretching, CeH stretching, and CeO stretching from 3315, 2922 and 999 cm1, respectively, characteristic to native pullulan, the widening of CeO stretching peak of composite hydrogels in the range of 945e1150 cm1 was ascribed to the etherification between pullulan and poly(ethylene glycol) diglycidyl ether [8]. In case of the crosslinking between the acryloyl group-modified nanogel consisting of cholesterol-bearing pullulan (CHP) and Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine. https://doi.org/10.1016/B978-0-323-95351-1.00002-8 Copyright © 2024 Elsevier Inc. All rights reserved.

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four-branched poly(ethylene glycol) with terminal thiol groups, the 1H-NMR spectrometry was useful to determine the degree of substitution of the acryloyl groups based on the signal ratio of the olefinic protons specific to acryloyl groups and the anomeric protons of pullulan glucose units [9]. The changes in the X-ray diffraction spectra of the pullulan, the interaction partners, and resulted pullulan-derived hydrogel provide information on the amorphous or crystalline nature of chemical compounds [9]. The thermal properties of hydrogels were monitored by differential scanning calorimetry, while the thermal stability of formulations consisted of pullulan-based hydrogels was evaluated by performing thermogravimetric analysis. Weight loss of hydrogels was observed due to vaporization of a water molecule followed by the degradation of the polysaccharide. By chemical transformation, the pullulan undergoes changes in the values of the glass transition temperature and the degradation temperature [10]. The morphology of tridimensional structure of hydrogels analyzed by scanning electron microscopy can attest to the porous nature of hydrogel and helps to investigate the distribution and dimensions of the pores. The dimensions and solution stability of particulated hydrogels are studied by dynamic light scattering and zeta-potential measurements [11]. Beside the physical appearance and clarity of the hydrogels, the degree of gelation, gelation time, and gelation temperature are investigated by researchers using the vial tilting or tube inversion method [9,12,13]. Rheological studies are performed to investigate the viscoelastic properties of hydrogels: storage modulus G’ and loss modulus G’. The settings of rheological experiments use protocols like these: (A) gelation kinetics: G’ & G’’ (Pa) versus time (s); constant 1% shear strain and 10 rad/s angular frequency; (B) amplitude sweep: G’ & G’’ (Pa) versus shear strain (%); angular frequency: 10 rad/s, strain: 0.1%e100%; (C) frequency sweep: G’ & G’’ (Pa) versus angular frequency (rad/ s); angular frequency: 0.1e100 rad/s, strain: 1%; (D) self-healing: G’ (Pa) versus time (s); and angular frequency: 10 rad/s, alternate strain: 0.1% and 400% for six cycles, 60 s each [11]. The drug loading into the polymeric matrix is a complex process, and its kinetic could be modulated with different models like pseudo-first-order and pseudo-second-order. In addition, the swelling capacity of a hydrogel and drug release studies under chemical and physical stimuli were monitored [11]. All these properties contributed to the performance of the hydrogel in a specific application. To recommend the safe use of these hydrogels in current medical practice, the researchers conducted a series of preliminary microbiological, immunological, histopathological, and cytotoxicity tests [14]. In this chapter, latest discoveries and applications of pullulan-based hydrogels were presented and reviewed. In the first part, emphasis was placed on those pullulan-based hydrogels with role of carrier for antiinflammatories, antidiabetics, therapeutic proteins, vaccines, and drugs for cancer therapy, wound healing, neurodegenerative diseases, and other neurological conditions. In the second part, the pullulan-based hydrogels for regenerative medicine applied in their pure form for reconstruction of dentin, cartilage, burn injuries, or as antiadhesion materials were discussed. In the third part, the pullulan derivatives hydrogels as supports for controlled and sustained release of drugs in regenerative medicine were evaluated.

10.2 Drug delivery systems 10.2.1 Antiinflammatory drugs Frequently, the liquid formulations like nanoemulsions and nanocapsules have not proper consistency for cutaneous application. Thus, the group of Ferrari Cervi [15] prepared from aqueous solutions a film based on polymeric component and plasticizers, namely pullulan, PEG 400, and sorbitol, as vehicle for nanoemulsions or nanocapsules containing pomegranate seeds oil. This natural drug with antiinflammatory and antioxidant effects was tested in the treatment of atopic dermatitis of laboratory mice, artificially induced by 2,4-dinitrochlorobenzene. By the incorporation of nanocapsules in pullulan film, the biological properties of pomegranate seeds oil were enhanced. Also, they concluded that glycerol was not suitable as plasticizer because the mechanical performances of films were improved in case of sorbitol instead. Osteoarthritis is a chronic degenerative disease of movable joints, accompanied by pain and sever disability. Colchicine is one of antiinflammatory drugs used to prevent deposition of calcium crystals in joints and to stop the progression of osteoarthritis by decreasing the cytokines production, which is related with the bone turnover markers involved in the remodeling of bone after resorption and replacement by a new bone with little changed shape. The research group of Mohamed et al. exploited new therapeutic option, more efficient than conventional therapies, which promoted the bioavailability of colchicine by transdermal route [16]. Thus, they developed an organic-inorganic hybrid delivery system loaded on cotton bandage. The fabric was dipped in mixed solution of carboxyethyl chitosan with colchicines-silica nanoparticles, and then treated with oxidized pullulan (OP) solution resulted from interaction of pullulan with sodium

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periodate. After that, the cotton fabric treated with resulted self-healing hydrogel has been portioned on poly(propylene) bags at room temperature for 2 h. In this way, patches loaded with 0.5 mg of drug per each cm2 resulted. It was supposed that the hybrid gel can be formed in situ on the surface of cotton fabric or by a Shiff base crosslinking reaction between the free aldehyde groups of pullulan and the amino groups of carboxyethyl chitosan by imine bonds as revealed the peak of 1629 cm1 from Fourier transform infrared (FT-IR) spectra. The silica particles were spherical in shape with pores diameter between 20 and 50 nm for mesoporous zones, and about 5 nm for microporous zone. Encapsulation efficiency for colchicine in mesoporous silica nanoparticles was of 67.4%. After drug encapsulation, the average size of silica particles increased from 129 to 167 nm, while the size distribution decreased from 0.32 to 0.22 value, and zeta potential changed from 26.6 to þ5.83 values. A positive zeta potential facilitated a better transdermal interaction of drug with abdominal skin of lab animals. The kinetic study for released drug from patches through the skin in a settled time was in good correlation with the Peppas model. The Peppas release exponent n ¼ 0.744 indicated an anomalous drug release mechanism governed by colchicine diffusion through hydrated pores of mesoporous silica nanoparticles and through the hybrid hydrogel matrix, too. The effect of the prepared patches on locomotor activity and oxidative stress biomarkers on monoiodoacetate (MIA)induced osteoarthritis in model rats was monitored. Using the new patches composed from hybrid hydrogel and drugloaded silica nanoparticles, the locomotor activity was improved with 74%, while the serum levels of nitric oxide and malondialdehyde were reduced by 28% and 34% and serum level of glutathione was increased by 23% compared with the positive control group composed from rats with MIA-induced osteoarthritis. All results are the basis of a new approach for an efficiently transdermal delivered therapy in case of osteoarthritis.

10.2.2 Antimicrobial drugs Antimicrobial drugs could be a part from hybrid hydrogels designed in the form of beads or membranes. Duceac and the team prepared hybrid structures with tunable pore size consisting of hydrogel beads based on chitosan (C) core covered of 6-carboxy pullulan by physical bonds (CPP) or 2,3-dialdehyde pullulan (CPT) by chemical bonds. It seems that the swelling degree of control chitosan beads was twice as high as that of the beads covered with OP [17]. These hydrogel beads were tested as hemostatic material and drug delivery system for bacitracin or neomycin, and even ibuprofen. Using several classical and empirical swelling models for the analysis of ibuprofen release mechanism, it was concluded that all three types of beads exhibited a non-Fickian behavior, known as an anomalous transport governed by polymer relaxation rather than diffusion of the solvent or drug into the network (Fig. 10.1). This limited swelling capacity of hybrid hydrogel beads was a consequence of decrease of the pore size by crosslinking reactions and of reducing of elasticity of networks by a limitation of flexibility of macromolecular chains (Fig. 10.2).

FIGURE 10.1 Ibuprofen release profiles performed in phosphate-buffered saline PBS, pH ¼ 7.4, at 37 C from C, CPP, and CPT hydrogel beads (C ¼ chitosan; CPP ¼ chitosan covered by periodate-oxidized pullulan; CPT ¼ chitosan covered by TEMPO-oxidized pullulan; TEMPO ¼ 2,2,6,6tetramethyl-1-piperidinyloxy). Reprinted from Duceac IA, Verestiuc L, Coroaba A, Arotaritei D, Coseri S. All-polysaccharide hydrogels for drug delivery applications: tunable chitosan beads surfaces via physical or chemical interactions, using oxidized pullulan. Int J Biol Macromol 2021;181:1047e62, Copyright (2021), with permission from Elsevier.

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FIGURE 10.2 SEM images of individual beads, their internal morphology, and the average bead and pore size dimensions (C ¼ chitosan; CPP ¼ chitosan covered by periodate-oxidized pullulan; CPT ¼ chitosan covered by TEMPO-oxidized pullulan; TEMPO ¼ 2,2,6,6-tetramethyl-1piperidinyloxy). Reprinted from Duceac IA, Verestiuc L, Coroaba A, Arotaritei D, Coseri S. All-polysaccharide hydrogels for drug delivery applications: tunable chitosan beads surfaces via physical or chemical interactions, using oxidized pullulan. Int J Biol Macromol 2021;181:1047e62, Copyright (2021), with permission from Elsevier.

The antibiotic-loading performance increased in order: 6.3%, 8.3%, 9.6% for C, CPP, CPT, respectively, and all types of beads recorded bactericidal activity against Staphylococcus aureus. Shafique et al. developed hydrogel membranes for wound healing based on pullulan, hyaluronic acid, poly(vinyl alcohol), and glycerol in different formulations, as support for embedding chitosan nanoparticles loaded with cefepime. The resulted wound dressings should be nonallergenic, nonadherent, and easy to remove without the damage of restored tissues. Along with the other components of hybrid hydrogels, responsible for a proper ventilation of wound and an optimum moist environment, this cephalosporin broadspectrum antibiotic provided the antibacterial protection. The hydrogel membrane was prepared by solvent-casting method, while lyophilized drug-loaded chitosan nanoparticles were added in a formulation during the formation of hydrogel membrane [18]. A medium particle size of 172 nm, a polydispersity index of 0.254, and zeta potential of þ27.8 mV for chitosan nanoparticles loaded with cefepime indicated a uniformly distribution by size and a good stability, properties specific to an ideal topical dressing. Scanning electron microscopy images revealed the surface morphologies of nanoparticles and hydrogel membrane, too. In cross-section of hydrogel membrane, it was observed a highly interconnected porous network generated by physical crosslinking of pullulan, hyaluronic acid, and poly(vinyl alcohol). Also, the small pores from developed membrane assured an oxygen permeability between 7 and 14 mg/L and water vapor’s transmission rate between 2000 and 2500 g/m(2)/day, but not allowed the passage of microorganisms. In this way, the superinfection of the wound was avoided, the wound was ventilated, and the drug was released in a controlled and sustained manner in proportion of up to 80%. The thickness of dried hydrogel membranes varied from 0.39 to 0.48 nm for all formulations. Generally, the gel fraction decreased by increasing the amount of polymers. In return, the swelling ratio of hydrogel membranes increased with increasing concentration of hyaluronic acid or poly(vinyl alcohol) and with decreasing of pullulan content. All solution properties of hydrogel membranes were in relation with hydrophilicity which was balanced by contribution of each chemical group of polymers like hydroxyl groups of hyaluronic acid, pullulan and poly(vinyl alcohol), carboxylic groups of hyaluronic acid, and hydroxymethyl groups of pullulan. Mechanical properties of optimized hydrogel membrane indicated a thickness of about 0.45 mm, a procentual elongation of 210% with 7.60 MPa tensile strength, peak load of 3.67 N, and the young’s modulus of 0.036 MPa. Their flexibility was enhanced by pullulan, while the strength and the stretching power were related with the presence of poly(vinyl alcohol). Thermogravimetric analysis and differential scanning calorimetry of hydrogel membranes recorded a total weight loss of 30% 5e570 C thermal interval, and endothermic peaks around 260 C and 360 C, suggesting that hydrogel membranes were thermally stable without being decomposed. Hydrogel films loaded with cefepime developed zone of inhibitions of 19 mm, 22, and 24 mm against Pseudomonas aeruginosa, Escherichia coli, and S. aureus. To replace the use of conventional eye drops and to allow a prolonged release of the drug, the researchers designed a new solid in situ gelling system composed from nanofibers for the treatment of topical ocular diseases [19]. The system

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FIGURE 10.3 Schematic overview over different AmpB solubilization techniques (Step I) and nanofiber compositions (Step II) with corresponding number of formulation. Reprinted from Goettel B, Lucas H, Syrowatka F, Knolle W, Kuntsche J, Heinzelmann J, et al. In situ gelling amphotericin B nanofibers: a new option for the treatment of keratomycosis. Front Bioeng Biotechnol 2020;8:600384, Copyright (2020), with permission from corresponding author Karsten Mäder, [email protected].

was based on gellan gum/pullulan electrospun nanofibers with 87.5% high porosity, narrow thickness distribution of 225e450, and 317 nm averaged diameter, which it is molded on the particular 3D shape of the eyeball and started to show gelling behavior immediately after administration, and finally was dissolved. This matrix will allow a homogeneous and prolonged drug distribution on cornea. Physiochemical properties of gellan gum can be enhanced by blending with pullulan to prepare not only hybrid hydrogels for antimicrobials delivery but also scaffolds for cell adhesion or tissue healing. Amphotericin B is an amphiphilic antifungal drug used against proliferation of some Ascomycetes species from Issatchenkia orientalis genus like anamorphous Candida krusei and Saccharomyces krusei involved in keratomycosis disease. This water-soluble drug was part of some formulations consisted from electrospun gellan gum/pullulan nanofibers designed by Goettel and his team [20]. To increase the drug encapsulation into in situ gelling nanofibers, different strategies to solubilize the drug were tested before electrospinning step, namely sodium cholate addition to drug (Ib), using of PLGA nanoparticles loaded with drug (Ic), or using drug-Eudragit polyelectrolyte complex (Id) (Fig. 10.3). Eudragit is the trade name of a cationic copolymer of methacrylic acid and methyl methacrylate which increases the mucoadhesive properties of drugs in gastrointestinal tract. Biodegradable and biocompatible PLGA is another copolymer of lactic acid and glycolic acid used as excipient in pharmaceutical formulations. All electrospun fibers showed mean fiber diameters in the 400e600 nm range. Only formulation (Id) based on polyelectrolyte complex Amphotericin B and Eudragit assured a proper drug loading of 0.68% for pullulan-gellan gum fibers of IId type. This value was comparable with the performance of existing eye drops in trade used for the treatment of keratitis. Microbiological plate diffusion tests testified a significant inhibition of I. orientalis growth only for complex-loaded fibers of (IId) type. In vitro cytotoxicity experiments with a multistratified epithelium cell model highlighted a superior tolerance of complex-loaded nanofibers compared with the conventional eye drop formulations. In addition, the authors tested the influence of an electron beam of 25 kGy used for sterilization of resulted nanofibers at room temperature. Drug content in complex-loaded fibers, after and before sterilization, was determined by HPLC, while modifications of molecular weights and polydispersity indices for pullulan and gellan gum after electron beam treatment were monitored by AF4 coupled with MALLS. Decreases of polydispersity index and of molecular weight after sterilization from 406 to 112 kDa for pullulan, and from 261 to 200 kDa in case of gellan gum, suggested a fragmentation of macromolecular chains such that their size distribution became narrower, but a sufficient resistance to this dried sterilization method.

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Usually, gels as semisolid formulations have low retention time on the epithelium and generate a leaking sensation and uneven loaded dose. The gels used for the treatment of superficial vaginal fungal infections should have mucoadhesive properties to overcome the mentioned disadvantage. The researchers proposed for the treatment of vulvovaginal candidiasis some formulations composed from clotrimazole-loaded Eudragit nanocapsules dispersed in a Pemulen (R) TR1/ pullulan-blended hydrogel [21]. The emulsifying and mucoadhesive Pemulen macromolecule is an anionic copolymer of C10eC30 alkyl acrylate crosslinked with allyl pentaerythritol and of acrylic acid blocks. Varying the proportion of the main components, the following semisolid formulations were obtained and labeled as follows: hydrogels containing clotrimazole-loaded nanocapsules, and pullulan at 3% and 1.5% (HGPP3-CTZ-NC and HGPP1.5-CTZ-NC); hydrogel containing clotrimazole-unloaded nanocapsules and pullulan at 3% (HGPP3-CTZ); and hydrogels containing clotrimazoleloaded nanocapsules, but without pullulan (HGPNC-CTZ). The formulations appeared homogeneous and were compatible with vaginal pH range which is between 4.5 and 5.5. The HGPP3-CTZ-NC and HGPP3-CTZ formulations maintained their initial 5.5 pH during 60 days of storage. Regarding their dimensions, the mean diameter of formulations was in the range of 250e320 nm with a polydispersity index of 0.25e0.6. Concerning the drug quantification in the hydrogels, the clotrimazole content of formulations was about 1 mg per gram of hydrogel. Analyzing the drug release profiles, it was observed that, after 8 h, the HGPP3-CTZ-NC hydrogel released 20.14  2.33 mg/cm2 clotrimazole, while the HGPP3-CTZ released 70.63  6.2 mg/cm2 clotrimazole. In relation to the mucoadhesion, it was observed that a higher concentration of pullulan induced a higher adhesion of formulation to the mucin gel in such manner that 55.09% of clotrimazole from HGPP3-CTZ-NC and 43.49% of drug from HGPP1.5- CTZ-NC adhered to the mucin gel. Based on the permeation or penetration ex vivo study on cow vagina mucosa, it was showed that the drug-loaded nanocapsules remained on the surface of vaginal mucosa. Using the inspired association between the two mucoadhesive polymers pemulen and pullulan, resulted an effective drug carrier system with promising results for local application and with minimal systemic absorption. To avoid the systemic toxicity induced by intravenous antibiotics, some authors proposed to use drug-loaded hydrogels wound dressing able to release a higher amount of drug to the wound site. Mert et al. synthesized novel hydrogels via UV crosslinking graft copolymerization of poly(acrylic acid-co-itaconic acid) onto pullulan using ammonium persulfate as the initiator and N,N0 -methylenebisacrylamide as the crosslinking agent [2]. Drug loading of Pu-g-p(AA-co-IA) resulted hydrogels were prepared by directly addition to the reaction medium of ampicillin sodium salt solution (0.1 wt%), a broadspectrum antibiotic highly soluble in water used as model drug. In thermal behavior of samples analyzed by DSC and TGA, specific thermal transitions were observed. Pure drug dehydrated in the 60e130 C thermal range and melted in the 228e231 C thermal range, while pullulan melted at 280 C and unloaded hydrogels recorded exothermic peaks at 137 C and 220 C. The crystalline drug was homogeneous dispersed through the polymer network because the sharp exothermic peak attributed to dehydration of ampicillin disappeared from the thermograms of drug-loaded hydrogels. Ampicillin-loaded hydrogels showed two zone of thermal stability; initially, they recorded a mass loss of 28% in 138e200 C thermal range, and then a major degradation with 45% mass loss in 325e450 C thermal range was observed. The prepared hydrogels were transparent enough to check the condition of the wound without removing the dressing. Also, according to MTT assay, they were nontoxic because 80% of mouse fibroblast cell lines L929 remained viable in the presence of grafted hydrogels. The pullulan-grafted copolymer hydrogels had swelling capacities of 700% at different pH values which simulated the skin layer (pH 7.4) and the moistened wound environment (pH 5.2 for wound exudate). From the series of mathematical models tested, it seems that ampicillin release profile best fit on Higuchi model related with a Fick’s law of diffusion. It was observed that, after seven days, drug-loaded pullulan-grafted copolymer hydrogels released 67% of ampicillin. Crystal violet was introduced in a biocompatible hydrogel system fabricated by one-pot incorporation of poly(dopamine) fibers in pullulan hydrogel [8]. This drug model has antifungal, antibacterial, and vermicide properties being used as topical antiseptic or as dye in Gram staining of bacteria. First, solutions of pullulan with molecular weight of 500 kDa in NaOH were mixed with NaOH solutions of dopamine hydrochloride and stirred at room temperature for 24 h. Then, poly(ethylene glycol) diglycidyl ether was added into these mixtures and promotes the crosslinking of pullulan, at the same time as the hydrogel reinforcement with the poly(dopamine) fibers resulted by oxidation (Fig. 10.4). The chemical crosslinking process was an etherification reaction of pullulan hydroxyl groups with epoxy groups of poly(ethylene glycol) diglycidyl ether. Nevertheless, crosslinking process by etherification reaction was competitively hindered by the hydrogen bonds between pullulan hydroxyl groups and phenolic and amine groups of poly(dopamine) fibers such that with the increase in the percentage of fibers, hydrogels with a lower degree of crosslinking resulted. The composite hydrogels with variable fiber percent and labeled as PHG-PDA1, PHG-PDA2, and PHG-PDA3 were rinsed with pure water until the unreacted hydrophilic crosslinker was removed and finally the mass fraction of reacted

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FIGURE 10.4 The whole process to prepare (A) poly(dopamine) PDA fibers, (B) pullulan hydrogel PHG, and (C) composite hydrogel PHG-PDAs. Reprinted from Su T, Zhao W, Wu L, Dong W, Qi X. Facile fabrication of functional hydrogels consisting of pullulan and polydopamine fibers for drug delivery. Int J Biol Macromol 2020;163:366e74, Copyright (2020), with permission from Elsevier.

crosslinking agent in pullulan hydrogel was estimated at approximately 9 wt%. By the introduction of poly(dopamine) fibers into 3D hydrogel networks, all physicochemical properties were improved compared with pure pullulan hydrogels. Thus, all samples had a rheological behavior typical to hydrogels because the value of the storage modulus (G0 ) was higher than loss modulus (G00 ) and presented high compressibility and elasticity when 80% strain was applied before fracture. With increasing PDA content, the elastic modulus decreased, while the average pore sizes increased from 231 mm up to 350 mm and the morphology of honeycomb-like compact structure with pores of about 205 mm in diameter of pullulan hydrogel changed with more rough structure specific to composite hydrogel. Studying the swelling and deswelling profiles, it was observed that the swelling rate of composite hydrogels in water proportionally increased with the content of poly(dopamine) fibers such that they reached an equilibrium state within 250 min. Also, the swelling ratio for these hydrogels in any other aqueous salted solution with NaCl, MgCl2, or AlCl3 was reduced.

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Finally, all these features of the new polymeric system allowed the controlled and efficient release of the drug. In case of new hydrogels composed from crosslinking pullulan and poly(dopamine) fibers, the kinetics of drug loading best fitted on the pseudo-second-order model and Langmuir isotherm model, suggesting that the drug was loaded mainly on the surface of hydrogels by chemisorption. The drug release from hydrogels was a pH-responsive process such that at pH 7.4 and pH 5.0, the cumulative release was 60.3% and 87%, respectively. In addition, it was demonstrated by cytotoxicity tests that the new hydrogels were nontoxic even they are used in a high concentration of 3.4 mg/mL.

10.2.3 Drugs for imagistic diagnostic Usually, the degradable nanogel-crosslinked hydrogels are used for scaffold materials for controlled drug release in regenerative medicine, but have low loading capacity in case of hydrophilic drugs. This disadvantage could be overcome by using liposomes which can encapsulate a relatively large amount of hydrophilic molecules. In another study, CHP nanogel was implied in design of hybrid gels used to entrapped a fluorophore model hydrophilic drug, namely 8aminonaphthalene-1,3,6-trisulfonic acid disodium salt ANTS, utilized as a neuronal tracer [22]. First, acryloyl-bearing cholesterol-bearing pullulan (CHPOA) nanogel was prepared from CHP and 2-(acryloyloxy)ethyl isocyanate (AOI). Then, ANTS was entrapped in multilayer liposomes based on cholesterol and dimyristoylphosphatidylcholine DMPC. Finally, CHPOA nanogel and drug-loaded DMPC liposomes were mixed with pentaerythritol tetra(mercaptoethyl)polyoxyethylene PEGSH as crosslinker to form hybrid gels. The release profile of liposomal drug was indirectly monitored by fluorescence spectroscopy and p-xylene-bis-pyridinium bromide DPX, a cationic dye quencher often coupled with ANTS. Actually, the hybrid hydrogel allowing a sequential dual release of carried molecules under physiological conditions because uncomplexed CHPOA nanogel was gradually released from hybrid hydrogel, followed by the release of nanogelcoated liposome complex.

10.2.4 Drugs for wound healing Being the largest human organ with a role in the first-line defense, the skin is often exposed to aggressive factors such as extreme temperatures, trauma, or chemical agents. The topical or systemic antibiotics, wound dressings, liquid dosage forms, or cell stem therapies represent available strategies to manage wound-healing process so as to overcome the phases of hemostasis and inflammation and to stimulate the proliferation of new cells for tissue remodeling. Depending of each type of wound, a specific physical form of dressing could be recommended to complete the wound healing: film, foam, ointment, hydrocolloid, sponge, nanofibers, or gel [23]. Hydrogels based on pullulan, dextran, chitosan, cellulose, hyaluronic acid, collagen, and gelatin are biocompatible, biodegradable, nontoxic materials with potential wound-healing efficacy. In the case of injured tissue, the bioactivity of the natural or synthetic polymer component in a topical hydrogel may be diminished due to peptidase and collagenase secreted by the wound so that healing is an uncertain and chaotic process, especially for diabetic patients. Therefore the use of in situ-forming hydrogels is preferred. These thermoresponsive hydrogels facilitate the drug encapsulation at room temperature and the gelation at body temperature and have some advantages: can be easily applied to irregularly shaped wounds, and the incorporated drug is released for a longer period of time. Shah et al. [24] fabricated in situ injectable hydrogels based on hyaluronic acid, pullulan, and Pluronic F127 (Poloxomers P407), loaded with curcumin (diferuloylmethane), a yellow pigment with antiinflammatory, antioxidant, and antibacterial properties. They used the cold method or solvent-casting method for hydrogel preparation. Hyaluronic acid, pullulan, and Pluronic F127 (Poloxomers P407) were dissolved in distilled water and formed self-assembling micelles, while drug was dissolved in methanol. All formulations were mixed and equilibrated at 4 C for 48 h in order to eliminate air bubbles and foam, and then gradually warmed from 25 C up to 35 C to allow the gelation by physical method. These hydrogels were freeze-dried, and powder samples were preserved for further studies. Compared with control hydrogels, the curcumin-loaded HA-Pu-based-F127 hydrogels recorded better rheological and mechanical properties and provided a sustained curcumin release in case of diabetic wounds after subcutaneous administration of injectable hydrogel, which was translated with a significant collagen regeneration, with organized formation of dermis and a fully developed epithelial layer. Focal adhesion kinase facilitates transformation of biochemical signals in mechanical forces and controls the cell mobility. In the complex wound repair process, the mechanical signals trigger a fibrotic response and the undesired scar formation. The local injections with focal adhesion kinase inhibitor FAKI have been shown to be ineffective in treating large wounds resulting from burns, explosions, or surgical excisions. Thus, the researchers proposed a hydrogel system based on pullulan and collagen to deliver this low molecular inhibitor on a bigger injured surface [25]. The antiscarring

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agent was encapsulated in porous hydrogel scaffolds by molecular imprinting technique, when a slow controlled and sustained drug release was desired, like in case of surgical excision wounds, or by surface incorporation method for a rapid release of drug from matrix to the superinflamed burn injuries which requires changing the drug-loaded dressings as often as possible. It was observed that the mechanical integrity of healed skin was improved after the treatment of excisional wounds with focal adhesion kinase inhibitor entrapped in hydrogel. Also, while the pullulan-collagen matrix promoted the healing, the focal adhesion kinase inhibitor inhibited the expression of inflammatory cytokine, especially in burn injuries. In conclusion, this system allowed wound healing deeply in dermal layers, on large surfaces and without scars.

10.2.5 Drugs for cancer therapy Due to the special structural and functional characteristics of the tumor cell, the hydrogels loaded with anticancer drugs must be smart stimuli sensitive. For example, 5-fluorouracil, recombinant mouse interleukin-12, and doxorubicin are molecules used for the treatment of fibrosarcoma, myeloma, and other types of cancer, included by researchers in various hydrogels, especially based on polysaccharides [4]. For localized treatment of tumors, Liang et al. developed injectable mucoadhesive and pH-responsive hydrogels based on OP and chitosan-grafted-dihydrocaffeic acid (CS-DA) via a Schiff base reaction [12]. These hydrogels were prepared and loaded with doxorubicin according to Scheme 10.1. Varying the concentration of crosslinker, namely 0.2%, 0.6%, and 1% (w/w) OP, three hybrid hydrogels were obtained: CS-DA/OP2, CS-DA/OP6, and CS-DA/OP10. The gelation time of these hydrogels (20  5 s, 9  3 s, and 4  2 s) determined by tube inversion method decreased with the increase in the content of crosslinker in hydrogel. After preliminary treatment of tissue with CS-DA gluing solution and lightly pressing the OP solution for 5e10 s at room temperature, a hydrogel formed spontaneously. In addition, the detachment stresses of CS-DA/OP hydrogels were twice as larger as for control CS/OP hydrogels. This adhesive strength test was the basis for in vivo gelation and tissue adhesion test. After subcutaneous injections of polymer solutions, the laboratory rat skin presented globular protuberances indicating a completely transformation into hydrogels hard to remove with forceps.

SCHEME 10.1 Synthesis of CS-DA/OP hydrogels and the drug encapsulation process (CS ¼ chitosan; DA ¼ dihydrocaffeic acid; OP ¼ oxidized pullulan). Reprinted from Liang Y, Zhao X, Ma PX, Guo B, Du Y, Han X. pH-responsive injectable hydrogels with mucosal adhesiveness based on chitosan-grafted-dihydrocaffeic acid and oxidized pullulan for localized drug delivery. J Colloid Interf Sci 2019;536:224e34, Copyright (2019), with permission from Elsevier.

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The swelling behavior of the unloaded CSDA/OP hydrogels was tested in simulated physiological microenvironments. The normal tissues are characterized by a pH ¼ 7.4, while extracellular pH of a tumor site ranges from 5.7 to 7.0. In pH 5.5 phosphate-buffered saline (PBS) solution, swelling ratio of 4597%, 3442%, and 2406% for CS-DA/OP2, CS-DA/OP6, and CS-DA/OP10 hydrogels decreased with the increase of crosslinking density in the hydrogels, while in pH 7.4 decreased to 1745%, 1467%, and 1238%, respectively. The structure of hydrogels was highly porous, with pore size of 305  45, 266  40, and 254  38 mm for CS-DA/OP2, CSDA/OP6, and CS-DA/OP10 in acidic medium, and pores of 170  27, 149  30, and 127  22 mm in pH 7.4 PBS. The release percentages of hydrophobic doxorubicin from CS-DA/OP10, CS-DA/OP6, and CS-DA/OP2 hydrogels after 6 h were 47%, 52%, and 63%, respectively, in accordance with the size of pores. After 32 h, all the hydrogels released about 76% of loaded drug. Because, for example, from CS-DA/OP6 hydrogel, at 37 C after 60 h, 52% of doxorubicin was released in pH 7.4 PBS and 87% in pH 5.5 P BS, it was concluded that these hydrogels released more drug around or inside tumoral site than in normal tissue. Then, the therapeutic activity of encapsulated and subsequently released doxorubicin from the hydrogels was tested on colon tumor cells (HCT116 cells) and was pointed out the anticancer effectiveness of drug-loaded hydrogels. The special properties of stimuli-responsive hybrid materials can be tuned by the modification of molecular structure or by variation of temperature, pH, ionic strength, or light. The hydrophilicity of pullulan not allows the loading with drugs. By modification of pullulan with self-organized thermoresponsive polymers like poly(N-isopropylacrylamide) and poly(L-lactide), it is possible to introduce hydrophobic segments into pullulan chains. In aqueous environment, resulted amphiphilic copolymers have a core-shell structure composed from a hydrophobic inner core and a hydrophilic outer shell. Nanogels derived from pullulan-g-poly(L-lactide) copolymers were designed by Seo et al. [26] in order to test the temperature-triggered release of doxorubicin from hydrogel-carrying drug. They used two type of poly(L-lactide) with different polymerization degree (0.81 and 1.42) and obtained two hydrogels labeled as PLP1 and PLP2. The mean hydrodynamic diameters of PLP1 and PLP2 nanogels were 121 and 163 nm at 25 C. These values fall within the maximum permissible particle size range used for the targeted treatment of solid tumors, being less than 200 nm. Due to hydrophobic lactide portions, the size of hydrogels varied more or less with temperature variations. Thus, at 32 C, the PLP2 hydrogels recorded higher dimensions, while the PLP1 hydrogels’ dimensions remained under 100 nm until 40 C. In the same time, turbidity of the PLP2 hydrogels at 37 C was more visible than at 25 C based on a secondary aggregation phenomenon between nanogels (Fig. 10.5). The critical association concentration of PLP2 was 0.02 mg/mL at 37 C and 0.06 mg/mL at 25 C, respectively, indicating an increased number of stable micelles with decrease in critical association concentration.

FIGURE 10.5 Molecular structure of the pullulan-g-poly(L-lactide) copolymer PLP and the schematic diagram for anticancer drug release from thermosensitive PLP nanogels by triggering temperature. Reprinted from Seo S, Lee C-S, Jung Y-S, Na K. Thermo-sensitivity and triggered drug release of polysaccharide nanogels derived from pullulan-g-poly(L-lactide) copolymers. Carbohydr Polym 2012;87:1105e11, Copyright (2012), with permission from Elsevier.

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For hydrogels based on from pullulan-g-poly(L-lactide) copolymers, the drug content of 4%e5% and loading efficiency about 40% were recorded. In vitro drug release tests suggesting a linear increased cumulative doxorubicin release from nanogels in the same sense with temperature such that faster, easier, and greater drug release occurred at higher temperatures when each particles became more compact. The thermoresponsive effect of doxorubicin-loaded PLP1 hydrogel on killing cancer HeLa cells was investigated at 37 and 42 C. The doxorubicin amount released at higher temperature after 6 h was more than at 37 C such that a stronger cytotoxic effect due to the released drug was observed at 42 C. In conclusion, doxorubicin was released from hydrogels, and then was internalized in the tumoral cells in such manner that the new thermosensitive self-assembled nanogels were effective for the destruction of cancer cells at higher temperature. Combination drug therapy became a standard clinical protocol because the benefits for patients come from the synergistic or summed-up action of the drugs that make up the hybrid system. In this way, for example, cisplatin combined with 5-fluorouracil, epirubicin, or doxorubicin showed higher efficiency against tumors and enhanced cytotoxicity than the single drug use. Thermoresponsive- or pH-dependent in situ hydrogels represent reservoirs for drugs which collapse under stimuli action and induce the controlled release of drugs upon injection in vivo. Cheng et al. prepared by Schiff base reaction in aqueous medium some hydrogels based on ε-poly(L-lysine), aldehyded pullulan, and branched poly(ethyleneimine) [27]. The ε-PL/A-Pul/BPEI hydrogels were loaded with cisplatin and doxorubicin by two ways: physical adsorption and Schiff base reaction. The temperature of the reaction solution influenced the gelation time. Thus, the reaction solutions stored at room temperature gelled in 38 s, while the reaction solutions stored at 4 C gelled in 280 s. An acidic environment with pH range of 5.7e7.0 is specific to solid tumor site and its adjacent tissues, while the physiological environment with pH 7.4 is specific to normal tissues. So, the authors watched few days the process of degradation of hydrogels in vitro at pH 7.0 and pH 7.4 and found the following: the degradation rate increased at acidic pH, and the adding of the drugs enhanced the weight loss rate and prolonged the degradation time. After subcutaneous injections of the reaction solutions, in situ-formed gelatinous globules appear in 5 min onto dorsal areas of laboratory rats, which then start to degrade in vivo in fouresix days. All hydrogels presented a macroporous structure with average pore sizes between 58 and 89 nm function of them drugloading degree. All free- and drug-loaded hydrogels reached a swelling equilibrium in about 400 min at room temperature such that the swelling ratio for ε-PL/A-Pul/BPEI hydrogels, cisplatin-loaded hydrogels, doxorubicin-loaded hydrogels, and dual drugs-loaded hydrogels were 1449%, 1353%, 1538%, and 1452%, respectively. The encapsulation efficiency of hybrid hydrogels varied from 47%, 55%, to 78% in case of cisplatin, doxorubicin, and dual-drugs system. Also, in vitro drug release ratio in acidic pH was higher than in physiological pH. Regardless of their form of encapsulation in mono or dual system, once released from the gels, the drugs maintain their antitumor activity almost unchanged. If the antitumor activity of released cisplatin was metastable and of doxorubicin increased with time, the best anticancer activity was recorded by the cisplatin-doxorubicin system due to their synergism. All these properties indicated that the dual drugs-loaded hydrogels are proper delivery systems for the acidic environment of tumors, improving thus the treatment of cancers.

10.2.6 Drugs for diabetic disease Along with other diseases of the century such as cancer, AIDS, or obesity, diabetes remains a disease with sociomedical challenges that requires the invention of controlled release systems of drugs included in transport systems, depending on the serum glucose level at a certain time. The hydrogel-type systems transporting the antiglycemic drugs were enabling to improve the diabetes therapy. Yi and the team prepared a smart hybrid hydrogel based on covalently modified carboxylated pullulan and concanavalin A in order to encapsulate an antidiabetic drug which inhibit the activity of a-glucosidase from small intestine and consequently reduce the postprandial glucose blood level, namely acarbose. Pullulan with carboxylic acid group substitution degree of 0.2 was chemically crosslinked by concanavalin A using an amidation reaction, and 1-ethyl-3-(3dimethylaminopropyl) carbodiimide and N-hydroxysuccinimide as coupling agents. The release of acarbose from hydrogels was tested in simulated physiological fluids with different pH values. It seems that acarbose-encapsulated hydrogel was stable in gastric fluid because only 20% of drug was released, while in small intestine fluid, the drug was released/detached in proportion of 80% [28]. The rheological measurements in dynamic and steady frequency scanning modes showed that this glucose-sensitive hydrogel loaded with acarbose was stable in gastric medium because G0 and G00 moduli were almost unchanged, while the hydrogels immersed in small intestine simulated fluid recorded an increase of G0 and G00 values in the frequency range of 0.1e10 Hz related with an increase of elasticity of the three-dimensional network.

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SCHEME 10.2 (A) Preparation routine of the carboxylated pullulan-concanavalin A hydrogel CPUL-ConA. (B) A proposed mechanism of smartly controlled release of insulin from the insulin/CPUL-ConA hydrogel. Reprinted from Lin K, Yi J, Mao X, Wu H, Zhang LM, Yang L. Glucose-sensitive hydrogels from covalently modified carboxylated pullulan and concanavalin A for smart controlled release of insulin. React Funct Polym 2019;139:112e19, Copyright (2019), with permission from Elsevier.

The enhanced elasticity of drug-loaded hydrogel provided to it an improved mechanical strength against the peristaltic movements of the stomach and small intestine. The same type of hydrogel based on pullulan derivative and concanavalin A was used by Lin et al. for loading and release of insulin [29]. Initially, pullulan was esterificated with succinic anhydride to produce a pullulan derivative with carboxyl groups that were subsequently covalently linked to amino groups of concanavalin A by an amidation reaction in the presence of 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide EDC and N-hydroxysuccinimide NHS and activation with MgCl2, CaCl2, and MnCl2. In this way, the system EDC/NHS consisting of a catalyst and a stabilizer do not allow the leakage of concanavalin A (Scheme 10.2). The swelling behavior of CPUL-ConA hydrogels was tested in PBS solution (pH 7.4) with different concentrations of glucose. Thus, the swelling ratio of hydrogels increased with the increase of glucose concentration, suggesting that activated concanavalin A residues specifically bonded the free molecules of glucose and incorporating them into hydrogels. In addition, CPUL-ConA hydrogels immersed for 24 h at 37 C in PBS solution (pH 7.4) containing 20 mmol/L glucose have undergone a gelesol transformation. The preparation of ConA-CPUL hydrogels loaded with different content of porcine insulin (0.6 or 10 mg/mL) was realized by simple immersion of hydrogels in insulin solutions and mixture homogenization for 2 h such that in situ insulin-loading efficiencies of 19 and 122 mg/mg were recorded. The morphological analysis investigated by SEM revealed a smooth crosslinked network structure and uniform pores on the surface of original hydrogels, but a more compact structure in case of insulin-loaded hydrogels, with interior pores bigger than surface pores and solid insulin particles present on the surface. From rheological point of view, all hydrogels presented a dominant elastic behavior, but at the same time, the storage (G0 ) and loss (G00 ) moduli values of the original hydrogels were larger than those of insulin-loaded hydrogels because the CPUL-ConA molecules became more free when insulin was added, and the network structure of hydrogels was affected. Also, the insulin from hydrogels induced a shear thinning of the hydrogels. In case of a diabetic patient with weight of 50e100 kg, the basal insulin dose is 25e50 U per day. The weight of porcine insulin released from the hydrogels could be converted to active units according to formula 1 U ¼ 38.46 mg of human insulin. The analysis of in vitro insulin release profiles of drug-loaded hydrogels in a medium that simulate the human hyperglycemic concentration showed that the hydrogel with 0.6 mg/mL porcine insulin does not succeed to meet the basal insulin dose, but the hydrogel containing 10 mg/mL insulin rapidly released the insulin in hyperglycemic medium, and the required basal insulin level was reached after about 1.2 h.

10.2.7 Drugs for neurodegenerative diseases and other neurological conditions Management of some diseases imposed the use of chronotherapy, meaning the administration of the drug on demand, in a controlled dose, when the symptoms become annoying. To prevent the disadvantages of oral administration of rivastigmine

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tartarate, a drug that combats the symptoms of Alzheimer’s disease, Patil and the team [30] proposed a novel electrically responsive transdermal delivery systems (ETDSs) for drugs. This system was composed from a series of layered elements in variable proportions: a poly(styrene) film as backing layer covered with a hydrogel reservoir for drug, and a ratecontrolling membrane for delivery of drug, which is in direct contact with the skin. The hydrogel reservoir was mainly represented by poly(acrylamide)-graft-pullulan copolymer prepared by free radical grafting reaction between pullulan and acrylamide. Subsequent alkaline hydrolysis of amide groups (eCONH2) of poly(acrylamide) into (eCOONa) groups gave the electro-responsive properties to the copolymer. To the copolymer thus hydrolyzed and then dried, methyl paraben, rivastigmine tartarate, 0.1 N HCl, and glutaraldehyde as crosslinking agent were added. After homogenization of mixture for 30 min, the obtained hydrogel reservoir was preserved in an air-tight container for future using. For the preparation of rate-controlling membrane, poly(ethylene glycol) 200 and glutaraldehyde were incorporated to the aqueous mixed solutions of native pullulan and poly(vinyl alcohol). The final solution was poured on glass plate, allowed to dehydrate about 48 h, and the film was stored in desiccators until subsequent use. Compared with the native pullulan, the main properties of hydrogel and membrane were improved. Thus, the thermal stability of grafted copolymer-based hydrogel was higher than pullulan. The pH range of developed translucent hydrogel reservoir was between 6.84 and 7.03, and drug content between 87% and 97%, depending on the proportion of glutaraldehyde. The thickness range (61 and 96 mm) and tensile strength of the rate-controlling membrane was directly proportional to the amount of glutaraldehyde. As the concentrations of poly(acrylamide)-graft-pullulan copolymer and of glutaraldehyde increased in hydrogel reservoir, the drug permeation tended to reduce. But applying pulsatile electrical stimuli of increasing intensity from 2 to 8 mA, the drug permeation rate increased because the negatively charged poly(acrylamide)-graft-pullulan copolymer collapsed and the loaded drug was released. Using this ETDSs instead of the traditional ones, a sustained plasma concentration of drug is expected to be maintained on demand, in a lesser dose and a lower frequency, and thus minimize the overall adverse effects. At the same time, the new technology could allow the incorporation and delivery of other molecules problematic from the point of view of their lipophilicity, ionic state, or molecule size. Pregabalin is an oral drug from gabapentinoids group which inhibits certain calcium channels. Having analgesic, anticonvulsant, and anxiolytic effects, it is mainly used to treat the neuropathic pain of diabetics and patients with zona zoster, epilepsy, fibromyalgia, and generalized anxiety disorder or preoperatory sedation in surgery. Cinay et al. developed stimuli-sensitive hybrid hydrogel systems for controlled delivery of pregabalin [31]. They integrated the CHPOA nanogel in the pH-responsive poly(methacrylic acid-g-ethylene glycol) P(MAA-g-EG hydrogels by two methods: bulk polymerization and surface-initiated polymerization. The second approach was advantageous for controlling the crosslinking density and thus resulting different swelling and drug release profiles. At pH 2.0, all hybrid hydrogels appeared opaque, but transparent and soft at pH 7.4. Their surfaces were porous facilitating the loading and release of drug. From morphological point of view, the hydrogels resulted from surface-initiated polymerization presented gradients in crosslinking density, and the pore sizes decrease with the nanogel addition. The gels polymerized in bulk have a variety of pore morphology, and the pore sizes increased with the addition of the CHPOA nanogel. The gel fraction values in all hybrid systems covered the range 50%e70%. Giving all these differences, the loading efficiencies of hydrogels were between 78% and 87%, and pregabalin released percent of 87%e99%. The human BJ fibroblast cells used to test the biocompatibility of hybrid hydrogels survived in a more number after 6 h, in the presence of the hydrogels produced by surface-initiated polymerization, in contrast with the gels from the other group.

10.2.8 Vaccine formulations Streptococcus pneumoniae is a pathogen antibody which colonizes the epithelium of airways and can trigger pneumonia, meningitis, otorhinolaryngological diseases, and even bacteremia in the human body. In medical practice, one of the routes of immunization against infectious diseases with respiratory transmission is to be nasal vaccination. An ideal nasal vaccine should deliver effectively and retain antigens in the nasal mucosa. In this regard, for example, some groups of researchers have developed studies for the design of a pneumococcal vaccine that combines the benefits of pneumococcal surface protein A (PspA) with those of a nontoxic polymeric transport system [32,33]. The nanohydrogel consisting of a cationic CHP, which forms a chaperone self-assembled nanogel in water, proved to be an ideal transportation system for PspA. The efficacy of the nanogel-based PspA nasal vaccine on laboratory animals was demonstrated by the following results: a reduced colonization and invasion by S. pneumoniae Xen10 in the upper and lower respiratory tracts; increased levels of PspA-specific serum immunoglobulin G (IgG); and the absence of curative protein in the olfactory bulbs or central nervous system after intranasal administration of vaccine. Until the use of these nasal vaccines in the human population, it was

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necessary to test them on nonhuman primates to demonstrate that the systemic and mucosal immune responses are induced and then the protective immunity is maintained for a period of time. Because the current injectable vaccines based on polysaccharides doesn’t succeed to develop an antigen-specific mucosal immune response, and consequently to inhibit the colonization of airways with S. pneumonia, some researchers proposed another drug delivery system based on PspA encapsulated in cationic cholesteryl pullulan nanogel, formulated as a nasal vaccine [34]. The cationic cholesteryl pullulan synthesized from pullulan contains 23% of amino residues and 1.3% of cholesterol residues, and has an estimated molecular weight of 120,000 Da, while the average molecular weight of nanogel is about 480,000 Da because each molecule of nanogel resulted from self-assembly of four molecule of modified pullulan. Actually, three PspA fusion constructs isolated from three bacterial strains were individually encapsulated in nanogel, and then were mixed together such that a trivalent nanogelPspA vaccine formulation resulted. The encapsulation ratio of PspA in the nanogelPspA formulation was of PspA1:PspA2:PspA3 ¼ 8:2:10 mg per dose ¼ 24 mL, reported for a nanogelPspA formulation of 1% nanogel in a 1.0 molecular ratio (nanogel:PspA). In this case, the nanogel performed the chaperone function, meaning that, after internalization of nanogel-PspA formulation by nasal epithelial cells, the protein component is released in their native form. This fact was confirmed by circular dichroism spectra of proteins which indicate a preservation of a-helical secondary structure after leaving the polymer transport system. It was found that the immunized lab mice with the new trivalent nasal vaccine formulation developed not only a systemic immune response, confirmed by detection of PspA-specific serum IgG antibodies, but even a mucosal immune response correlated with mucosal IgA antibodies. Using the same immunologically inert nanohydrogel of cholesteryl pullulan, which self-assembles in water, other authors developed a cancer vaccine based on the encapsulation of a synthetic long peptide antigen (LPA) with selectivity for medullary macrophages. The cholesteryl pullulan forms a stable complex with the peptide by hydrophobic interaction, helping its solubilization and long-term stabilization. After subcutaneous injections to laboratory animals, the selective nanogel-based vaccine was transported to the draining lymph node and absorbed by macrophages localized into the core of lymph nodes. The medullary macrophages participate in T cell immunity, stimulating a strong antitumor antigen-specific CD8þ T cells response. Experimental results indicated that the new cancer vaccine inhibited in vivo the tumor growth both in the prophylactic and therapeutic protocols. The second role of pullulan-based particulated hydrogel was to protect the antigens from overdegradation in phagosomes [35].

10.2.9 Therapeutic proteins Therapeutic proteins are successfully incorporated in self-assembled polysaccharides gels through hydrophobic interactions forming complexes with high colloidal stability. The self-assembled molecules are stimuli-sensitive and allow targeted drug delivery. In addition, the nanoscale polymer networks of the hydrogels can incorporate the therapeutic proteins in their bioactive shape minimizing their denaturation, aggregation, and enzymatic degradation. Using a bottom-up approach, Tahara et al. developed acryloyl group-modified polysaccharide nanogels as building blocks of a macrogel [36]. First, they prepared CHPOA by the modification of CHP in the presence of 2- AOI and di-Nbutyltin-dilaurate (DBTDL). In the second step, the CHPOA nanogels, self-assembled in water due to their hydrophobic cholesteryl groups, were photopolymerized using light with 365 nm-wavelength, in the presence of poly(ethylene glycol diacrylate) PEGDA as crosslinker and Irgacure 2959 as biocompatible photoinitiator of radical polymerization. The ester bonds linked the pullulan chains of CHPOA nanogels and the poly(ethylene glycol) chains of PEGDA. Thus, by crosslinking of CHPOA nanogels resulted novel CHPOA hydrogel particles and CHPOA macrogels. The macrogel has nanodomains and designable crosslinking points. Varying the concentrations of CHPOA nanogels and PEGDA crosslinker resulted nanogel-crosslinked hydrogel nanoparticles with diameters between 47 and 107 nm. This size range recommends them for future preparation of injectable hydrogel scaffolds for regenerative medicine or tissue engineering. To analyze the release profile of a model protein from the new hydrogels, the nanogel-crosslinked hydrogels containing insulin were incubated in PBS solutions with pH 7.4, 8.1, and 9.0. It was observed that protein was gradually released over seven days, and at higher pH values, the release rate increased. The authors suggested that actually the nanogels based on complexes of insulin with CHPOA were released from nanogel-crosslinked hydrogel by hydrolysis. The bottom-up method was also used by Shimoda et al. to prepare nanoparticles comprising assemblies containing CHPOA nanogel and pentaerythritol tetra(mercaptoethyl) polyoxyethylene PEGSH in variable proportion [37]. With the increase of solution concentration of resulted CHPOA-PEGSH nanoparticles from 2 to 5 mg/mL to 10 mg/mL, the solution state of system containing hybrid particles with size of 70e150 nm has undergone a transition to the gel state. Under higher concentrations which exceed 20 mg/mL, the CHPOA-PEGSH forms macrogels.

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The new resulted nanogel-crosslinked nanoparticles obtained from CHPOA nanogel and PEGSH crosslinker showed a “raspberry-like” morphology. In addition, the crosslinked hydrogels hydrolyzed under physiological conditions resulting CHPOA nanogels gradually released from the crosslinked constructs. The authors proposed the using of macrogels as scaffolds for tissue engineering, and application of nanogel-crosslinked nanoparticles for the preparation of injectable nanocarriers for controlled release of proteins. Morimoto et al. prepared a new self-assembled stimuli-responsive cholesteryl pullulan nanogel as protein delivery vehicle. It is about the acid-labile cholesterol-modified pullulan (acL-CHP) resulting from a laborious procedure presented as following: (a) Azido-modified pullulan (Pullulan-N3) with a substitution degree of 3.6 was the product of reaction between pullulan, 1,10 -carbonyldiimidazole and azidopropylamine; (b) Acid-stable cholesterol derivative resulted from the reaction of N,N0 -dicyclohexylcarbodiimide with 2-(prop-2ynyloxy)-acetic acid, cholesterol, and N,N0 -dimethyl-4aminopyridine; (c) Acid-labile cholesterol derivative was the product of reaction between acid stable cholesterol derivative with Tebbe reagent; (d) Reaction of azido-modified pullulan Pullulan-N3 with the acid-labile cholesterol derivative in the presence of sodium ascorbate, copper (II) sulfate pentahydride, and triethylamine solutions generated the acid-labile cholesterol derivative bearing pullulan (acL-CHP) via click reaction [38]. The solution properties of acS-CHP nanogel acL-CHP nanogel in water were monitored by dynamic light scattering. The precursor particles have size of 37 nm in diameter, while acL-CHP particles have 53 nm. Both types of nanogels have polydispersity index of about 0.2. They observed that the acL-CHP nanogel formed a stable complex under physiological pH conditions with bovine serum albumin BSA, and after acidification of medium, the self-assembled nanogel released the protein cargo. Also, this type of stimuli-responsive amphiphilic polymer was nontoxic. All these properties of hybrid nanogels recommended them for tissue engineering and drug delivery applications. Knowing that the pullulan-based nanogel works as a protein stabilizer, pullulan has a poor spinnability, and the fibers can be preserved for a long time, Shimoda et al. fabricated polysaccharide nanogel containing gelatin fibers by electrospinning a solution composed from CHPOA nanogels and gelatin type A. The natural crosslinker Genipin was used to enhance the stability of gelatin and nanogel containing gelatin fibers, and was a substitute for others synthetic crosslinkers like EDC or EDC-NHS system. Then, the horseradish peroxidase was encapsulated in these hybrid fibers [39]. The size of self-assembled nanogels uniformly dispersed throughout the nanofibers was about 60e80 nm. As the gelatin degraded in the presence of collagenase, the stably incorporated nanogels were gradually released from the fibers. After complexation of the protein with nanogels, the formed entities have dimensions of 20e30 nm. The tests showed that after electrospinning, the horseradish peroxidase remained stable in dry and cold conditions, without or with nanogels. The hybrid fibers consisted from pullulan-based nanogels and gelatin represented a new platform for stable encapsulation of other proteins, too.

10.3 Regenerative medicine 10.3.1 Hydrogels for burn injuries Pathophysiology of burn injury includes phenomena of necrosis, stasis, and hyperemia of affected zones which require special management because of hypoxia, oxidative stress, and inflammation processes. Over time, various methods of treating burns have been addressed, including injections of pluripotent stem cells into the injured area, but whose viability was quite low due to the rather hostile environment. To improve their viability and enhance proregenerative capacity, some researchers used biomimetic scaffolds that can be seeded with stem cells. Skin engineering appeals to biomaterials like alginate, hyaluronic acid, polyethylene glycol, collagen, elastin, and fibrin capable to mimic the extracellular matrix, mainly composed from type I collagen fibers arranged in a three-dimensional network, and its components, and finally to render the structural architecture of healthy skin. To design dermal synthetic or native substitutes/scaffolds for the replacement of injured skin, specialists used one of the following approaches/techniques, more or less expensive or complex: cryogenic processes, gas foaming, powder sintering electrospinning, salt-induced phase inversion, and leaching [40]. Barerra et al. fabricated a 5% collagen-pullulan hydrogel in order to seed it with adipose-derived stem cell (ASC). This composite was used to treat burn scars and scar contractures in case of a murine contact burn model. Crosslinking between pullulan (molecular weight 200,000) and rat tail collagen type I was performed with sodium trimetaphosphate (STMP) in the presence of NaOH and KCl (as porogen agent for in-gel crystallization). The hydrogel films were dehydrated with ethyl alcohol, dried overnight, washed in PBS until pH of 7, and stored at 4 C until further use [41]. Dehydrated hydrogels cut

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into circles were seeded capillary with solution of ACSs in sterile PBS, and then applied per wounds. Comparing the wounds treated with ASC-seeded hydrogel with untreated wounds, wounds treated only with hydrogel, and wounds treated only with ACSs, and using histopathological data, quantitative polymerase chain reaction and enzyme-linked immunosorbent assay tests, the authors concluded that the new treatment accelerated the reepitelization and vascularization of the injured zone, and improved the quality of collagen matrix. Cellularized skin substitutes are preferred for the treatment of chronic skin or burned wounds instead of acellular substitutes in order to speed up the healing process, to reduce the inflammation and to avoid the overinfection. Nicholas et al. prepared a skin substitute composed of a three-dimensional porous hybrid structure based on pullulan and gelatin, in which they incorporated in vitro human normal skin fibroblasts and human epithelial keratinocytes by centrifugation in the presence of laminin [42]. To create the pullulan-gelatin hydrogel, they used solvent casting-particulate leaching and freezedrying methodology. The hybrid hydrogel had an average pore size of 62 nm allowing the passage and in volume adhesion of the human dermal fibroblasts with diameter of 10e15 nm. The pullulan-gelatin hydrogels submersed in microbial pullulanase solution showed over 50% degradation after 20 min. Mechanical characteristics of pullulan-gelatin hydrogel affected their cellular properties. In case of using these new cellularized skin substitute, the neo-dermis thickness increased up to 204 mm, a double value compared to the use of acellular substitutes. In addition, the angiogenesis increased and the macrophage infiltration decreased.

10.3.2 Hydrogels with antiadhesion and tissue regeneration properties Tissue regeneration is dependent on the presence of an adequate amount of growth factors which synergistically act to improve the tissue regeneration and facilitate the formation of new nerve, periodontal tissue, bone, cartilage, and functional muscle [43,44]. In order to prevent postsurgical tissue adhesion, the specialists appeal to various pharmaceuticals like heparin, warfarin, steroids, urokinase, and direct thrombin inhibitors or adhesion prevention barriers in the form of films, gels, and solutions. Because the products currently exist on the specialized market tend to aggregate or fluidize, are nonbiodegradable and difficult to apply, Bang et al. proposed a new injectable antiadhesion barrier based on carboxymethyl cellulose CMC and pullulan [13]. In this sense, first they have modified the carboxymethyl cellulose with tyramine by a coupling reaction in the presence of EDC/NHS system. The modified carboxymethyl cellulose sponge was stored until its mixing with pullulan solution in order to prepare the hybrid hydrogel. The crosslinking reaction between the two polysaccharides was enzymatically mediated using different proportions of horseradish peroxidase (HRP) and hydrogen peroxide (H2O2). Generally, the gelation time of new product decreased as the amount of horseradish peroxidase increased. For example, the new hydrogels gelated in 30e35 s when horseradish peroxidase was settled at 223 units per 10 g of product regardless of the amount of hydrogen peroxide. This value of gelation time will prevent the nozzle clogging during a surgical protocol. The optimized formulations have enhanced properties due to the adhesion of pullulan and injectability of tyraminemodified carboxymethyl cellulose. Compared with general hydrogels, the storage modulus G0 of the modified CMCpullulan hydrogel was greater (about 60 kPa). As the concentration of horseradish peroxidase increased, the hybrid hydrogels degraded slower. Also, the weight loss of hydrogels increased with the increase of amount of hydrogen peroxide. In addition, the hydrogel biodegradation was not affected by the amount of horseradish peroxidase, but with 10 mL/10 g of hydrogen peroxide, the hydrogel starts to change its structure on the seventh day until on the 28th day it completely collapsed. According to the tests of biocompatibility and proliferation of mouse embryonic fibroblasts in the presence of new hybrid hydrogel, the cell viability was around 80%e82% and the material was nontoxic, which recommends it to be used in normal and laparoscopy surgery protocols as an antiadhesive barrier. Some biocompatible and biodegradable polymers themselves exhibit tissue regenerating properties, so they do not need to be loaded with drugs. An and the team prepared hydrogels by physical mixing in different proportions of carboxymethyl cellulose and pullulan, respectively [45]. Resulted solutions, labeled as ST 2.25-0, ST 2.5-5, ST 2.5-10, were designed to prevent the surgical intraoperative and postoperative complications generated by tissue adhesion. Both biopolymers are soluble in water and proper to manufacture safe, nontoxic, and hemocompatible antiadhesion agents. The rheological analysis, hemolysis assay, and adhesion force tests have been completed with the assessment of therapeutic efficacy of the new hydrogels. The values of the storage modulus G0 and loss modulus G00 at both 25 C and 37 C were higher as the concentration of pullulan and carboxymethyl cellulose increased. Among all antiadhesion agents, these ones with the highest concentration of carboxymethyl cellulose and pullulan, namely ST with 2.5:10 (w/v) ratio, recorded the highest adhesive values under 100 and 300 gf adhesive forces. To check their in vitro residence stability in physiological

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conditions, the adhesion agents were preserved in solutions of PBS, with or without the addition of 20% fetal bovine serum (FBS) for 35 days. Sample containing only carboxymethyl cellulose ST 2.25e0 has been almost fully eroded within 10 days in PBS containing 20% FBS solution and 20 days in PBS, while sample with maximum of pullulan ST 2.5e10 started to degrade after 15 days in PBS solution, highlighting the controlling role of pullulan in the hydrogel degradation profile. The survival rate of tested fibroblast and epithelial cell lines in the presence of the hydrogels was higher than 90% as toxicological tests revealed. Compared with Medicurtain, a standard antiadhesion agent used in surgical protocols, the peritoneal administration of carboxymethyl cellulose-pullulan hydrogels to lab rat’s models with injured peritoneum and cecum produced a healing of abdominal walls, recording the zero adhesion score. Type-1 diabetes is a chronic autoimmune disease in which the functioning of insulin secreting beta cells from pancreas is disturbed. Often insulin therapy is a commonly used approach in the clinical treatment of diabetics, and rarely, from economical reasons, the pancreas/islet transplantation is proposed. Usually, the immunosuppressive drugs used in transplant procedure negatively influence the function of transplanted beta cells. In addition, almost 30e50% of viable pancreas islet transplanted from one patient to another not survived. So, these disadvantages could be offset by the use of MIN6 cell line consisted of beta cell from a mouse insulinoma. To increase the viability in the culture medium and to improve the function of these cells in physiological conditions, some researchers proposed a nanothin coating of insulin secreting beta cell aggregates with various advanced functional biomaterials. Recently, Bal and team prepared nanogels composed from pullulan chain as backbone, cholesterol group, and acrylate group as side chains, which self-assembled due to hydrophobic cholesterol [46]. Its precursor, cholesteryl pullulan is an immunologically inert material. Then, during the incubation in aqueous physiologically compatible solution, the MIN6 pseudoislets were covered with multilayers of hydrogels in such manner crosslinked interactions between acrylate groups of hydrogels and thiols of proteins from cell membrane resulted. The viability tests of covered cells were so encouraging that they were considered by transplant specialists. It seems that this new engineering approach has a potential in design of more tolerable grafts.

10.3.3 Hydrogels for cartilage tissue engineering Because of poor mechanical strength and their fast degradation rate, the pullulan hydrogels are not proper for articular cartilage repair. Consequently, some authors fabricated biocompatible hybrid hydrogels by simply mixing solutions of methacrylated pullulan PulMA with degree of methylation of 10%, poly(ethylene glycol)diacrylate PEGDA with Mn ¼ 700 Da in various proportions, and lithium phenyl(2,4,6-trimethylbenzoyl)phosphinate (LAP) as UV photoinitiator such that to obtain photocrosslinking PulMA/PEGDA hydrogels (Fig. 10.6). They observed that with the addition of PEGDA up to 15%, the hybrid hydrogels had pore size range of 200e440 mm with well-interconnected three-dimensional porous structure and increased wall thickness, an increased storage modulus G’ of 16.0
103 Pa, and an increased compressive modulus of 1.17  0.17 MPa. Beside elastic behavior and good mechanical properties, these hydrogels displayed a decreased swelling ratio of 404  21% and a slower degradation rate, properties which promote them as promising scaffolds for cartilage repair and regeneration. In this regard, the rabbit’s mesenchymal stem cells (MSCs) encapsulated in PulMA/PEGDA hydrogels with 15% PEGDA content remained viable in proportion of 85% after three days. Also, the production of glycosaminoglycans GAG significantly increased, suggesting that MSCs were able to produce cartilage extracellular matrix within hybrid hydrogel [47].

10.3.4 Hydrogels for dentin tissue engineering To restore and regenerate the dental mineralized tissues destroyed by caries, some researchers used 3D hybrid scaffolds composites like poly-caprolactone/bioactive borate glass/Pluronic F127 hydrogel, collagen hydrogel crosslinked with cinnamaldehyde, or nanofibrous gelatin/magnesium phosphate [48e50]. Processing the results of its predecessors on composite scaffolds based on biopolymeric hydrogels and inorganic nanoparticles with applications in regenerative endodontics, Rad et al. [1] prepared a novel nanobiocomposite scaffold for the regeneration of dentin, following few steps. The mixture between solutions of cellulose acetate (30 kDa and degree of substitution of 2.49), pullulan oxidized with sodium iodate and gelatin in acetic acid at 40 C was cooled allowing crosslinking of pullulan with gelatin, and then phase separation. The lyophilized resulted product was redissolved in acetic acid/water at 40 C and mixed with KCl and boron (B) modified bioactive glass nanoparticles BG-NPs and recooling. In this stage, resulted scaffolds were immersed in glutaraldehyde solution for crosslinking. Unreacted glutaraldehyde was neutralized with glycine solution. The new scaffolds presented an aligned tubular morphology, with tube diameters of 11 mm, and were conditioned with solutions that mimic body fluids and allow the mineralization with calcium phosphate deposits. The water-absorption capacity of scaffolds during one month of incubation in phosphate buffer solution decreased in time and was in accordance with their

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FIGURE 10.6 Scheme of pullulan methacrylation by the transesterification reaction of the MA to the hydroxyl groups of the pullulan chains under pH 8, and PulMA/PEGDA hybrid hydrogel formation by UV photocrosslinking (methacrylated pullulan PulMA; methacrylic anhydride MA; poly(ethylene glycol) diacrylate PEGDA; lithium phenyl(2,4,6-trimethylbenzoyl)phosphinate LAP). Reprinted from Qin X, He R, Chen H, et al. Methacrylated pullulan/ polyethylene (glycol) diacrylate composite hydrogel for cartilage tissue engineering. J Biomater Sci Polym Ed 2021;32:1057e71, Copyright (2021), with permission from Taylor & Francis Ltd., www.tandfonline.com.

weight loss. In application stage, the new scaffolds were seeded with human dental pulp stem cells which subsequently differentiated into cells characteristic of dentin. The new biomaterial tested for dentin tissue engineering allowed the preparation of dentin constructs with superior biological properties confirmed by elevated levels of odontoblastic markers and mineral deposition.

10.4 Controlled and sustained release of drugs in regenerative medicine The sustained release of drugs from pullulan-based hydrogels was exploited for regeneration of bones, muscles, and vascular endothelium. Bone injuries or fractures associated with chronic pain are long-term disabilities which require specific treatments like systemic delivery of antiinflammatory agents and local release of growth factors to prevent inflammation, irritation, implant rejection, or bone loss, and stimulate a fasten bone healing. Specialty studies have been directed toward developing new biomaterials for tissues regeneration applications, with limited side effects, high stability, and prolonged half-time to control and sustain the local release of therapeutic agents. In bone/muscle/endothelial restoration therapy, the growth factors like prostaglandin E1, platelet-derived growth factor (PDGF-BB), human bone morphogenetic protein 2 (BMP2), human fibroblast growth factor 18 (FGF18), fibroblast growth factor (FGF), and vascular endothelial growth factor (VEGF) are administered individually or in combination of two to induce a synergistic effect [5]. Tissue engineering can also be applied to resolve bone defects. For this purpose, various materials in the form of hydrogel, loaded with growth factors are used that will favor osteogenesis. The encapsulation and release of growth factors from the hydrogel matrix is safer and with fewer side effects than drug systemic administration.

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Beside implants, the dental rehabilitation appeals to guided tissue regeneration procedure using materials for horizontal and vertical augmentation of defective bone. In this sense, over times the specialists used diverse biocompatible and biodegradable materials in the form of resorbable or nonsorbable membranes based on collagen, alginate, chitin-chitosan, methyl cellulose, poly(glycolic) and poly(lactic) acids, and poly(tetrafluoroethylene). There is a tendency to replace the membranes with hydrogels for bone regeneration. The CHP is a base for the design of such hydrogels. This hydrophobized polysaccharide has the abilities to self-assemble forming stable and monodisperse hydrogel nanoparticles and to trap inside it both hydrophilic molecules and hydrophobic molecules. For example, during the substitution process of the cholesterol groups, the CHP succeeds to incorporate growth factors and performs a molecular chaperone role. Miyahara et al. used rat calvaria with bone defect to evaluate the performance of two types of guided tissue regeneration membranes [51]. The first membrane was prepared by the combination of two types of bovine collagen derived from dermis tissue and tendon, respectively, and hexamethylene diisocyanate (HMDIC) as crosslinking agent. The second membrane was a nanogel-crosslinking hydrogel composed from CHPOA, resulted from the reaction of CHP and 2-AOI, and subsequently crosslinked with a poly(ethylene glycol) derivative bearing thiol groups PEGSH. The radiographic analysis and histological evaluation of calvaria samples treated with these biosorbable membranes were completed with the immunological tests. The concentration of the platelet-derived growth factor (PDGF) after the incubation of a test human serum with each membrane was monitored. The concentration of dimer PDGF-BB was lowest in case of incubation with CHP nanogel membrane suggesting that the endogenous protein was stored/trapped in material. In this way, CHP nanogel membrane alone stimulated bone regeneration and a smooth newly formed mature bone was observed in almost the entire defected area of calvaria under the membrane. In case of using collagen-based membrane, the surface of new bone was irregular. Preliminary results have been encouraging, but to recommend CHP nanogel membranes in clinical practice, researchers should optimize the storage of the material for a long time at room temperature and improve its mechanical properties by combining it with another biodegradable polymer. In another study, Fujioka-Kobayashi et al. used the same fast-degradable CHPOA nanogel-crosslinking hydrogel, obtained by the reaction of CHPOA nanogel with PEGSH crosslinker, to create delivery systems for another growth factors used for in vivo bone defect restoration [52]. The concentration of CHPOA nanogel in the resulted hydrogels was 20 mg/ mL. Each sample of hydrogel was individually loaded with recombinant BMP2, recombinant FGF18, and a 1:1 (w:w) mixture of the two growth factors. By disintegration of nanogels from hydrogel matrix, the growth factors were sustained released, followed by their interaction with plasma proteins like BSA. The complexes between growth factors and plasma proteins prolonged the action and maintained the concentration of growth factors in zone of bone defects, thus contributing to the bone healing. Because no one of protein from FGF family was effective alone, the researchers decide to test the combination of FGF18 with BMP2 to evaluate the osteogenetic effects in a calvarial bone defect. These growth factors cooperatively promoted the in vivo bone formation during three weeks. Actually, the FGF18 enhanced the osteo-inductive activity of BMP2 suggesting a synergistic effect on bone healing in vivo. Using bottom-up method, Charoenlarp and the team synthesized different hydrogels in the form of disc and based on PEGSH-crosslinked CHPOA nanogels, which they conditioned in different ways: conventionally drying; freeze drying; RGD peptide adding; and freeze-drying treatment. The resulted hydrogels were labeled as NanoClik disc, NanoCliP disc, and RGD-NanoCliP disc, respectively [5]. The freeze-drying procedure induced a porous structure into hydrogels. These hydrogels with different morphology were loaded with human fibroblast growth factor 18 (hFGF18) and used as support materials for bone-healing experiments. The best result for mouse calvarial defects was obtained with RGD-NanoCliP disc because the RGD peptide improved the cellular adhesion of the hFGF18 and kept it for a prolonged time to the defect site. Thus, the two molecules worked synergistically for bone restoration. The CHPOA nanogels crosslinked by pentaerythritol tetra (mercaptoethyl) poly(oxyethylene) also were used to design injectable NanoClik nanoparticles as carriers for W9 synthetic peptide [53]. This peptide drug prevents the bone loss by the inhibition of osteoclastogenesis. The reference material was CHP nanogels loaded with W9. The W9 peptide integrated and subsequently released from CHP nanogels subcutaneous injected to a lot of laboratory animals fed with a low calcium diet had a little effect on bone loss. In case of NanoClik-W9 system, a better W9 sustained release and an inhibitory effect of peptide on bone loss were recorded. These results derived from the fact that the unloaded NanoClik nanoparticles remained in circulation longer time compared with unloaded CHP nanogels, meaning that they were more stable in vivo. Beside antioxidant and antiinflammatory corticosteroid properties, dexamethasone is an osteogenic inducer, too. Chauhan et al. [54] developed a pullulan-based biodegradable hydrogel loaded with dexamethasone for potential application in bone regeneration. First, PEG-dexamethasone conjugates with different content of drug (0.50 mmol or 0.050 mmol) were prepared by reaction of 8-arm PEG hydrazine (10% w/v) with dexamethasone in methanol with the addition of acetic acid. The mixture was precipitated with diethyl ether, dried, purified, and lyophilized until its

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FIGURE 10.7 Fabrication of injectable, self-healing, and biodegradable pullulan-PEG hydrogels via hydrazone linkages for the sustained release of dexamethasone covalently attached to the PEG in the hydrogel matrix. Reprinted from Chauhan N, Gupta P, Arora L, Pal D, Singh Y. Dexamethasoneloaded, injectable pullulan-poly(ethylene glycol) hydrogels for bone tissue regeneration in chronic inflammatory conditions. Materi Sci Eng C 2021;130:112463, Copyright (2021), with permission from Elsevier.

incorporation in hydrogels. Second, crosslinking of OP (1% w/v) in the presence of 8-arm PEG hydrazine-dexamethasone conjugates resulted in the formation of injectable polymeric hydrogels based on hydrazone linkages at room temperature and pH 6.5 (Fig. 10.7). Dex-loaded hydrogels and free-drug hydrogels recorded a highly macroporous and interconnected structure, with pores with diameter of 45e48 mm (Fig. 10.8). The physiological (7.4) and wound (6.5) pH were tested, and authors concluded that the release process of PEG-Dex conjugates conformed/fitted to different kinetic models in accordance with pH and drug concentration: KorsmeyerePeppas exponential model, and consequently a quasi Fickian mode of release of PEG-Dex conjugates, at 6.5 pH for both type composition of hydrogels; Higuchi and KorsmeyerePeppas models at 7.4 pH for hydrogel with low content of drug; and Higuchi model at 7.4 pH for hydrogel with high content of drug. To explore the degradation properties of hydrogels, some measurements of weights for hydrogel preserved at different pH were realized. Compared with other chemically crosslinked hydrogels, these hydrogels were intact and stable, even after 30 days, necessary conditions for tissue engineering applications. The rheological studies revealed that hydrogels possess more elastic-like nature as the storage modulus was higher than the loss modulus. Also, hydrogels were mechanically robust because of a complete recovery of initial shape/status after six

FIGURE 10.8 Morphology of lyophilized pullulan-PEG hydrogels (SEM images; Scale bar 50 and 10 mm). Reprinted from Chauhan N, Gupta P, Arora L, Pal D, Singh Y. Dexamethasone-loaded, injectable pullulan-poly(ethylene glycol) hydrogels for bone tissue regeneration in chronic inflammatory conditions. Materi Sci Eng C 2021;130:112463, Copyright (2021), with permission from Elsevier.

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alternate cycles of extreme (400%) and low (1%) strains of 60 s each. In this case, the self-healing properties of hydrogels were due to the dynamic hydrazone linkages. In vitro biological tests suggested that dexamethasone-loaded hydrogels promoted early osteogenic differentiation, enhanced mineralization, and improved the vitality and proliferation of specific cells such that could be recommended as potential scaffolding material for bone implants. In the soft tissue engineering, other growth factors come into play, but similar hydrogel carriers could be used. Cutiongco et al. tested some scaffolds as platforms for fibroblast cell attachment and for controlled release of VEGF [55]. The plain scaffold consisted from pullulan-dextran hydrogel. The second tridimensional scaffold was a composite between the pullulan-dextran hydrogel and the interfacial polyelectrolyte complexation IPC fibers produced from chitosan and alginate. The spatial distribution of IPC fibers in composite controlled the future distribution of incorporated cells. The incorporation of fibronectin improved the cell adhesion and proliferation on composite scaffold. In addition, using different proportions of alginate and heparin, the VEGF incorporation efficiency was optimized. The VEGF molecules were sustained released from composite scaffold over seven days, preserving its biological activity even after leaving the matrix. The experimental results suggesting that a combination between extracellular matrix proteins and growth factors in interfacial polyelectrolyte complexation IPC type fibers could be favorable for in vitro coculture composed from endothelial cells and for the mimicking of a tridimensional construct. Using a fixed concentration of silk fibroin and variable concentrations of pullulan derivative of 0, 1.5, 3, 6, and 12 mg/ mL, Li and the team developed biocompatible injectable in situ-forming hydrogels crosslinked in the presence of horseradish peroxidase and hydrogen peroxide [56]. In this way, the physicochemical properties of the resulted hydrogels were modulated. The new hydrogels were proposed as scaffolds for multipotent MSC with potential applications in musculoskeletal tissue engineering. The tyramine-substituted carboxymethylated pullulan TA-CMPL was synthesized by a coupling reaction between carboxylic groups of polysaccharide and the amine groups of the low molecular compound using EDC-NHS activation system. The gelation time for silk fibroin hydrogels, pullulan derivative hydrogels, and hybrid hydrogels with 12 mg/mL TACMPL were 1 h, 6, and 12 min, respectively. With the increase of TA-CMPL concentration in system, the network crosslinking density increased and the equilibrium water uptake decreased. The values of compressive modulus of the hydrogels varied from 7 to 71 kPa, while the storage modulus ranged from 200 to 1470 Pa. All rheological and mechanical characteristics of hydrogels fall in the range of that of musculoskeletal system. After seven days, the rabbit MSCs seeded on the hybrid hydrogels were homogeneous dispersed and indicated a chondrogenic differentiation. CHP with 1.2 substitution degree, synthesized starting from pullulan with molecular weight of 1
105 g/mol was the start material for the preparation of nanogel-crosslinked porous NanoCliP gel [9]. In the first step, to CHP solution in DMSO, 4-(4,6-dimethoxy-1,3,5-triazin-2-yl)-4-methylmorpholinium chloride (DMT-MM) was added, followed by the addition of N, N-diisopropylethylamine DIPEA as catalyst and acrylic acid. After mixture stirring for 22 h and purification for four days, the acryloyl-cholesteryl-bearing pullulan (CHPA) with 9.3 substitution degree of acryloyl groups per 100 glucoside units resulted. In the second step, solutions of CHPA and four-branched poly(ethylene glycol) with terminal thiol groups PEGSH in PBS were mixed and incubated for 1 h at 37 C. The nanogel-crosslinked porous NanoCliP gel, with microsize pores and resulted by Michael addition reaction mechanism, was gradually frozen to increase its porosity of gel from 46% to 72% (Fig. 10.9). The freeze-dried nanogel-crosslinked porous FD NanoCliP gel was loaded with a model protein, namely fluorescein isothiocyanate (FITC)-labeled human insulin. The release behavior of FTIC-insulin from nanogels was investigated in PBS with and without 10% FBS. It was found that the encapsulated protein was gradually released from FD NanoCliP gel in proportion of 80%, while 20% already was removed in the preexperimental washing step. NanoCliP gel and FD NanoCliP gel were degraded in 10 days in PBS with 10% FBS adding at 37 C, but degradation has occurred in 20 days for FD NanoCliP gel immersed in PBS. 3D structure of FD NanoCliP gel allowed the myoblasts to penetrate and proliferate into the interconnected macropores of gel to a depth of at least 60 mm and not prevented the cell differentiation. These in vitro results encouraged the researchers to apply this hydrogel in vivo. Thus, they used the hydrogel as scaffolding material in the tongue surgery. Compared with the empty gel, 3D muscle regeneration of mouse tongue was improved if the FD NanoCliP gel implanted in the surgical hole contained myoblasts. This has been highlighted by the increase in the number of newly regenerated myofibers. Thereby, the studied hydrogel had two distinct roles: biocompatible vehicle for sustained release of protein and suitable scaffold for myoblasts, which recommended using it in drug delivery or regenerative medicine.

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FIGURE 10.9 Preparation of nanogel-crosslinked porous gel NanoCliP (CHPA ¼ acryloyl-cholesteryl-bearing pullulan; PEGSH ¼ four-branched poly(ethylene glycol) with terminal thiol groups). Reprinted from Kinoshita N, Sasaki Y, Marukawa E, Hirose R, Sawada S-I, Harada H, et al. Crosslinked nanogel-based porous hydrogel as a functional scaffold for tongue muscle regeneration. J Biomat Science Polym Ed 2020;31:1254e71, Copyright (2020), with permission from Taylor & Francis Ltd., www.tandfonline.com.

10.5 Conclusions and perspective remarks This polysaccharide continues to be exploited because it is economically accessible and has special properties. The mechanical properties, biocompatibility, biodegradation, and cytotoxicity of pullulan are improved by chemical modification. Especially the C6 site from the chemical structure of pullulan is susceptible to chemical modification. Pullulan derivatives can self-associate, and their amphiphilic character favors the loading of different drugs in such carriers. The release of drug from pullulan-based hydrogels could be stimuli-responsive. If the drug is loaded and then released in a controlled manner at the target site, the dose of the drug is reduced than if it had been injected as such, thus avoiding toxicity, which is even greater with systemic/parenteral administration. The coadministration of drugs aimed the “as low possible drug dose and side effects” criteria. After the drug release, the polymeric matrix is biodegraded. Sometimes, the pullulan-based hydrogels play a double role: carrier for drugs and scaffold for cells. Thus, the tridimensional hydrogel networks and stem cells are involved in the newest techniques of tissues regeneration. Whether used as drug carriers or as scaffolds for cells, the pullulan-based hydrogels require additional biological studies to be subsequently recommended for human subjects.

Acknowledgments Thanks to all those who did not believe in me and consequently motivated me to move on.

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[35] Muraoka D, Harada N, Hayashi T, et al. Nanogel-based immunologically stealth vaccine targets macrophages in the medulla of lymph node and induces potent antitumor immunity. ACS Nano 2014;8:9209e18. [36] Tahara Y, Kosuge S, Sawada S, et al. Nanogel bottom-up gel biomaterials for protein delivery: Photopolymerization of an acryloyl-modified polysaccharide nanogel macromonomer. React Funct Polym 2013;73:958e64. [37] Shimoda A, Sawada S, Kano A, et al. Dual crosslinked hydrogel nanoparticles by nanogel bottom-up method for sustained-release delivery. Colloids Surf B Biointerfaces 2012;99:38e44. [38] Morimoto N, Hirano S, Takahashi H, et al. Self-assembled pH-sensitive cholesteryl Pullulan nanogel as a protein delivery vehicle. Biomacromol 2013;14:56e63. [39] Shimoda A, Chen Y, Akiyoshi K. Nanogel containing electrospun nanofibers as a platform for stable loading of proteins. RSC Adv 2016;6:40811e7. [40] Wong VW, Rustad KC, Galvez MG, et al. Engineered pullulan-collagen composite dermal hydrogels improve early cutaneous wound healing. Tissue Eng 2011;17:631. [41] Barrera JA, Trotsyuk AA, Maan ZN, et al. Adipose-derived stromal cells seeded in pullulan-collagen hydrogels improve healing in murine burns. Tissue Eng Part A 2021;27:844e56. [42] Nicholas MN, Jeschke MG, Amini-Nik S. Cellularized bilayer pullulan-gelatin hydrogel for skin regeneration. Tissue Eng Part A 2016;22:754e64. [43] Chen F-M, Zhang M, Wu Z-F. Toward delivery of multiple growth factors in tissue engineering. Biomaterials 2010;31:6279e308. [44] Lee K, Silva EA, Mooney DJ. Growth factor delivery-based tissue engineering: general approaches and a review of recent developments. J R Soc Interface 2011;8:153e70. [45] An JM, Shahriar SMS, Hasan MN, et al. Carboxymethyl cellulose, Pluronic, and pullulan-based compositions efficiently enhance antiadhesion and tissue regeneration properties without using any drug molecules. ACS Appl Mater Interfaces 2021;13:15992e6006. [46] Bal T, Oran DC, Sasaki Y, et al. Sequential coating of insulin secreting beta cells within multilayers of polysaccharide nanogels. Macromol Biosci 2018;18:1800001. [47] Qin X, He R, Chen H, et al. Methacrylated pullulan/polyethylene (glycol) diacrylate composite hydrogel for cartilage tissue engineering. J Biomater Sci Polym Ed 2021;32:1057e71. [48] Kolan K, Liua Y, Baldridge J, Murphy C, Semon J, Day D, et al. Solvent based 3D printing of biopolymer/bioactive glass composite and hydrogel for tissue engineering applications. Procedia CIRP 2017;65:38e43. [49] Kwon Y, Lee S, Hwang Y, Rosa V, Lee K, Min K. Behavior of human dental pulp cells cultured in a collagen hydrogel scaffold cross-linked with cinnamaldehyde. Int Endod J 2017;50:58e66. [50] Qu T, Jing J, Jiang Y, Taylor RJ, Feng JQ, Geiger B, et al. Magnesium-containing nanostructured hybrid scaffolds for enhanced dentin regeneration. Tissue Eng Part A 2014;20:2422e33. [51] Miyahara T, Nyan M, Shimoda A, et al. Exploitation of a novel polysaccharide nanogel cross-linking membrane for guided bone regeneration (GBR). J Tissue Eng Regen Med 2012;6:666e72. [52] Fujioka-Kobayashi M, Ota MS, Shimoda A, et al. Cholesteryl group- and acryloyl group-bearing pullulan nanogel to deliver BMP2 and FGF18 for bone tissue engineering. Biomaterials 2012;33:7613e20. [53] Sato T, Alles N, Khan M, et al. Nanogel-crosslinked nanoparticles increase the inhibitory effects of W9 synthetic peptide on bone loss in a murine bone resorption model. Int J Nanomed 2015;10:3459e73. [54] Chauhan N, Gupta P, Arora L, Pal D, Singh Y. Dexamethasone-loaded, injectable pullulan-poly(ethylene glycol) hydrogels for bone tissue regeneration in chronic inflammatory conditions. Materi Sci Eng C 2021;130:112463. [55] Cutiongco MFA, Tan MH, Ng MYK, et al. Composite pullulan-dextran polysaccharide scaffold with interfacial polyelectrolyte complexation fibers: a platform with enhanced cell interaction and spatial distribution. Acta Biomater 2014;10:4410e8. [56] Li T, Song X, Weng C, et al. Enzymatically crosslinked and mechanically tunable silk fibroin/pullulan hydrogels for mesenchymal stem cells delivery. Int J Biol Macromol 2018;115:300e7.

Chapter 11

Hydrogels based on levan A´lvaro Gonza´lez-Garcinun˜o1, Antonio Tabernero1 and Eva M. Martı´n del Valle1, 2 1

Department of Chemical Engineering, University of Salamanca, Salamanca, Spain; 2Institute of Biomedicine of Salamanca, IBSAL, Salamanca,

Spain

11.1 Introduction The use of biopolymers in hydrogel production is becoming more important due to their biocompatibility and biodegradability properties. The most commonly used biopolymers, such as glycogen, hyaluronic acid, chitosan, dextrans, or gellan gum, are reviewed in other chapters of this book. However, there are new polysaccharides with interesting properties that have not been studied in depth yet, such as schizophyllan or scleroglucan. Levan is one of these newly found polymers, a homopolysaccharide formed by fructose residues that is produced by different bacteria genera and some plants. The growing interest of the academic community in this polymer over the last decade is evidenced by the number of scientific publications that mention this polymer in their abstract, title, or keywords. The number of papers related to levan was 19 in 2010, and reached 87 in 2020 (data obtained from Scopus), underlining the interest of the scientific community in this promising tool and its various biomedical applications. This chapter discusses the properties of this polymer, the main mechanisms for levan production, and its potential applications in the biomedicine field. The chapter will focus on the latest research on the synthesis of levan-based hydrogels, their properties, and their applications.

11.1.1 Levan structure and mechanism of formation Levan is a homopolysaccharide composed of fructose residues linked by b(2e6) bonds with occasional b(2e1) linked branch chains. Levan is also commonly known as fructan, but this is a more general term involving other molecules such as inulin. The name levan was postulated by Greig-Smith in 1901 due to its levorotatory properties in polarized light. Levan is obtained from sucrose, using levansucrase as the catalytic enzyme. The mechanism of action of levansucrase is still under study, but it seems to consist of a two-step reaction: hydrolysis and transfructosylation, with the substrate usually being sucrose. First, the molecule of sucrose is hydrolyzed into glucose and fructose and, subsequently, the enzyme polymerizes the residues of fructose using the previously obtained molecule of glucose as a primer. Therefore, levan chains always start with a glucose residue linked to a variable number of fructose residues. The size of the levan molecule depends on the source, the enzyme used for synthesis, and the environmental conditions, ranging from the lowest molecular weight (one residue of glucose and two residues of fructose, known as 6-kestose) to the highest molecular weight (around 104 kDa) [1]. Fig. 11.1 shows the glucose primer and the b(2e1) branches of a levan molecule. The mechanism of catalysis involves three levansucrase residues, following a ping-pong mechanism according to the model proposed by Chambert et al. [2,3]. This model implies that the first substrate (sucrose) changes the conformational structure, which induces the reaction with the second substrate (fructose). This hypothesis is based on the formation of an intermediate between the enzyme and a fructosyl residue, which reduces the activation energy and, in turn, allows the incorporation of another fructosyl residue. Therefore, the reaction occurs as follows: there are two sucrose molecules in the active center, where one of them acts as a fructosyl donor and the other as an acceptor (after hydrolysis). The nucleophilic attack can happen through oxygen one or oxygen six, forming 1-kestose or 6-kestose, respectively. When 6-kestose is formed, another fructosyl residue can start the attack and continue with the polymerization process. Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine. https://doi.org/10.1016/B978-0-323-95351-1.00012-0 Copyright © 2024 Elsevier Inc. All rights reserved.

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FIGURE 11.1 Levan chemical structure.

In this mechanism, there are main three amino acids in catalysis: glutamic acid at position 342, aspartic acid at position 86, and aspartic acid at position 247. The latter amino acid leads to the transitional state formation (enzyme-fructosyl), allowing the polymerization to continue [4]. A mutagenesis study considering six different mutations has been published. Nevertheless, the wild strains still give the best yield in production. Site-directed mutagenesis modifies the chain length, obtaining a wide range of molecular weights [5].

11.1.2 Levan properties and associated potential applications A wide range of applications has been reported in the last years for this biopolymer [6], most of them related to its physicochemical properties. The three most interesting properties of levan (solubility, tensile strength, and viscosity) have been summarized below: -

-

-

Solubility: Levan is soluble in water and some organic solvents such as dimethylsulfoxide (DMSO). It is insoluble in the vast majority of organic solvents such as methanol, 2-propanol, ethanol (used for its precipitation), or dimethyl formamide. To be precise, this is not a solution process because levan chains are not solvated by water molecules. Instead, there is a phenomenon of chain self-assembly due to thermodynamic equilibrium. Above critical concentration, known as CAC (critical aggregation concentration), levan chains are able to rearrange themselves to expose their hydrophilic residues to the solvent (water), while the hydrophobic ones avoid contact with water. This results in the formation of a colloid with a variable number of chains [7] and size, depending on different factors. This organization in solution has been used several times to create drug delivery systems to anchor different hydrophobic drugs within them. Moreover, levan presents no swelling in water, and, therefore, the hydration process of the polymer can be considered negligible [8]. Tensile strength: The polymer structure includes branches (which cause cohesive forces) and hydroxyl groups (which can interact with different molecules). Due to this structure, levan presents high tensile strength in aluminum (more than 1500 psi) and adequate shear strength in plastics (sometimes higher than some petroleum-based composites) [9]. Viscosity: This polymer does not show the usual properties of polysaccharides in terms of viscosity. It has a remarkably low intrinsic viscosity compared to other similar high molecular weight polymers. This could be due to the dispersion that levan shows in water (instead of dissolution) as has been mentioned in the solubility section. Stivala and Bahary [10] reported an intrinsic viscosity from 0.1 to 0.2 dL/g depending on the solvent and temperature. The pH and ionic strength do not have a significant effect on the viscosity of the solution, whereas temperature plays an essential role above 70
C [11]. Levan solution shows plastic behavior at concentrations above 50% weight [8]. This behavior makes this polymer particularly valuable as it addresses one of the main drawbacks for the biomedical application of exopolysaccharides: the handling of high viscosity fluids. Other interesting properties of levan are biocompatibility and biodegradability. Moreover, this polymer does not cause ocular irritation and cannot be easily hydrolyzed. Taking advantage of these properties, levan has been suggested as a promising tool for different biomedical applications.

Hydrogels based on levan Chapter | 11

-

-

-

-

-

-

-

177

Antiinflammatory: Different studies have proved the potential use of levan as an antiinflammatory agent. First, Liu et al. [12] showed that a high concentration of levan (around 0.8 g/L) was required to observe this effect in splenocytes, whereas more recent studies have observed the same effect using lower concentrations (10 ppm), achieving a higher antiinflammatory effect compared to reference drugs such as diclofenac sodium [13]. In vitro results showed that levan increases the level of TNFa, promoting the mobilization of T and B cells as well as immunoglobulins [14]. Antitumor: The potential antitumor effect of levan is currently unclear, and there are only a few studies on this topic. Leibovici et al. [15] showed how levan induced the action of B-cells, controlling tumor proliferation. In 2000, Calazans et al. [16] showed how levan from Zymomonas mobilis can inhibit sarcoma cell proliferation. Four years later, Yoon et al. [17] established an association between this inhibition and the degree of branching by analyzing the effect on gastric and liver cells. Nowadays, research efforts focus on clarifying the role of levan in tumor cell proliferation [18]. Antioxidant: Levan has been proposed as an antioxidant as it promotes an increase in the glutathione-S-transferase expression, as well as an increase in the activity of catalase and superoxide dismutase enzymes. All these enzymes are involved in the conservation of the cell’s oxidative state [19,20]. Therefore, the use of levan could prevent different diseases related to oxidative stress such as atherosclerosis or hypercholesterolemia. Support for cell proliferation: Its good tensile strength and the absence of swelling make levan an interesting polymer to use as a scaffold for tissue engineering. However, this strategy has not been explored in detail, and only a few studies on this property have been published. For example, Erginer et al. [21] used sulfated levan from Halomonas as a scaffold for cardiac cells because of its anticoagulant properties (similar to heparin). Tabernero et al. [22] showed that levan can be introduced in membranes by using supercritical CO2, modifying the mechanical and biological properties of membranes. Prebiotic: Fructooligosaccharides (FOSs) were identified several years ago as prebiotics, that is, substances that promote the growth of bacteria populations in the human gut [23]. Levan biopolymers are structurally similar to FOS, with the main difference being the molecular weight, which is considerably higher due to the length of the chains. Some studies [24,25] have observed that levan also promotes bacteria growth, particularly, in some genera such as Faecalibacterium, Bifidobacterium, and Lactobacilli. The first one is considered the most beneficial for humans because of its antiinflammatory properties, while the second and third ones are the main genera promoted by classical prebiotics. However, to the best of our knowledge, the reason or mechanisms for these effects are still not clear. Film formation: Different film structures can be formed by mixing levan with other reagents. For example, Barone and Medunets [26] created a film with an adequate plasticity degree by mixing levan and glycerol, followed by an extrusion process. Chen et al. [27] mixed levan and montmorillonite, creating a film with thermal stability and increased glass transition temperature. Recently, Osman et al. [28] prepared a thermosensitive gel by modifying the polymer with N-isopropylacrylamide (NIPA) to control the release of 5 aminosalicylic acid (5ASA). Drug delivery systems: The use of levan as a carrier for drug delivery is an alternative that has been explored in recent years, motivated by the possibility of specific targeting of levan nanoparticles on fructose receptors (GLUT5) and their overexpression in some tumor cells such as liver, breast, or colorectal cells. Kim et al. [29] considered the use of levan for indocyanine green release to test this specificity in mice (in vivo studies). Levan can form nanoparticles because of its ability to self-assemble and shield its hydrophobic residues when a critical concentration is reached [30]. Different drugs have been anchored or encapsulated to levan particles for a controlled release: vancomycin [31], bovine serum albumin [32], 5-fluorouracil [33], silver nitrate [34], doxorubicin [35], resveratrol [36], or curcumin [37].

11.2 Production mechanisms Typically, levan has been produced from microbial fermentation by culturing different microorganisms that express the enzyme levansucrase and can, therefore, produce the polymer in a culture medium with excess sucrose. However, over the last years, different alternatives have been found for the production of levan. These methods are based on the overexpression and purification of this enzyme so that it can be used alone in a cell-free system, without the need for an entire microbial culture. This section summarizes different strategies for microbial levan production and cell-free synthesis, concluding with the method for levan purification in both types of production.

11.2.1 Microbial production For microbial production, it is essential for microorganisms to express the levansucrase enzyme. This enzyme is translated from mRNA into ribosomes and accumulates in the periplasm, where it acquires its final form. From this space, the enzyme is released into the external medium. There are two different types of enzyme secretion depending on the bacteria type.

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In gram-positive bacteria, secretion consists of two steps: first, the signal peptide is eliminated and then, the final folding takes place to obtain the functional structure [38]. In gram-negative bacteria, there is no final folding and the process is only controlled by the signal peptide. In both cases, the signal for enzyme release is an acidic environment for bacteria growth (pH between 5 and 6) [39]. The first description of levan production by microbial fermentation was presented in 1943, where Bacillus subtilis was used as the strain and sucrose as the substrate. Today, there are more than 80 papers that explore levan production using different microorganisms and substrates. Despite this wide range of culture conditions, there are some common parameters in all the studies: -

Stirring: from 100 to 200 rpm to ensure the correct mixing with no cell damage. Temperature: from 25 to 40
C Culture pH: from 5 to 7

When considering the culture time, several differences are observed among the reported studies: from 21 h [40] to 6 days [41]. Different levels of productivity have been reported in the studies, showing values of approximately 1 g/L$h. The highest yield was achieved with B. subtilis natto using 400 g/L sucrose as substrate, reaching a productivity of about 7 g/L$h [42].

11.2.2 Cell-free synthesis Recombinant techniques have allowed the overexpression of some enzymes in bacteria or yeast, and thus the possibility of using only the enzyme without microbial growth in a cell-free system. The study with the highest yield was carried out by Kang et al. [43], reaching more than 14,000 U/L of levansucrase from Pichia pastoris. Consequently, the first steps in levan synthesis (microorganism selection, inoculum culture, microbial growth, removal of cells from culture .) are not necessary. This fact leads to a reduction in operating costs as well as a reduction in some risks associated with microbial culture such as contamination with other strains, presence of undesired metabolites, etc. [44]. However, the process of obtaining a highly pure enzyme involves elevated costs, meaning that in most cases, the typical forms of production are still preferred. The alternative to reducing costs is an immobilization process to allow several production cycles with the same enzyme. González-Garcinuño et al. [45] explored this possibility with immobilization in an agarose porous monolith and alginate-calcium particles, reducing the limitations of mass transfer phenomena. Other support materials that have been studied are chitosan or DEAE-cellulose [46,47], which showed adequate yields of immobilization as well as a similar catalytic rate to nonimmobilized enzymes (increasing the stability of the enzyme in some cases). The design of a cell-free system for levan production only requires the presence of substrate (sucrose in most cases), the enzyme levansucrase, and a buffer to ensure pH stability and ionic strength. Szwengiel et al. [48] suggested that catalysis is enhanced by the presence of manganese ions (Mn2þ) in the medium, which acts as an enzyme cofactor, providing stability to the intermediate and promoting polymer chain extension.

11.2.3 Levan purification Levan purification depends on the method used for synthesis (microbial production or cell-free system). In both cases, the final step involves precipitation with ethanol, which is common and essential to obtain a polymer with a high degree of purity [49]. In microbial production, it is necessary to clear the culture, removing the bacteria that synthesized the polymer. To do this, deep filtration or centrifugation is a viable alternative, but centrifugation is usually preferred as it avoids cake formation and the pellet can be easily removed. After centrifugation, the levan in the supernatant is purified. The supernatant is treated with a solution of methanol in water (20% v/v) to promote the deactivation of other enzymes and to destroy possible bacteria that may have not been removed during the centrifugation process, obtaining a product without bacterial contamination. This mixture is dialyzed against water using an MWCO 500 Da membrane to remove all components of the culture medium [50]. Next, the precipitation process is indistinctive for both methods. Each volume of the content of the dialysis bag or the reaction medium (in the case of cell-free systems) is mixed with three volumes of ethanol 96% to promote polymer precipitation. In some cases, some amount of calcium chloride can be added to enhance precipitation or it can be kept frozen (20
C) overnight to obtain better yields [51]. Following precipitation, the supernatant is discarded after centrifugation and the pellet (purified polymer) is dried by lyophilization. Fig. 11.2 summarizes the purification process for both synthesis methods.

Hydrogels based on levan Chapter | 11

FIGURE 11.2

179

Methods for levan purification from microbial culture and cell-free system.

11.3 Levan-based hydrogels As explained above, due to its particular structure, levan is an uncharged water-soluble exopolysaccharide and is considered a nongelling material. This phenomenon is a consequence of the interactions that allow the polymer to rearrange into nanoparticles when dispersed in water. Moreover, this is also important when it comes to the rheology of levan solutions. The rheological behavior provides information on whether levan solutions show elastic or viscous behavior and whether they can be considered gels.

11.3.1 Rheology of levan solutions The rheology of the levan solutions depends on the polymeric source. However, some common results have been observed. In general, aqueous solutions of levan show a Newtonian behavior up to a high concentration. This phenomenon was found in solutions of levan produced from Bacillus sp. (Newtonian behavior up to 30 wt%) as is illustrated in Fig. 11.3 [52], from Erwinia amylovora (up to 8 wt%) [53] and from an enzymatic synthesis (up to 5 wt%) [54]. This is supported by Benigar et al. [55], who showed a similar Newtonian behavior of different aqueous solutions of levan (from Z. mobilis, Escherichia herbicola, and B. subtilis) at 1 wt%. However, they found that, at 8 wt%, levan from B. subtilis maintained Newtonian behavior, while levan from E. herbicola and Z. mobilis acquired pseudoplastic characteristics. Finally, aqueous solutions of levan from B. subtilis AF17 showed a shear-thinning effect over a wide range of concentrations (from 0.1 wt% to 2.0 wt%) [56]. Oscillatory analysis, although less explored, also gives contradictory results about the polymeric behavior of aqueous solutions of levan. Particularly, Benigar et al. [55], found that at low concentrations, levan from B. subtilis, Z. mobilis, and E. herbicola had a strong elastic behavior (gel-like) since the storage modulus was always higher than the loss modulus, showing solutions with a predominant solid behavior. However, at high concentrations, in solutions of levan from E. herbicola and Z. mobilis, the loss modulus can become higher than the storage modulus (depending on the strain) and

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FIGURE 11.3 Viscosities of aqueous levan solutions from Bacillus sp. at different concentrations at 20
C. Reprinted from Arvidson SA, Rinehart T, Gadala-Maria F. Concentration regimes of solutions of levan polysaccharide from Bacillus sp. Carbohydr Polym 2006;65:144e9, Copyright (2006), with permission from Elsevier.

has a strong strain dependence. Finally, Blanco-López et al. [54] observed the typical spectrum for polymeric solutions of enzymatic levan (higher loss modulus than storage modulus), whereas Hundschell et al. [57] found both phenomena (solidlike and fluid-like behavior) depending on the size and polydispersity of the levan in aqueous solutions. These results can be explained by taking into account how the levan assembles into nanoparticles in water, as well as the interactions between the polymer chains. However, the concentration at which the polymer rearranges into nanoparticles depends on the microorganism used and the production methodology [58]. Therefore, experimental conditions and microorganisms play an important role in levan structure, particle size, and polydispersity, and as a consequence, in the rheological behavior [57]. In any case, although sometimes aqueous levan solutions can have a predominant elastic behavior, the difference between storage and viscous modulus is usually not high.

11.3.2 Type of hydrogels: A brief summary of classification and preparation techniques There are different categories of hydrogels depending on the type of crosslinking or the physical state of the hydrogel (solid, semisolid, or liquid). These characteristics are key when choosing the final application of the material. More information about the classification of hydrogels can be found in Refs. [59,60]. Hydrogels are essentially formed by crosslinked networks. Hydrogels can be crosslinked physically or chemically via different methodologies. Physical hydrogels are produced without using crosslinkers and are based on the interactions between the different chains. These hydrogels are usually reversible, and the experimental conditions and how the hydrogels are obtained (i.e., pH modification) can be used to tune the final properties. Technique such as freeze-thawing cycles [poly(vinyl alcohol)] is the typical example [61], stereocomplexation [poly(lactic acid)] as an example [62], or even the use of heat can be used to synthesize physical hydrogels [63]. On the other hand, chemically crosslinked hydrogels are characterized by the existence of covalent bonds, forming a strong network, that can be tuned by modifying gelation time, polymer functionalization, etc. Covalent bonds can be formed using different methodologies, such as the use of a crosslinker (glutaraldehyde [GA]), enzymatic reaction, radical polymerization, grafting, or even radiation [59,60]. Hydrogels can also be classified depending on their physical properties and their physical state. This classification is crucial to determine the final application. Solid hydrogels have a strong network and are usually produced by covalent crosslinks, being solid at room temperature and having an appropriate swelling capacity [64]. Semisolid hydrogels are

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usually made of a combination of polymers (one of them must be natural, such as plant gum) [65]. This combination, with a high molecular weight, gives important adhesive interactions and interfacial forces to the hydrogel. Finally, liquid hydrogels are liquid at room temperature. However, under certain conditions, they can experience an elastic phase, conferring useful properties to the material [66].

11.3.3 Levan-based hydrogels The synthesis of levan-based hydrogels depends on their polymeric structure and properties. Levan is an uncharged polysaccharide, so both physical and chemical network crosslinking are complex processes. For these reasons, levan is considered a nongelling polysaccharide. In spite of this fact, several strategies can be followed to produce levan-based hydrogels. Table 11.1 summarizes the results of levan hydrogels. This table was built after searching in Scopus with the words levan þ hydrogel (only 14 documents were found). After using different crosslinkers, Demirci et al. [67] showed that only 1,4-butanediol diglycidyl ether (BDE) was able to crosslink the levan polymeric chains, forming a hydrogel with a strong solid-like behavior and a large storage modulus. In this context, Fig. 11.4 illustrates a dried levan hydrogel that can be obtained by using this crosslinker. However, to date, the most common method to produce levan-based hydrogels is blending it with another polymer to promote interactions between the different polymeric chains and/or adding another compound that can be crosslinked with another polymer, forming a new network. As an example, levan was blended with gellan gum [69] under alkaline conditions. These conditions promoted an increase in functional groups, mainly from gellan gum, which could be then crosslinked with calcium chloride. A similar strategy was followed with levan and poly(vinyl alcohol), using GA as a crosslinker, forming a polymeric network between the two polymers and the GA [68]. Alternatively, levan and carboxymethylcellulose were blended with pluronic acid (F127), forming a gel at a temperature of 37
C [72], which was used to produce a thermosensitive hydrogel. These thermosensitive properties were also obtained with methacrylated levan that was crosslinked with NIPA. Fig. 11.5 shows the methodology to perform the previous synthesis, since levan must be firstly carboxymethylated to produce the metacrylated levan (ML) after. The final step is the crosslinking of the ML with the NIPA to produce the thermosensitive levan hydrogel [28]. This last article also shows another strategy to produce hydrogels from levan. In this case, the functional groups of the polymer were modified to synthesize levan derivatives that could be crosslinked with the respective compound. In fact, several methacrylated levans were photocrosslinked with different photoinitiation systems and irradiation units to synthesize hydrogels with tunable properties, while phosphonated and hydrolyzed levan were crosslinked with BDDE [70]. Concerning applications in medicine, levan hydrogels showed usually a positive result in terms of biocompatibility. In vitro and in vivo studies indicated that these materials were not cytotoxic toward different cell lines or toward mice [70,72]. However, only one article was found with in vivo tests with mice, and in vivo biocompatibility must be studied in more detail. As an example, Fig. 11.6 illustrates the cytocompatibility results (WST-1 assay) with methacrylated levan (using methacrylic anhydride) (LEV-MA) and with different photoinitiators, such as eosin Y (EY), lithium phenyl-2,4,6trimethylbenzoylphosphinate (LAP), and Irgacure 2959 (I2959). It can be observed how no cytotoxic effects were produced in eight days toward mouse fibroblast 3T3 cells. Levan hydrogels have also a high degree of swelling, which can be a key feature for wound dressing or antiwrinkle purposes. Moreover, it has been shown that some properties, such as injectability, adhesivity, elasticity, and strength, can be modified depending on the composition of the hydrogel [70e72]. Finally, levan hydrogels have been proposed as drug delivery systems [67], as their release can be controlled depending on the material composition or even the stimuli (e.g., temperature) [28], thus broadening the potential applications of these hydrogels in several fields. Nonetheless, as shown in Table 11.1, more studies need to be carried out on levan-based hydrogels, proposing new blends and methodologies to synthesize them. This is a mandatory task for the future, as the characteristics of the structure and functional groups of levan hinder the hydrogel synthesis. Moreover, future studies are needed to test the biocompatibility of levan-based hydrogels both in vitro and in vivo. In this context, the effect of levan on the surface properties of hydrogels (i.e., contact angle) can be important to determine the uses of these hydrogels in biomedicine. Also, there is a lack of experiments regarding the degradation of levan-based hydrogels under physiological conditions, which may be important depending on the final application.

11.4 Conclusions and future perspectives Although levan is considered a nongelling polysaccharide due to its particular structure, different methodologies have been proposed to produce hydrogels. So far, the only positive results have been achieved by using BDDE as a crosslinker.

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TABLE 11.1 Levan-based hydrogels. Crosslinker

Encapsulated drug

Application

Particularities

References

Levan

BDDE

Amphotericin B

Antifungal

High degree of swelling In vitro cell assay In vitro antifungal activity Pore diameter 5e50 microns High storage modulus (105 Pa)

[67]

Levan

GA and PVA

e

Capture influenza virus

High swelling ratio Capture the virus with swabs and air sampler The addition of PVA increased the strength Pore size 10e50 microns

[68]

Levan and gellan gum

Polymeric blend and CaCl2 and alkaline conditions

e

e

Strong solid-like behavior Possibility of developing injectable systems Possibility of controlling swelling ratio At least 97% of water retention

[69]

Methacrylated levan

NIPA

5-Aminosalicylic acid

Temperaturesensitive hydrogel

In vitro cell assay Volume phase transition temperature was closer to the physiological temperature

[28]

Several methacrylated levans

Photocrosslinking

e

e

In vitro biocompatibility Control of hydrogel properties depending on the hydrogel-forming experimental conditions

[70]

Hydrolyzed and phosphonated levan

BDDE

Resveratrol

Tissue engineering

In vitro biocompatibility Improved adhesive properties

[71]

Levan

Blend of levan with carboxymethylcellulose and pluronic acid

Dermal filler

Elastic modulus higher than a hyaluronic acid hydrogel Levan improved biocompatibility in in vivo experiments with mice Improved fibroblasts proliferation Injectable hydrogel (gel state at 37
C) with thermosensitive properties

[72]

Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine

Compounds

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FIGURE 11.4 Dried levan hydrogel after using BDE as crosslinker. Reprinted from Demirci T, Hasköylü ME, Eroglu MS, Hemberger J, Öner ET. Levan-based hydrogels for controlled release of Amphotericin B for dermal local antifungal therapy of Candidiasis. Eur J Pharm Sci 2020;145:105255, Copyright (2020), with permission from Elsevier.

FIGURE 11.5 Methodology to produce levan methacrylate hydrogel and its crosslinking with NIPA. Adapted with permission from Osman et al. [28]. Reprinted from Osman A, Oner ET, Eroglu MS. Novel levan and pNIPA temperature sensitive hydrogels for 5-ASA controlled release. Carbohydr Polym 2017;165:61e70, Copyright (2017), with permission from Elsevier.

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FIGURE 11.6 In vitro cytocompatibility results with methacrylated levan hydrogels. Adapted with permission from Berg et al. [70]. Reprinted from Berg A, Oner ET, Combie J, Schneider B, Ellinger R, Weisser J, et al. Formation of new, cytocompatible hydrogels based on photochemically crosslinkable levan methacrylates. Int J Biol Macromol 2018;107:2312e9, Copyright (2018), with permission from Elsevier.

However, levan has been successfully included in different hydrogel networks via polymer blending or by synthesizing an interpenetrated network. Moreover, the functional groups present in levan can be modified to produce polymeric derivatives that can be crosslinked with more conventional products. Due to the advantages of its properties, levan could become a new source for the production of hydrogels and the development of new materials in biomedicine or the basic materials industry. In fact, levan has been proposed as a substitute for several polymers (e.g., hyaluronic acid) due to its biocompatibility, adhesive and swelling properties, as well as the possibility of being used to form drug delivery systems. However, there is still a lack of results concerning in vitro and in vivo assays, as well as mechanical tests, which may be helpful to identify future applications of levan-based hydrogels. Nevertheless, the production of levan hydrogels has been studied for the last 10 years and more research is currently underway. Furthermore, to clarify and design new routes for producing levan-based hydrogels, rheological studies have to be performed to identify different polymeric behaviors in solution and how they can modify the proposed methodologies to produce hydrogels.

References [1] Nasir A, Sattar F, Ashfaq I, Lindemann SR, Chen MH, Van den Ende W, et al. Production and characterization of a high molecular weight levan and fructooligosaccharides from a rhizospheric isolate of Bacillus aryabhattai. LWT 2020;123:109093. [2] Chambert R, Gonzy-Treboul G. Levansucrase of Bacillus subtilis: kinetic and thermodynamic aspects of transfructosylation process. Eur J Biochem 1976;62:55e64. [3] Chambert R, Treboul G, Dedonder R. Kinetic studies of levansucrase of Bacillus subtilis. Eur J Biochem 1974;41:235e300. [4] Meng G, Fütterer K. Structural framework of fructosyl transfer in Bacillus subtilis levansucrases. Nat Struct Biol 2003;10:935e41. [5] He C, Yang Y, Zhao R, Qu J, Jin L, Lu L, et al. Rational designed mutagenesis of levansucrase from Bacillus licheniformis 8-37-0-1 for product specificity study. Appl Microbiol Biotechnol 2018;102:3217e28. [6] González-Garcinuño A, Tabernero A, Domínguez A, Galán MA, Martín del Valle E. Levan and levansucrases: polymer, enzyme, microorganisms and biomedical applications. Biocatal Biotransformation 2018;36:233e44. [7] González-Garcinuño A, Tabernero A, Marcelo G, Martín del Valle E. A comprehensive study on levan nanoparticles formation: kinetics and selfassembly modelling. Int J Biol Macromol 2020;147:1089e98. [8] Ullrich M. Bacterial polysaccharides: current innovations and future trends. Norfolk: Horizon Scientific Press; 2009. [9] Rehm BH. Microbial production of biopolymers and polymer precursors. Applications and perspectives. Norfolk: Horizon Scientific Press; 2009. [10] Stivala SS, Bahary NS. Save-diluted-solution parameters of the levan of Streptococcus salivarius in various solvents. Carbohydr Res 1978;67:17e21. [11] Vina I, Karsakevich A, Gonta S, Linde R, Bekers M. Influence of some physicochemical factors on the viscosity of aqueous levan solutions of Zymomonas mobilis. Acta Biotechnol 1998;18:167e74. [12] Liu C, Lu J, Lu L, Liu Y, Wang F, Xiao M. Isolation, structural characterization and immunological activity of an exopolisaccharide produced by Bacillus licheniformis 8-37-0-1. Bioresour Technol 2010;101:5528e33. [13] Srikanth R, Sundhar Reddy CHSS, Siddartha G, Ramaiah MJ, Uppuluri KB. Review on production, characterization and applications of microbial levan. Carbohydr Polym 2015;12:102e14. [14] Xu X, Gao C, Liu Z, Wu J, Han J, Yan M, et al. Characterization of the levan produced by Paenibacillus bovis sp. nov BD3526 and its immunological activity. Carbohydr Polym 2016;144:178e86. [15] Leibovici J, Kopel S, Siegal A, Gal-Mor O. Effect of tumor inhibitory and stimulatory doses of levan, alone and in combination with cyclophosphamide, on spleen and lymph nodes. Int J Immunopharmacol 1986;8:391e403. [16] Calazans GMT, Lima RC, de França FP, Lopes CE. Molecular weight and antitumour activity of Zymomonas mobilis levans. Int J Biol Macromol 2000;27:245e7.

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[17] Yoon EJ, Yoo SH, Cha J, Lee HG. Effect of levan’s branching structure on antitumor activity. Int J Biol Macromol 2004;34:191e4. [18] Sarilmiser HK, Öner ET. Investigation of anti-cancer activity of linear and aldehyde-activated levan from Halomonas smyrnensis AAD6t. Biochem Eng J 2014;92:28e34. [19] Esawy MA, Amer H, Gamal-Eldeen AM, El Enshasy HA, Helmy WA, Abozeid MAM, et al. Scaling up, characterization of levan and its inhibitory role in carcinogenesis initiation stage. Carbohydr Polym 2013;95:578e87. [20] Dahech I, Harrabi B, Hamden K, Feki A, Mejdoub H, Belghith H, et al. Antioxidant effect of nondigestible levan and its impact on cardiovascular disease and atherosclerosis. Int J Biol Macromol 2013;58:281e6. [21] Erginer M, Akcay A, Coskunkan B, Morova T, Rende D, Bucak S, et al. Sulfated levan from Halomonas smyrnensis as a bioactive, heparin-mimetic glycan for cardiac tissue engineering applications. Carbohydr Polym 2016;149:289e96. [22] Tabernero A, Baldino L, González-Garcinuño A, Cardea S, Martín del Valle E, Reverchon E. Supercritical CO2 assisted formation of composite membranes containing an amphiphilic fructose-based polymer. J CO2 Util 2019;34:274e81. [23] Markowiak P, Slizewska K. Effects of probiotics, prebiotics and synbiotics on human health. Nutrients 2017;9:1021. [24] Yang Y, Galle S, Le MHA, Ziljstra RT, Gänzle MG. Feed fermentation with reuteran and levan-producing Lactobacillus reuteri reduces colonization of weanling pigs by enterotoxigenic Escherichia coli. Appl Environ Microbiol 2015;81:5743e52. [25] Adamberg K, Tomson K, Talve T, Pudova K, Puurand M, Visnapun T, et al. Levan enhances associated growth of Bacteroides, Escherichia, Streptococcus and Faecalibacterium in fecal microbiota. PLoS One 2015;10(12):e0144042. [26] Barone JR, Medunets M. Thermally processed levan polymers. Carbohydr Polym 2007;69:554e61. [27] Chen X, Gao H, Ploehn HJ. Montmorillonite-levan nanocomposites with improved thermal and mechanical properties. Carbohydr Polym 2014;101:565e73. [28] Osman A, Öner ET, Eroglu MS. Novel levan and pNIPA temperature sensitive hydrogels for 5-ASA controlled release. Carbohydr Polym 2017;165:61e70. [29] Kim SJ, Baem PK, Chung BH. Self-assembled levan nanoparticles for targeted breast cancer imaging. Chem Commun 2015;51:107e10. [30] Renuart E, Viney C. Biological fibrous materials: self-assembled structures and optimised properties. In: Elices M, editor. Structural biological materials. Oxford: Pergamon/Elsevier Science; 2000. p. 221e67. [31] Demir Sezer A, Sarilmiser K, Rayaman E, Cevikbas A, Öner ET, Akbuga J. Development and characterization of vancomycin-loaded levan based microparticular system for drug delivery. Pharmaceut Dev Technol 2015;25:1e8. [32] Demir Sezer A, Kazak H, Öner ET, Akbuga J. Levan-based nanocarrier for peptide and protein drug delivery optimization and influence of experimental parameters on the nanoparticles characteristics. Carbohydr Polym 2011;84:358e63. [33] Tabernero A, González-Garcinuño A, Sánchez-Álvarez JM, Galán MA, Martín del Valle E. Development of a nanoparticle system based on a fructose polymer: stability and drug release studies. Carbohydr Polym 2017;160:26e33. [34] González-Garcinuño A, Masa R, Hernández M, Domínguez A, Tabernero A, Martín del Valle E. Levan-capped silver nanoparticles for bactericidal formulations: release and activity modelling. Int J Mol Sci 2019;20:1502. [35] Akturk O. The anticancer activity of doxorubicin-loaded levan-functionalized gold nanoparticles synthesized by laser ablation. Int J Biol Macromol 2022;196:72e85. [36] Cinan E, Cesur S, Haskoylu ME, Gunduz O, Oner ET. Resveratrol-loaded levan nanoparticles produced by electrohydrodynamic atomization technique. Nanomaterials 2021;11:2582. [37] Richa R, Roy Choudhury A. Exploration of polysaccharide based nanoemulsions for stabilization and entrapment of curcumin. Int J Biol Macromol 2020;156:1287e96. [38] Donot F, Fontana A, Baccon J, Schorr-Galindo S. Microbial exopolysaccharides: main examples of synthesis, excretion, genetics and extraction. Carbohydr Polym 2012;87:951e2. [39] Venugopal V. Marine polysaccharides: food applications. Boca Raton, FL: CRC Press; 2011. [40] Shih L, Yu YT. Simultaneous and selective production of Levan and poly (g-glutamic acid) by Bacillus subtilis. Biotechnol Lett 2005;27:103e6. [41] Linde D, Rodríguez-Colinas B, Estévez M, Poveda A, Plou FJ, Fernández-Lobato M. Analysis of neofructooligosaccarides production mediated by the extracelular b-fructofuranosidase from Xanthophyllomyces dendrorhous. Bioresour Technol 2012;109:123e30. [42] dos Santos LF, Bazani Cabral De Melo FC, Martins Paiva WJ, Borsato D, Corradi Custodio Da Silva ML, Pedrine Colabone Celligoi MA. Characterization and optimization of levan production by Bacillus subtilis natto. Rom Biotechnol Lett 2013;18:8413e22. [43] Kang HK, Yun SI, Lim TY, Xia YM, Kim D. Cloning of levansucrase from Leuconostoc mesenteroides and its expression in Pichia pastoris. Food Sci Biotechnol 2011;20:277e81. [44] Bersaneti GT, Baldo C, Colabone Celligoi MAP. Immobilization of levansucrase: strategies and biotechnological applications. J Chil Chem Soc 2019;64:4377e81. [45] González-Garcinuño A, Ruiz S, Sánchez-Muñoz A, Tabernero A, Martín del Valle E. Biotechnological strategies to produce levan: mass transfer and techno-economical evaluation. Chem Eng Process 2019;141:107529. [46] Lorenzoni ASG, Aydos LF, Klein MP, Ayub MAZ, Rodrigues RC, Hertz PF. Continuous production of fructooligosaccharides and invert sugar by chitosan immobilized enzymes: comparison between in fluidized and packed bed reactors. J Mol Catal B Enzym 2015;111:51e5. [47] Esawy MA, Mahmoud DAR, Fattah AFA. Immobilisation of Bacillus subtilis NRC33a levansucrase and some studies on its properties. Braz J Chem Eng 2008;25:237e46. [48] Szwengiel A, Goderska K, Gumienna M. Synthesis of B(2-6)-linked fructan with a partially purified levansucrase from Bacillus subtilis. J Mol Catal B Enzym 2016;131:1e9.

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[49] Jathore NR, Bule MV, Tilay AV, Annapure US. Microbial levan from Pseudomonas fluorescens: characterization and medium optimization for enhanced production. Food Sci Biotechnol 2012;21:1045e53. [50] Ua-Arak T, Jakob F, Vogel RF. Fermentation pH modulates the size distributions and functional properties of Gluconobacter albidus TMW2.1191 levan. Front Microbiol 2017;8:807. [51] Franken J, Brandt BA, Tai SL, Bauer FF. Biosynthesis of levan, a bacterial extracellular polysaccharide in the yeast Saccharomyces cerevisae. PLoS One 2013;8:e77499. [52] Arvidson SA, Rinehart BT, Gadala-María F. Concentration regimes of solutions of levan polysaccharide from Bacillus sp. Carbohydr Polym 2006;65:144e9. [53] Peng J, Xu W, Ni D, Zhang W, Zhang T, Guang C, et al. Preparation of a novel water-soluble gel from Erwinia amylovora levan. Int J Biol Macromol 2019;122:469e78. [54] Blanco-López M, González-Garcinuño A, Tabernero A, Martín del Valle EM. Steady and oscillatory shear flow behaviour of different polysaccharides with Laponite. Polymers 2021;13:966. [55] Benigar E, Dogsa I, Stopar D, Jamnik A, Cigic IK, Tomsic M. Structure and dynamics of a polysaccharide matrix: aqueous solutions of bacterial levan. Langmuir 2014;30:4172e82. [56] Bouallegue A, Chaari F, Casillo A, Corsaro MM, Bachoual R, Ellouz-Chaabouni S. Levan produced by Bacillus subtilis AF17: thermal, functional and rheological properties. J Food Meas Char 2022;16:440e7. [57] Hundschell CS, Braun A, Wefers D, Vogel RF, Jakob F. Size-dependent variability in flow and viscoelastic behaviour of levan produced by Gluconobacter albidus TMW 2.1191. Foods 2020;9:192. [58] González-Garcinuño A, Tabernero A, Marcelo G, Sebastián V, Arruebo M, Santamaría J, et al. Differences in levan nanoparticles depending on their synthesis route: microbial vs cell-free systems. Int J Biol Macromol 2019;137:62e8. [59] Kaith BS, Singh A, Sharma AK, Sud D. Hydrogels: synthesis, classification, properties, and potential applications-A brief review. J Polym Environ 2021;29:3827e41. [60] Varaprasad K, Raghavendra GM, Jayaramudu T, Yallapu MM, Sadiku R. A mini review on hydrogels classifications and recent developments in miscellaneous applications. Mater Sci Eng C 2017;79:958e71. [61] Hassan CM, Peppas NA. Cellular PVA hydrogels produced by freeze/thawing. J Appl Polym Sci 2000;76:2075e9. [62] De Jong S, De Smedt S, Demeester J, Van Nostrum C. Biodegradable hydrogels based on stereocomplex formation between lactic acid oligomers grafted to dextran. J Contr Release 2001;72:47e56. [63] Aoki H, Assaf SA, Katayama T, Phillips GO. Characterization and properties of Acacia Senegal (L.) Willd. var. Senegal with enhanced properties (Acacia (sen) SUPER GUMÔ ): part 2dmechanism of the maturation process. Food Hydrocoll 2007;21:329e37. [64] Jayaramudu T, Raghavendra GM, Varaprasad K, Sadiku R, Ramam K, Raju KM. Iota-carrageenan-based biodegradable Ag0 nanocomposite hydrogels for the inactivation. Carbohydr Polym 2013;95:188e94. [65] Nep EI, Conway BR. Grewia gum 2: mucoadhesive properties of compacts and Gels. Trop J Pharm Res 2011;10:393e401. [66] Liow SS, Dou Q, Kai D, Karim AA, Zhang K, Xu F, et al. Thermogels: in situ gelling biomaterial. ACS Biomater Sci Eng 2016;2:295e316. [67] Demirci T, Hasköylü ME, Eroglu MS, Hemberger J, Öner ET. Levan-based hydrogels for controlled release of Amphotericin B for dermal local antifungal therapy of Candidiasis. Eur J Pharm Sci 2020;145:105255. [68] Kim SJ, Bae PK, Choi M, Keem JO, Chung W, Shin YB. Fabrication and application of levan-PVA hydrogel for effective influenza virus capture. ACS Appl Mater Interfaces 2020;12:29103e9. [69] Nair R, Choudhury AR. Synthesis and rheological characterization of a novel shear thinning levan gellan hydrogel. Int J Biol Macromol 2020;159:922e30. [70] Berg A, Öner ET, Combie J, Schneider B, Ellinger R, Weisser J, et al. Formation of new, cytocompatible hydrogels based on photochemically crosslinkable levan methacrylates. Int J Biol Macromol 2018;107:2312e9. [71] Selvi SS, Hasköylu ME, Genç S, Öner ET. Synthesis and characterization of levan hydrogels and their use for resveratrol release. J Bioact Compat Polym 2021;36:464e80. [72] Choi WI, Hwang Y, Sahu A, Min K, Sung D, Tae G, et al. An injectable and physical levan-based hydrogel as a dermal filler for soft tissue augmentation. Biomater Sci 2018;6:2627e38.

Chapter 12

Hydrogels based on schizophyllan Yachen Hou1 and Jingan Li2 First Affiliated Hospital of Zhengzhou University, Zhengzhou, China; 2School of Materials Science and Engineering & Henan Key Laboratory of

1

Advanced Magnesium Alloy & Key Laboratory of Materials Processing and Mold Technology (Ministry of Education), Zhengzhou University, Zhengzhou, China

12.1 Introduction Kikumoto et al. discovered the treatment effect of schizophyllan (SPG) firstly [1]. The SPG belongs to the glucose polysaccharide family and can be obtained from a variety of sources, such as fungi, bacteria, algae, lichens, and some plants. SPG is a nonionic water-soluble hyperglycan with a b-(1 3)-linked main chain, and each b-(1 6)-linked glucose side chain on every three residues [2]. It is formed by three independent 6/1 helices stabilized by hydrogen bonds, and b-(1 6)D-glucose side chain protruding outwards, as shown in Fig. 12.1A and B [3]. The most notable feature of SPG compared to other polysaccharides is its triple helix structure, formed by self-assembly of three b-glucan chains via hydrogen bonds. While the triple helix structure of SPG can be dissociated by changing the solvent type or pH condition to form a single chain [4]. This part mainly introduces the preparation, structure, and gelation behavior of SPG.

12.1.1 Preparation and structure Hot water extraction (HWE), ultrasonic-assisted extraction (UAE), supercritical fluid extraction (SFE), and enzymeassisted extraction (EAE) were the extraction manners of the SPG [5]. In order to obtain high purity SPG more effectively, the hot water-supercritical CO2 binary system and ethanol-ion exchange separation precipitation and size exclusion chromatography were established [6]. SPG is a rod-like triple helix state in water at pH 7, while the condition converts to strongly alkaline solution (pH > 13) or DMSO solution, the SPG will disperse into a single random curly state [7]. Norisuye et al. analyzed the triple helix configuration of SPG in water [8,9], and HyunBae et al. observed the rod-like structure and helicity of SPG by atomic force microscopy and scanning electron microscopy [10]. Splice pleated polysaccharides have no charge, and the formation of ion pairs is not the cause of complexation. Hydrogen bonds are formed between glucose and the base portion [11]. Some researches thought the conformation transition of the SPG is related to the hydrogen bond of water [12]. The water owns different phases in different status, and the triple helix arrangement of SPG was embraced by bound water, structured water, and loosely structured water, in Fig. 12.1C. These different water structures are separated by fuzzy interfaces around three helical layers, forming different phases with specific thermodynamic properties [13,14]. The SPG solution exhibits two highly synergistic configuration transformations after heat to 6 or 135
C, while after adding DMSO (content increased to over 87%) or NaOH (pH > 13), the hydrogen bond of the triplex structure will be resolved [15]. While the SPG cannot return to the original triple helix state after the condition restores [16]. Therefore, the structural reconstruction of its SPG (DRSPG) requires the use of denaturation and renaturation (DR) method: (1) the SPG is dissolved in DMSO and then dialysis with H2O; (2) dissolve with NaOH and then neutralize with HCl; (3) heat to above 135 and cool to room temperature [17]. In addition, some study showed the DRSPG chains form of linear hairpins, particles, and complex clusters [18]. Studies on the DR processes of DRSPG chains will help clarify the SPG superhelix phenomenon, which is the basic behavior of DRSPG as drug or gene delivery [19].

Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine. https://doi.org/10.1016/B978-0-323-95351-1.00003-X Copyright © 2024 Elsevier Inc. All rights reserved.

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FIGURE 12.1 Schematic representation of (A) chemical structure of SPG; (B) the triple helix of SPG [3]; (C) cross-sectional vision of the SPG triple helix in water. The four circles demonstrated the boundaries of the structures: helix core-bound water (1.68 nm, yellow), bound water-structured water (2.05 nm, blue), structured water-loosely structured water (2.31 nm, purple), and loosely structured water-free water (2.79 nm, green) [3,13]. (A) and (B) Reprinted from Itou T, Teramoto A, Matsuo T, Suga H. Isotope effect on the orderedisorder transition in aqueous schizophyllan. Carbohydr Res 1987;160:243e57, Copyright (1987), with permission from Elsevier; (C) Reprinted from Kazuto Y, Akio T, Naotake N, Kikuchi K, Miyazaki Y, Sorai M. Static water structure detected by heat capacity measurements on aqueous solutions of a triple-helical polysaccharide schizophyllan. Biomacromolecules 2003;4(5):1348e56, Copyright (2003), with permission from American Chemical Society.

12.1.2 Gelation behavior The gelation procedure can be divided into two kinds according to the methods of cross-linking [20]. One is the chemical gel by the cross-linking via covalent bond to form polymer chain [21]. The other is the physical gel, which mainly through hydrogen bond hydrophobic interaction and coulomb interaction and other noncovalent bond association to form the connection region [22]. Physical gels can be further divided into four types according to their temperature dependence: (1) cold-set gels, such as agarose and carrageenan, are formed by cooling aqueous solutions; (2) heat-curing gel: such as methyl cellulose, broccoli, konjac, glucomannan, and other heated gel solution; (3) reentrant gels, such as xyloglucan; (4) reverse reentrant gels, such as a mixed solution of methyl cellulose and gelatin. SPG is a kind of cold-solidified gel, which can form weak gel after cooled in aqueous solution below 6
C [23,24]. To increase the capacity and strength of SPG gels, small molecules such as borax or sorbitol can be added in gel [25]. The ability of borax to promote gel is due to the reaction between borate ions and SPG side chain’s hydroxyl groups to promote its gelation behavior in the solution [26]. The gelation process of SPG was monitored by dynamic viscoelastic measurement, and it was found that the gelation rate was faster when the content of borax was high, while the concentration of SPG was higher under the temperature was low [27]. However, the use of borax to promote the cross-linking of SPG gel is nonpermanent and controlled by complexation equilibrium [28]. The gelling mechanism of SPG and sorbitol is controversial [29]. One study suggests that the presence of sorbitol may reduce the mobility of water molecules, as showed in Fig. 12.2A. The entanglement degree of SPG chain is greatly enhanced and push SPG molecules together to form a three-dimensional gel network [30]. Another study on the gelation mechanism showed that the gelation of sorbitol aqueous solution is combined with a highly conformational transition from disordered condition to ordered condition of SPG’s triple helix, as showed in Fig. 12.2B [31]. Under the condition of sorbitol content stabilization, the gel temperature of SPG-sorbitol solution was not affected by SPG concentration, and the gel temperature cut down mildly with the decrease of SPG molecular weight (MW) [32]. The gelation temperature increased significantly with the increase of sorbitol content, but the enthalpy during gelation had nothing to do with sorbitol content [33]. While Fang et al. argued that the water-based SPG-sorbitol gel acts like a very concentrated solution that does not flow over the commonly observed time range and should be called a structured liquid rather than a weak gel [34]. In addition to the addition of small molecules, co-gels of SPG and protein are also ways to improve the SPG gel capacity [35]. Some studies have used periodate oxidation method to locally modify the SPG side chain to form periodate oxidation SPG (POSPG), which can form elastic gel with gelatin and has good mechanical properties [36]. The co-gel of

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FIGURE 12.2 Illustration of the gelation mechanisms of SPG/sorbitol aqueous solutions: (A) the gelation mechanism via the aggregation of SPG triple helices [30]; (B) the gelation mechanism via the association of broken parts of SPG triple helices caused by sorbitol [31]. (A) Reprinted from Fuchs T, Richtering W, Burchard W, Kajiwara K, Kitamura S. Gel point in physical gels: rheology and light scattering from thermoreversibly gelling schizophyllan. Polym Gels Netw 1998;5(6):541e59, Copyright (1998), with permission from Elsevier.

POSPG with protein is based on the Schiff base reaction between the amino group and the aldehyde group on the POSPG chain [37]. Therefore, this complexation method greatly broadens the application of SPG as gel and provides a possible approach for using SPG as tissue repair [38].

12.2 Biological function Features of SPG confirmed first was antitumor and immunomodulatory functions, which can enhance the immune activity of combined chemotherapy and prevent tumor metastasis [39]. A review of the studies over the past decades indicated that SPG has the advantages of good biosafety and tolerability [40]. When used in combination with chemotherapy or radiation, the immune suppression associated with the treatment process can be reduced and the production of white blood cells can be accelerated to improve the recovery of cancer patients, with the side effect that patients are more prone to nausea [41,42]. One of the key mechanisms of SPG’s biological activity is its ability to trigger phagocytosis by macrophages, a product of the long evolution of the body’s immune system against bacterial or fungal pathogens [43e46]. Since 2019, Givosiran, a siRNA drug combined with the sugar derivative N-acetylgalactosamine (GalNac), has received FDA approval [47]. The first human Phase 1 study (Study No. 73001-C111) also confirmed the safety of sicD40SPG in humans [48]. These researches increased the interest of gene therapy, and the status and importance of SPG in drug delivery field were also indicated. As shown in Fig. 12.3, the most common antitumor mechanisms of SPG include cell cycle arrest, antiangiogenesis, and apoptosis, all of which play a direct tumor-killing ability or immunomodulatory role, protect the immune system, and induce indirect tumor-killing effect [49e51]. The reason why SPG has such excellent properties is that its spiral structure forms a one-dimensional hydrophobic cavity during the renaturation process of SPG, which could contain functional nanomolecular and nanoparticles by an induced-fit methods [52]. A single conjugated polymer or molecular combination could be integrated into a onedimensional intermediate space to form water-soluble one-dimensional nanocomposites [53]. SPG single chains can be used as a good gene carrier because they are linked together by medium hybrid hydrogen bonds in the same way as nucleic acids. In addition, the specific action between SPG and Dectin-1 increases the cell targeting of SPG. This section reviews the antitumor properties and drug delivery capacity of SPG.

12.2.1 Antineoplastic activity The host-mediated antitumor activity of SPG aqueous solution against 180 sarcoma was demonstrated for the first time by Komatsu et al. [54]. In SPG-treated mice, the increase of IL-2 and interferon-g produced by peripheral blood monocytes further confirmed its anticancer effect, while in thymectomy mice, the loss of antitumor activity of SPG confirmed that its antitumor activity was related to T cells [2]. Since 1986, SPG has been used as an anticancer drug in Japan for its good clinical treatment effect. The effects have been observed in patients with lung cancer, cervical cancer, and gastric cancer [55]. The antitumor effect of SPG is host-mediated and connected with the activation of macrophages as well as the T lymphocytes and could increase the interferon production [56]. To explore the clinical effects, stochastic research of patients with gastric cancer was studied. The results showed that the SPG combined with conventional chemotherapy (tegafur or mitomycin C and 5-fluorouracil) improved median survival in 367 patients, and SPG-assisted

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FIGURE 12.3 The immunomodulation, prevention, and direct antitumor activity of SPG.

immunochemotherapy was randomized to 326 patients with gastric cancer and enhanced the survival in gastric cancer patients at stage III [57]. A randomized clinical study was conducted in 220 patients with stage II or III cervical cancer who underwent radiotherapy. Especially, the survival of patients who taking up SPG was significantly longer compared with normal patients at stage II [2]. In a prospective randomized clinical trial of 312 patients cured with a combination of surgical chemoradiotherapy (fluorouracil) and SPG, five-year survival was significantly longer in patients treated with SPG than in patients not cured with SPG [58]. SPG has also been demonstrated that cure the effect of the extended lifetime of patients with head and neck cancer [59]. Fifteen patients received adjuvant immunotherapy with SPG, and the five-year cumulative survival rate was 86.7% and 73.4% in the SPG taking group and control group, respectively [60]. The SPG group could quickly recover the cellular immune function damaged by radiotherapy and chemotherapy [28,59]. In addition, SPG aqueous solution combined with chemotherapy drugs could against a variety of tumors in mice and rats efficiently, such as AMC-60 fibrosarcoma, MM-46 cancer, MH-134 cancer, BC-47 bladder tumor, and A-755 breast cancer [61]. There is evidence that the immunomodulatory activity of SPG is related to the MW, and the higher the MW of SPG, the greater the effect on the immune system [62]. High MW dextran has a more stable structure and can be recognized directly by specific receptors on the surface of immune cells. Another influencing factor is the residence time of high MW dextran in the intestinal system and the gradual slowing of degradation metabolism [63]. b-glucan has low MW and short side chains (3)-D-glucan schizophyllan to deliver antisense-oligonucleotides with avoiding lysosomal degradation. Biomaterials 2005;26(23):4866e73. [36] Fang Y, Takahashi R, Nishinari K. Protein/polysaccharide cogel formation based on gelatin and chemically modified schizophyllan. Biomacromolecules 2005;6(6):3202e8.

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Further reading [1] Lehtovaara BC, Gu FX. Pharmacological, structural, and drug delivery properties and applications of 1,3-b-glucans. J Agric Food Chem 2011;59(13):6813e28.

Chapter 13

Curdlan based hydrogels Natasha Aquinas, Ramananda Bhat M and Subbalaxmi Selvaraj Department of Biotechnology, Manipal Institute of Technology, Manipal Academy of Higher Education (MAHE), Manipal, Karnataka, India

13.1 Overview Polymers are materials made up of repeating units which form a macromolecule. On the basis of their origin, they are classified as synthetic polymers, natural polymers, and semisynthetic polymers. Natural polymers are a class of polymers which are acquired from natural sources such as plants, animals, and microorganisms. The advantage of natural polymers is that they are comparatively cost effective, biodegradable, and biocompatible too. They are mainly grouped into six classes: proteins, polysaccharides, polyesters, polynucleotides, polyisoprenes, and lignin [1]. Bacterial polymers constitute four major groups: polysaccharides, polyesters, polyamides, and inorganic polyanhydrides [2]. Polysaccharides are known to be one of the most abundant renewable resources in the world. Polysaccharides can be subdivided into broad categories as: exopolysaccharides, which are secreted outside the cells or synthesized extracellularly by certain enzymes which are anchored to the cell wall. Some examples of exopolysaccharides include xanthan, dextran, cellulose, alginate, hyaluronic acid, curdlan; capsular polysaccharides such as K30 antigen and intracellular polysaccharide, for example, glycogen. Polysaccharides can also be grouped based on their chains, that is, branched or straight chains. Polysaccharides that comprise monomers made of D-glucose and linked by glycosidic bonds are known as glucans. The bacterial exopolysaccharide, curdlan is a b-1,3 glucan which obtained immense importance over the recent years due to its varied properties and applications. Its rheological properties, solubility, and, especially, simple structure are some of its highlights. Curdlan was discovered by accident in 1962. After great efforts by Harada and his coworkers in 1962 to isolate microorganisms that could make use of petrochemical materials, they finally isolated an organism from soil that could grow on media which consisted of only one carbon sourced10% ethylene glycol. They named this particular organism as Alcaligenes faecalis var myxogenes 10C3. This organism gave rise to a new b-glucan which consisted of about 10% succinic acid and, hence, they called it succinoglucan. Later one day, they noticed that the medium did not become viscous nor did it produce any succinoglucan. However, all the glucose in the medium was being consumed. This drove them to investigate if any other product was being formed instead of succinoglucan. They examined this special product and deduced that it was a neutral polysaccharide. It was this neutral polysaccharide which was called curdlan in the year 1966 [3]. A mutant strain of 10C3 was isolated from the stock culture that could produce just curdlan. Therefore, it was by sheer luck and chance that they were successful in isolating an organism that could produce curdlan. Curdlan is a linear, neutral glucan which is constituted exclusively of b-1,3 glycosidic linkages without any branching. Curdlan’s structure is shown in Fig. 13.1. Due to its potential to curdle when heated it was named as curdlan [4].

13.2 Biosynthesis of curdlan The operon crdASC consists of four genes which are crdA, crdS, crdC, and crdR. Biosynthesis of curdlan needs these four genes. Their exact role in biosynthesis is not very clear. However, crdS regulates the catalytic subunit of b-1,3-glucan synthase, whereas the expression of the operon is mainly due to the crdR gene [5]. The biosynthesis of curdlan begins when glucose moves inside the cell through an active transporter such as the PEP-PTS (phosphoenolpyruvate glucose phosphotransferase) or permease. It is noted that most sugars gain entry into the cell by the PEP-PTS pathway which give rise to sugar-6-phosphates [6]. In the same way, on entering the cell, glucose is phosphorylated to glucose-6-phosphate by Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine. https://doi.org/10.1016/B978-0-323-95351-1.00005-3 Copyright © 2024 Elsevier Inc. All rights reserved.

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FIGURE 13.1 Chemical structure of curdlan showing b-1,3 glycosidic bonds.

enzyme hexokinase. Glucose-6-phosphate is then converted to glucose-1-phosphate. This is brought about by phosphoglucomutase. Uridine diphosphate (UDP)-glucose is one of the main precursors for the synthesis of curdlan. UDP-glucose and lipid-P-glucose are synthesized from glucose-1-phosphate. This reaction is catalyzed by uridine triphosphate (UTP) and UDP glucose phosphorylase. Next step is polymerization which is brought about by enzyme curdlan synthase. It moves one molecule of glucose from UDP glucose to the nascent polymer chain, which produces a UDP molecule [7]. This UDP is again transformed back to UTP by UDP-kinase and through ATP which is derived either from the glycolysis pathway or the tricarboxylic acid cycle (TCA) [8] (Fig. 13.2).

FIGURE 13.2 Metabolic pathway of curdlan biosynthesis: 1: hexokinase, 2: phosphoglucomutase, 3: UDP glucose phosphorylase, 4: glycosyltransferase, 5: curdlan synthase, 6: glucose-6-phosphate dehydrogenase, 7: 6-phosphogluconate dehydrogenase, 8: ribose phosphate diphosphokinase, 9: orotate phosphoribosyl transferase and orotidine-5-phosphate decarboxylase, 10: uridylate kinase, 11: UDP kinase.

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Curdlan is a relatively lesser-known polysaccharide which has immense potential in various industries. It was approved to be used in the food industry in the year 1996 by the U.S. Food and Drug administration (FDA). Subsequently in 1996, the United States of America catapulted curdlan into the market. It was named Pureglucan. In 1989, it was also accepted for use in Taiwan, Japan, and Korea [9]. One of its key hallmarks is that it is nontoxic and biodegradable to humans and the environment. Because of this, there is a rise in host of applications. To highlight a few, curdlan was used as a thickener and gelling agent in food items. It has been exploited as a stabilizer in food. It also has the ability to mimic fat, meat, and seafood [10]. The pharmaceutical industry has been utilizing it in the encapsulation of certain drugs such as theophylline, prednisolone, and indomethacin [11].

13.3 Polymer properties with structural features related to hydrogel formation Hydrogels are mainly three-dimensional, hydrophilic polymeric matrices swollen in an aqueous medium. Depending on the hydrogel, they can retain their structural integrity even while expanding their form [12]. Based on some parameters like charge, origin, cross-linking, physical state, configuration, and composition, hydrogels are classified into various categories. The two major classes of hydrogels on the basis of origin are polysaccharide-based hydrogels and protein-based hydrogels. Classification based on the nature of cross-linking include physical hydrogels and chemical hydrogels. Depending on the charges present on the polymer, five types of hydrogel groups are evident. These are nonionic (neutral), anionic, cationic, amphoteric, and zwitterionic. On the basis of their physical state and configuration, hydrogels are classified as solid, semisolid, liquid and crystalline, amorphous, and semicrystalline, respectively. According to the polymer composition, hydrogels are classified as homopolymeric hydrogel, copolymeric hydrogel, semiinterpenetrating network, and interpenetrating network [13]. It must be noted that a particular polymer can fall into multiple categories of classification as mentioned above and is not restricted to any one of them. Curdlan, for example, comes under physical hydrogels under the umbrella of classification based on the nature of cross-linking. It can also be grouped under polysaccharide-based hydrogels and nonionic hydrogels. Curdlan’s unique property of gel formation has led to fascination and interest toward it. After processing, it is seen as a colorless, odorless, tasteless powder insoluble in water and alcohol. However, it is soluble in alkali [14]. When in aqueous solution, two types of gels can be formed namely, thermo-reversible gel and thermo-irreversible gel. The types of gels that are formed are due to different heating temperatures. When the aqueous solution temperature is raised to around 55
C, it forms thermo-reversible gel which is also known as low set gel. However, if the temperature is around 80
C, thermoirreversible gel is formed which is also known as high set gel [4]. The thermo-reversible gel can form a solution again on heating. However, the thermo-irreversible gel does not reform into solution. The main reason for this gelation property can be ascribed to curdlan’s structure. At different temperatures, curdlan’s structure undergoes transformations such that different types of gels are formed. At room temperature, it is a combination of single helices as well as triple helices which are loosely intertwined, whereas at higher temperatures, the structure is more condensed and forms rod-like triple helices [15]. The pH required for gelation is between a broad range of 2 and 10 because of which curdlan has applications in varied fields of food, pharmaceutical, and biomedical sectors. Generally, an ideal hydrogel must possess characteristics such as high fluid absorption, good porosity, photostability, biodegradability, low cost of manufacturing, and storage. It can be said that curdlan hydrogels have a good balance of these characteristics which makes its use widespread.

13.4 Importance of curdlan hydrogel in drug delivery Curdlan triggers certain antimicrobial responses such as phagocytosis and generation of ROS (reactive oxygen species). It has the ability to also stimulate downstream signaling in cells such as dendrites, neutrophils, and macrophages that can then activate NF-kB, mitogen-activated protein kinases, along with a few NFAT transcription factors [11]. It is possible to modify curdlan in a number of ways to prepare grafted curdlan, hydrogels, and nanocomposites. Some properties that curdlan hydrogels include are high flexibility, its ability to swell, as well as biocompatibility. It is because of these properties curdlan hydrogels have potential to be used in pharmaceutical, agricultural, drug delivery, wastewater treatment plants, and biomedical applications [16]. As it has the capability to manage and sustain drug release, it is implemented as a drug delivery carrier. Various examples of different studies done in this field have been discussed briefly in the later sections of the chapter. Many hydrogels require the addition of cross-linkers to actually form a hydrogel. These cross-linking agents could contribute to toxicity. At times, there is also the formation of certain toxic by-products due to the addition of cross-linking agents [17]. Curdlan hydrogels can bypass this problem as they do not necessarily require cross-linking agents to form

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hydrogels and have the inherent property of physical gelation which is attributed to heating temperatures. Therefore, curdlan hydrogels play a very vital role in several fields, especially in drug delivery where humans are directly involved.

13.5 Method of preparation of curdlan hydrogel The preparation of native curdlan hydrogel is simple as mentioned in the previous sections. Heating aqueous solutions of curdlan to approximately 55 and 80
C forms the thermo-reversible and the thermo-irreversible gel, respectively. Owing to the existence of several hydroxyl groups in curdlan, there are strong hydrogen bonds which are present within the molecule rendering it insoluble in water. However, the hydrogen bonds in curdlan can be broken by the addition of any alkaline substance so that it can dissolve in water [18]. To make curdlan-based biomaterials, two different strategies can be employed. The first one is chemical modification of curdlan which involves methods such as carboxymethylation, phosphorylation, and sulfation among others. These modifications of curdlan give rise to curdlan derivatives which have increased solubility in water, thereby making it more useful by increasing its functionality. The second strategy is by using curdlan after the thermal gelation process, freeze-drying method, blended with other materials [19]. The best way to produce curdlan hydrogel is by combining both the above methods such that the modified hydrogel works at its optimum level in terms of biological activity, elasticity, flexibility, and mechanical strength.

13.5.1 Carboxymethylation Carboxymethylation of curdlan helps in improving its water solubility without altering its bioactivity. This broadens its range of applications in commercial settings [9]. A recent study developed pH-responsive micelles made up of carboxymethyl curdlan and were amphiphilic in nature for the delivery of curcumin. Octenyl succinic anhydride (OSA) being an amphiphilic derivative with great properties such as pH responsiveness with greater encapsulation efficiency was used to make an amphiphilic polymer namely OSA-CMCD which is OSA-modified carboxymethyl curdlan. This biopolymer was mainly made to encapsulate and release curcumin. Micelles were prepared for encapsulation purpose, and these micelles showed a larger size after curcumin encapsulation. Maximum encapsulation efficiency was 85.49%. It was noted that the curcumin release was gradual at pH 1.2 or 6.8 in PBS, whereas at pH 7.4, the release was 68% depicting that the drug release is pH-dependent. The potential use of these OSA-CMCD micelles might be in targeted delivery for hydrophobic compounds in the colon [20]. Another study characterized unique carboxymethylated curdlan (CMC)-silica hybrid hydrogels. This work utilized the solegel process to make these hybrid hydrogels of carboxymethyl-curdlan (CMCD) and silica [tetrakis (2-hydroxyethyl) orthosilicatedTHEOS]. In order to find the drug release behavior of this unique hybrid hydrogel, bovine serum albumin (BSA) was used as the model protein. Mechanical strength of gels could be varied by tuning the concentrations of CMCD and THEOS. Elevated concentration of CMCD and THEOS gave rise to stiffer gels. It was seen that the hydrogel allowed easy diffusion of the model protein, BSA through its large and well-connected network. It is believed that the CMCD/THEOS hybrid hydrogels could be used in drug release work for many biomedical applications [21]. Carboxymethyl curdlan is also being used for the preparation and development of nanoparticles which can be used in various biomedical applications. Preparation of gold nanoparticles by using a green synthesis approach was carried out, wherein CMC was used as the reducing as well as stabilizing agent to make CMC-stabilized gold nanoparticles. Gold nanoparticles have a disadvantage of being cytotoxic in biomedical applications. Therefore, an MTT assay was executed using human breast cancer cell line MCF-7 to ascertain the in vitro cytotoxicity of the prepared CMC-stabilized AuNPs. It was seen that these nanoparticles have no to negligible amount of cytotoxicity toward MCF-7 cell line. These CMCAuNPs have carboxylic groups on their surface which are deemed to be nontoxic. Hence, it can be potential carriers in drug delivery for cancer therapies [22]. Another novel type of nanoparticle was prepared from deoxycholic acid (DOCA) and carboxymethyl curdlan. DOCA is the principal component of bile acid, and it can form micelles in water. DOCA attached to carboxymethyl curdlan leads to the formation of self-assembled nanoparticles which are in the range of 192e347 nm. This particular nanoparticle may be useful in some pharmaceutical applications such as antitumor drug delivery [23]. DOCA-modified carboxymethylated curdlan (DCMC) conjugate was studied as a novel carrier of Epirubicin. Epirubicin is an anticancer, anthracycline drug which is employed to treat various carcinomas. However, it comes with the aftermath of certain problems such as some allergic reactions, blood problems, and cardiotoxicity. In order to combat these problems, epirubicin has been used as a model drug in the DCMC conjugate to sustain its release, elevate its therapeutic index, and try to reduce its toxicity. In vitro and In vivo cytotoxicity studies were done to evaluate the effect of these DCMC nanoparticles. For in vitro assessment, MTT assay was done using human breast cancer cell line (MCF-7), and for in vivo assessment, fiveesix weeks old mice were used. Epirubicin-DCMC nanoparticles (EDNs) could sustain drug

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release. In vivo studies showed that EDNs were more cytotoxic when compared with free Epirubicin. In vivo studies demonstrated that the conjugate did not give rise to any unpredictable outcomes or side effects. It was also seen that the cumulation of epirubicin in the tumor was increased, whereas its absorption in the heart was reduced. The takeaway from this study was that these nanoparticles embedded with epirubicin could suppress the tumor in vivo and could also be a good antitumor drug carrier [24] Many researchers have pointed out that curdlan derivatives such as CMC have very low cytotoxicity. Therefore, the antitumor mechanism of curdlan is not dependent on cell killing but more on immunological regulation [25].

13.5.2 Phosphorylation Phosphorylated curdlan also has great potential. Ionic hydrogels were prepared by chemically cross-linking curdlan and its derivative phosphorylated curdlan. The cross-linking agent used here was 1,4-BDDE, and tetracycline hydrochloride (TCH) was taken as the model drug for release studies as it is a well-known antibiotic against various bacterial infections. The drug release was greatly influenced by the hydrogel compositions. It took about 3.5 h to attain drug release equilibrium. The results were very promising, and these ionic hydrogels could be used in local drug delivery systems [12]. Micro and nanocapsule fabrications which can be used for the encapsulation of different materials are of great interest in the food, biomedical, and pharmaceutical industries. For capsule preparation, several materials have been used such as organic polymers and inorganic compounds among some others. Natural polysaccharides have gained popularity in this field over the last few years. A layer-by-layer assembly was employed to make microcapsules which consisted of hydrolyzed collagen (HC) and curdlan phosphate (PCurd). Negatively charged silica particles were used as the central core around which a layer of HC was coated followed by a layer of PCurd. After the silica core dissolved, hollow capsules were obtained. These microcapsules can be used in medical and pharmaceutical applications for the purpose of encapsulation [26].

13.5.3 Sulfation Another modification of curdlan is curdlan sulfate. Curdlan sulfate has proven to be effective in anti-AIDS therapy, vaccine adjuvants, antitumor immunotherapy, and as drug carrier. The causative agent of acquired immunodeficiency syndrome (AIDS) is the human immunodeficiency virus (HIV). HIV is of two types, that is, HIV-1 and HIV-2. Azidothymidine, dideoxyinosine, and dideoxycytidine are nucleoside analogs which cease DNA chain synthesis from the viral RNA by inhibiting the reverse transcriptase enzyme. But this comes with some extremely serious after effects such as bone marrow toxicity. It must also be noted that it gives rise to the development of azidothymidine-resistant viruses over time rendering them useless in the near future. Sulfated polysaccharides such as curdlan sulfate have shown to have anti-AIDS activity. Curdlan was sulfated to curdlan sulfate with piperidine-N-sulfonic acid. When a concentration of 3.3 mg/mL was achieved, curdlan sulfates completely inhibited the infection. It was expected to act as an anti-AIDS drug which could be used in the early stages of infection for both HIV-1 and HIV-2[27]. Curdlan sulfate has also been recently used to make curdlan sulfate-O-linked quaternized chitosan nanoparticles which could be used as potential adjuvants in the future. The nasal cavity acts as the first line of defense. Usually foreign materials/antigens which are inhaled are presented to the lymphocytes by the antigen-presenting cells (APCs) which will in turn give rise to their differentiation, proliferation, and activation. However, antigens which are nonadherent and soluble are not blocked. The only intranasal vaccine approved for use against influenza, FluMist, has shown to have very low efficiency 46%e58% along with side effects. The way to combat this problem is by utilizing nanosized materials such as nanoparticles, liposomes, and nanogels. These nanosized biomaterials can be used to effectively carry soluble antigens which can be recognized by APCs, thereby promoting strong immunity [28]. Curdlan recognizes the surfaces of immune cells via the dectin 1 receptor, and curdlan sulfate shows good immunoenhancement in vitro. It has shown to elicit strong immune response in mice. However, this polysaccharide can barely be absorbed by the epithelia, thereby limiting its application in mucosal immunotherapy, which is why, in this particular study, chitosan which is a known mucosal absorption promoter has been used. O-HTCC (O-(2-hydroxy) propyl-3-trimethyl ammonium chitosan chloride was selected to be the carrier as well as internalize ovalbumin (OVA), the model antigen. Curdlan sulfate acted as the immunological stimulant and the cross-linking agent for O-HTCC to form CS/O-HTCC nanoparticles. For conducting studies in vivo, female BALB/c mice in the ages between six and eight weeks were used. It was seen that the nanoparticles improved phagocytosis and also heightened the influx of dendritic cells, macrophages into the spleen, which was succeeded by the activation of the APCs. The results were promising, and it showed that these CS/O-HTCC nanoparticles could be intranasal vaccine adjuvants in future [29].

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A novel idea to attack cancer cells can be through dendritic cell vaccine immunotherapy which mainly consists of applying tumor cell lysate (TCL) pulsed with dendritic cells to elicit an antigen-specific response which attacks the cancer cells. But immunostimulants are needed to increase this effect of response. Therefore, curdlan sulfate was used to help mature the TCL-pulsed DCs. It was studied that the curdlan sulfate-DC vaccine elevated the CD8þ T cells, CD80, both MHC I and MHC II expression in tumor tissues and downregulated the TGF-b transcription. It was noted that the vaccine reduced the tumor burden in mice and prolonged their survival. It was concluded that curdlan sulfate (type 3) is worth trying for tumor immunotherapy as it a potential candidate for DC vaccines [30]. Curdlan sulfateechitosan nanoparticles polyelectrolyte complex was developed to study its effect for intracellular drug delivery using Rifampicin. Tuberculosis is mainly caused by Mycobacterium tuberculosis which infects almost 23% of the world population. Mycobacteria reside within macrophages and granulomas. High doses of drugs are used to get rid of these, but this in turn leads to drug resistance. Therefore, to address this problem, curdlan sulfateechitosan nanoparticles were prepared. Rifampicin was encapsulated into these nanoparticles. Its effectiveness was studied on the Mycobacterium smegmatis-infected macrophages. It was noted that drug release was sustained over a given period. Bactericidal activity was observed, and they could target the infected macrophages. This study showed that curdlan sulfateechitosan nanoparticles can be used for macrophage-targeted drug delivery and can also take care of the rising drug resistance problem [31]. Most polysaccharide-derived hydrogels are very weak and fragile in nature due to the absence of antifracture feature which limits their applications. Hence, to upgrade the strength of the polysaccharide-based hydrogels, they were usually cross-linked with other synthetic polymers during preparation [32]. While this method might be good at increasing the strength of the hydrogel, it may also alter the biological function of the hydrogel, and this might greatly affect its application in the real-world scenario.

13.6 Evaluation parameter in animal model for curdlan hydrogel Curdlan hydrogels have been tested for their effect and efficacy in mice and rat models mainly for the purpose of wound healing. An animal study was conducted wherein curdlan/AgNP nanofibril hydrogels were developed for wound healing. This was studied in a mice skin wound model which was infected by bacteria. First, curdlan-stabilized AgNPs were prepared and then curdlan hydrogel which consisted of nanofibrils and AgNPs were made. These silver-embedded curdlan hydrogels were checked for different activities such as antifracture, antiinflammatory, and antibacterial along with the wound-healing study. The cytotoxicity of this particular hydrogel was tested on fibroblast cells. It was seen that the fibroblast cells were easily adhering onto the curdlan nanofibrous hydrogels, which means, it promoted cell affinity and adherence. This could be due to a similar stiffness exhibited by curdlan and collagen. Since curdlan hydrogels have the ability to sustain release, AgNPs are slowly released from the silver-embedded curdlan nanofibril hydrogels. This avoids sudden accumulation of Agþ and reduces cytotoxicity. The cells that were treated with the hydrogels and the untreated cells both showed the same level of viability which confirmed that the hydrogels were not cytotoxic. The hydrogels prohibited the growth of Escherichia coli and Staphylococcus aureus, and the bacterial inhibition rate increased as the Ag concentration in the hydrogels increased. Proinflammatory cytokines such as IL-6, NO, IL-6, COX-2, and TNF-a levels in LPS-induced macrophages were reduced on treatment with the hydrogel, but increased expression of antiinflammatory factors wouldn’t be affected in any way. The antifracture property of this hydrogel was also evaluated. It was seen that the fracture strain exhibited was around 350%, the stress was about 0.2 MPa, and the water was 97%. These values were higher as compared to other polysaccharide-based hydrogels for the same water content. A mice skin model with S. aureus wound infection was used for the study on the effect of hydrogel on wound infection. The wound was subjected to hydrogel treatment and no treatment. Without treatment, the wound would heal itself slowly, but even up to the 10th day, the wound had not closed. However, in the presence of the said hydrogel, the healing process was pronounced, and the wound started closure on the sixth day itself [32]. All these characteristics of this hydrogel make it an ideal material for clinical wound healing. The use of curdlan hydrogels as adhesion barriers was evaluated in postoperative peritoneal adhesion in rats. Hydrogels of curdlan, gellan, and their mixture were used in this study. RAW 264.7 macrophage cells were used for checking the cytotoxicity of these hydrogels. The hydrogels did not show any toxicity up to 1% (v/v) concentration. In fact, cell proliferation was increased on addition of the hydrogel. This shows that the hydrogel can be prudently used in the human body. To study the antiadhesion effect in rats, the hydrogels were used to cover the damaged peritonea and ceca. After seven days, the adhesion levels were evaluated. It was seen that the curdlan/gellan mixture hydrogel was the best at reducing adhesion between the peritoneum and the cecum [33]. The second best at reducing adhesion was found to be

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curdlan hydrogel followed by gellan hydrogel. Therefore, this mixture of curdlan/gellan hydrogel might be a good option for adhesion barrier in postoperative adhesion.

13.7 Application of curdlan-based hydrogel There are several studies that have concentrated on polysaccharide-based hydrogels for applications of drug delivery, scaffolds, and other medical uses. However, work on curdlan-based hydrogels is limited and is also not very well known as compared to other polymeric hydrogels. Nonetheless, we have tried to outline the research in various domains which use curdlan hydrogels. The main target of effective drug delivery is to see that the drug reaches the target site in the body and thereafter provides therapeutic quantity of the drug. It must also provide this effect for the entire duration of the treatment for positive results. Curdlan hydrogels being naturally obtained degrade biologically within the body which eliminates the need to take out a drug delivery system after it has released the drug [34].

13.7.1 Drug delivery application 13.7.1.1 Oral drug delivery The pharmacokinetics of the drug theophylline was studied by in vitro and in vivo by preparing theophylline-containing curdlan tablets. Theophylline is effective against COPD (chronic obstructive pulmonary disease). It helps in clearing the airway and breathing easier. Two different types of tablets were prepared based on the tablet surface area. One had a larger surface area (Tab L) whereas the other one had a smaller surface area (Tab S). Tab L had a faster drug release rate of below 24 h as compared to Tab S which was around 32 h. This was probably due to the larger surface area of Tab L. In addition to the in vitro study, the same was checked in vivo too in five healthy volunteers. The tablets were administered orally to the volunteers, and it was observed that the drug release profiles were very similar to that of the in vitro drug release [35]. The reduced bioavailability of Tab S advocated that complete drug release wasn’t achieved at the time of gastrointestinal passage. In conclusion, though curdlan improved the controlled release of theophylline, Tab L was the better choice due to its larger surface area. Indomethacin is an NSAID which is used against osteoarthritis, rheumatoid arthritis, gout, and joint pain; prednisolone is used for certain allergic conditions, breathing problems, and arthritis; and salbutamol sulfate is used to treat wheezing and other COPD-related problems. Curdlan gel was also utilized for the encapsulation of certain drugs like indomethacin, prednisolone, and salbutamol sulfate. This was done by first preparing gel suppositories followed by conducting in vitro drug release studies. Curdlan was added to the drug suspension to get a drug-containing curdlan suspension. Different concentrations of drug solutions were prepared, and the suppository was made. The fastest release was seen for salbutamol sulfate followed by prednisolone and indomethacin. In conclusion, curdlan gel was suitable for sustained drug release and it was found to be diffusion-controlled [36]. This facilitates the bioavailability in the lower rectum and bypasses the firstpass mechanism by the liver when taken orally or injected [11]. Curdlan gum was employed to develop a regioselective drug delivery system for lamotrigine, an antiepileptic drug. For this purpose, floating lamotrigine tablets were made which consisted of curdlan gum and HPMC K100M. The tablets showed desirable release properties in vitro. One of the formulations showed drug release for 12 h which made it a suitable alternative to the conventional therapy [37].

13.7.1.2 Dental drug delivery Curdlan hydrogels have also gained importance in dental field wherein they have been employed for the treatment of periodontal problems. In the last two decades, the use of polymers in dental drug delivery systems has picked up pace due to polymer properties which make them ideal candidates. The polymer in question must be able to support the drug in action or control the release kinetics. Periodontitis is an inflammatory disease which effects the periodontal tissue due to pathogenic microbes. It is characterized by tearing down the supporting teeth structures [38]. The bacteria that are responsible for causing the destruction of these supporting tissues are mainly Porphyromonas gingivalis, Tannerella forsythia, Fusobacterium nucleatum, etc. The most common formulation for local administration in the buccal cavity is liquid or semisolid because they are easy to administer. However, it comes with a disadvantage of poor retention, and hence the therapeutic effect is suboptimal. To overcome such disadvantages, polymers have been used. A functionalized curdlandpolydopamine (PDA) composite hydrogel was made for antimicrobial delivery against periodontitis [17]. Antibacterial activity was checked using S. aureus and E. coli, whereas the antimicrobial chosen was acetate chlorhexidine (CHX). It was observed that there was synergistic antibacterial activity, and the bacteriostatic rate was almost 99.9%. The

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local release of the drug further enhanced the bactericidal activity. These results are very promising for dental patients and, therefore, in vivo studies are need of the hour.

13.7.1.3 Transdermal drug delivery Curdlan hydrogels have been tried as a remedy for bacterial wound infections as well. When a person is wounded, he/she may develop bacterial infections at the site of the wound. Such bacterial infections are mainly caused by microbes such as S. aureus, P. aeruginosa, E. coli, and E. faecalis which may give rise to improper wound healing and may also lead to sepsis and death in some serious cases. Managing bacterial colonization and infections requires the use of appropriate antibacterial dressing. Such dressings should be nontoxic, nonadherent, and nonallergic to the human body. They must also have the capability to absorb exudates from the wound and maintain an environment with moisture at the wound site. But, more importantly, bacterial infections at the wound should be fended off [39]. Therefore, dressings are always made better by the addition of antimicrobials, antiseptics, and metals. Commonly used metal ions are silver, copper, and zinc. Among these, copper has shown great wound-healing ability. Not only does it have a wide spectrum of antimicrobial activity but it also promotes wound healing and stabilization of skin proteins [40]. Hence, curdlan-based hydrogels augmented with copper ions were made as dressing materials to evaluate their potential against bacterial colonization in wounds. The curdlan-based hydrogels with copper ions were made by two methods viz. ion-exchange dialysis against copper ions and freeze-drying. The biomaterial had a good porous structure, absorbed the wound exudate, and showed the ability to release good amounts of copper ions into the surrounding environment. Moreover, it also reduced the growth of the test bacterial strains, S. aureus and E. coli [41]. Curdlan-based hydrogel incorporated with copper ions was compared with the commercially available dressingdKALTOSTAT . It showed similar absorbent properties. But it possessed much better fibroblast viability than KALTOSTAT [42]. Therefore, it can be deduced that such hydrogels can be a good alternative to the already existing dressing materials. Curdlan-agarose hydrogels with gentamicin have also been prepared to study their effect as wound-dressing biomaterial on highly exuding wounds in a rat. To avoid any damage to the biomaterial, the dressing was sewn into the abscess cavity. After seven days, it was seen that the dressing fell off along with the scab. After 10 days, reepithelialization process was complete and on the 14th day, remainder scab fell off, and skin healing was completed [43] This biomaterial was not only nontoxic to human fibroblast cells but also showed antibacterial activity against S. aureus and P. aeruginosa. It also reduced scarring in the animal. Nanofibers of curdlan along with polyethylene oxide (PEO) have been made by the process of electrospinning. Curdlan needs to be blended with other polymers for successful electrospinning. PEO is nontoxic, biocompatible, and costeffective. Hence, it was used to prepare nanofibers along with a cross-linking agent called glutaraldehyde (GA). The polymer blend was then loaded with a model drug, TCH. TCH can be used to control infections and also be useful in skin and bone-related applications. The noncross-linked nanofibers initially showed a burst of TCH in 2 h then in 5 h followed by full release, whereas the cross-linked nanofibers showed a slower release rate of up to 24 h. It was also seen that the cross-linked TCH-loaded curdlan/PEO showed an effective antibacterial effect against E. coli growth [44] Similarly, composite hydrogels of curdlan and b-cyclodextrin have been prepared with epichlorohydrin (ECH) and ethylene glycol diglycidyl ether (EGDGE) as cross-linkers and loaded with sodium salicylate for release studies. b-cyclodextrin is also a natural polymer which is produced as a result of starch hydrolysis. Sodium salicylate is used as an antipyretic and also as antirheumatic drug. This composite hydrogel could sustain sodium salicylate release for more than 8 h. It was also observed that the hydrogels prepared with EDGE had better activity than the ECH hydrogels. The addition of b-cyclodextrin significantly improved performance of the hydrogel [18].

13.7.1.4 Protein delivery vectors Curdlan hydrogels have also shown to be effective as protein delivery vectors. Proteins also show good therapeutic effects, but due to their large size, enzymatic degradation, and other problems such as chemical and physical instabilities, their activity gets affected. Hydrogels have been studied for their capability to carry proteins to increase their bioavailability and stability. Histidine is an amino acid which has side groups with pH-responsive properties. Therefore, pH-responsive hydrogels were made from curdlan derivatives and Boc-histidine. These hydrogels were loaded with BSA to study protein drug delivery. The protein drug release rate improved with increase in pH, that is, at almost neutral pH of 7.4, the release rate was increased as compared to pH 4. Cytotoxicity evaluation was also done for these hydrogels by increasing the concentration. It was seen that viability of the HepG2 cell line went from 106% at 50 mg/mL to 90% at 800 mg/mL [45]. Hence, it can be said that the cytotoxicity of the curdlan-histidine hydrogels was very low, which makes them a potential candidate for biomedical applications.

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Exopolysaccharides such as curdlan are used in the preparation of hydrogels because of their advantageous biological properties. Moreover, they can be chemically modified to make curdlan hydrogels water soluble. Therefore, the purpose of the said work was to prepare curdlan and phosphorylated curdlan hydrogels by using 1,4-BDDE as a chemical cross-linker. TCH was chosen for drug loading as well as release studies. The drug release greatly increased from 48% to 87% when phosphorylated curdlan was increased and equilibrium was reached after 3.5 h. These results show that ionic hydrogels can be successfully utilized in local drug delivery systems [12].

13.7.1.5 Cryogels Apart from hydrogels, attempts have been made to prepare cryogels from curdlan. Cryogels are one of the lightest solid materials available. It is made from hydrogels as the precursor wherein the liquid in the hydrogel is freezed below glass transition or melting point, later succeeded by sublimation. Two types of cryogels were prepared from curdlan hydrogels. One was with PEO and other included PEO/cellulose nanofibrils. The release of diclofenac sodium from these cryogels was studied, and it was noted that release was better with the curdlan/PEO/CNF cryogel [44] This study might be the starting point of a new field of curdlan gels which broadens the scope of applications for curdlan.

13.7.2 Tissue engineering application Curdlan’s rheological properties make it an exemplary candidate for use in tissue engineering studies. However, not many reports have been published in this stream.

13.7.2.1 Skin tissue regeneration Curdlan and polyvinyl alcohol (PVA) were chemically cross-linked using formaldehyde to form 3D scaffolds. PVA was deliquesced in deionized water and then heated. It was stirred at 300 rpm at 90
C for 30 min until a colloid was formed. After the PVA solution was reduced to room temperature, curdlan solution was poured and mixed followed by the addition of H2SO4 and formaldehyde. This sample was then added to a tube and placed in a hot air oven at 80
C for 50 min. These formed 3D scaffolds were washed in distilled water to take off any residues. In vitro and in vivo studies took place with this scaffold. Scaffolds were subcutaneously embedded in SpragueeDawley rats. After 14 days, these rats were sacrificed in order to remove the scaffolds from within. There was no issue on the skin of the rats. The scaffold was covered using dermal tissue, and there was no significant inflammation seen. Histology also revealed that there was no dissimilarity between the control and the implanted rats. The scaffold had very good decomposability and was biocompatible. These findings confirm that curdlan/PVA scaffolds can be used in future tissue engineering studies [46]. Fabricated curdlan hydrogels were made by cross-linking using EGDGE as the cross-linker. Furthermore, curdlan organogels were formed by acetylation of curdlan hydrogels. It was found that curdlan hydrogels were soft, flexible, and tough. For polysaccharide-based hydrogels, curdlan is the only one which has all three properties concurrently. For example, agar gels are soft but not flexible or tough. Organogels were impervious to continuous drying and swelling. Curdlan hydrogels would be useful as artificial soft tissue analogs in tissue engineering studies, whereas organogels can be useful in a variety of soft material science [47]. Ion-exchange dialysis method was employed to entrap model protein BSA in curdlan hydrogels to check if the hydrogel can be an effective protein carrier for delivery. It was seen that curdlan hydrogel had very high entrapment efficiency for proteins. It exhibited a very orderly release of BSA up to four weeks, without altering the properties of the protein. Therefore, curdlan hydrogels seem to be a potential protein delivery system in tissue engineering research [48].

13.8 Conclusion Curdlan, as mentioned in this chapter, has a multitude of applications. However, because of its water insolubility, it would seem like its applications could be curtailed. Nonetheless, either by chemical modifications curdlan can be made water soluble or by addition of appropriate amount of NaOH. Curdlan or its hydrogels are not very well known, but it has now started gaining pace and importance in the pharmaceutical and biomedical sectors as outlined in this chapter. As a carrier for drug delivery and for applications in tissue engineering, it is a very suitable option because of its varied properties such as biodegradability, biocompatibility, and nontoxicity. One of its most promising features is that it can control and sustain drug release which would be of great benefit to the biomedical field when patients are concerned. It might also be an alternative to cartilage and soft tissues, but more studies have to be done with respect to this aspect. Therefore, in conclusion, curdlan hydrogels hold a very broad and positive scope for the future of drug delivery and tissue engineering.

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Chapter 14

Chitosan nanogel for drug delivery and regenerative medicine ¨ zhan Aytekin Neslihan Kayra and Ali O Genetics and Bioengineering Department, Engineering Faculty, Yeditepe University, Istanbul, Turkey

14.1 Introduction 14.1.1 Chitosan Chitin is a linear polysaccharide, polymer of b-(1e4)-N-acetyl-D-glucosamine monomers and found in cell walls of fungi and algae, exoskeleton’s of insects, mollusks, and crustacean’s shells (Fig. 14.1). Chitosan (CS) is a linear polysaccharide and N-deacetylated form of chitin, composed of b-(1e4) linked 2-amino-2 deoxy-D-glucose monomers. Chitosan possesses hydroxyl and amino groups due to the glucosamine unit in the chain of chitosan. These groups are crucial for crosslink reactions because they are reactive sites of chitosan [1]. Because the hydroxyl and amide groups can facilitate the formation of hydrogen and covalent bonds [2]. Additionally, it is the only polycationic polymer in nature [3]. This feature presents easy binding with negatively charged compounds without chemical reactions [1]. Although chitin is not soluble in organic solvents and water due to the strong intermolecular hydrogen bonds, while chitosan can be soluble in aqueous acid solutions; because of free amino groups. Chitosan is a weak poly-base with around 6.5 pKa value, so its charge density can alter in the pH range between 6 and 6.5. Because of this particular characteristic, chitosan can respond to alteration of pH in the environment and soluble and insoluble forms of the chitosan can be obtained at pH between 6 and 6.5 [4]. While the high charge density of chitosan causes polyelectrolyte formation at pH levels below the pKa, the low charge density of chitosan at neutral pH conditions results in low toxicity and contribution to the release of intracellular biomolecules [4].

FIGURE 14.1 Chitosan structure and preparation from biological resources. revised from Tian B, Liu Y, Liu J. Chitosan-based nanoscale and nonnanoscale delivery systems for anticancer drugs: a review. Eur Polym J 2021;154:110533. https://doi.org/10.1016/j.eurpolymj.2021.110533 and permission will be taken from Tian B, Liu Y, Liu J. Chitosan-based nanoscale and non-nanoscale delivery systems for anticancer drugs: a review. Eur Polym J 2021;154:110533. https://doi.org/10.1016/j.eurpolymj.2021.110533. Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine. https://doi.org/10.1016/B978-0-323-95351-1.00018-1 Copyright © 2024 Elsevier Inc. All rights reserved.

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Biodegradable, biocompatible, and nontoxic properties of chitosan increase the opportunity of usage in various fields such as in the food industry in pharmaceutical for drug delivery systems, in biomedical for artificial skin and bones [3], and in wastewater treatments systems as coagulants [5]. Recently, among these applications, nanogel form of chitosan for drug delivery and regenerative medicine has acquired attraction because of its outstanding characteristics of chitosan. For example, the cationic nature of chitosan provides mucosal adhesion [2], controlled drug release, in situ gelation, and improvement in permeation [6].

14.1.2 Chitosan nanogels Nanogel is a class of hydrogels that are a three-dimensional hydrophilic network of polymers that can be produced by chemical or physical cross-linking. Nanogels differ from hydrogels in terms of their size, stability, and responsiveness [8]. The nanosized properties, which are approximately between 1 and 100 nm or 20 and 200 nm, increase the chance of usage of nanogels in numerous areas in contrast to hydrogels. Nanogels can be considered as a combination of hydrogels and nanoparticles. Nanogels can be made from synthetic and natural polymers. Among these natural polymers, polysaccharides are promising materials because of their enhanced biodegradability. As well as, biocompatibility, low cost due to its abundancy in nature, lox toxicity, and functionality are other important features of polysaccharides that make them attractive for nanogel formulations. In this regard, among these polysaccharide-based biopolymers, chitosan and its derivatives are promising polymers for nanogel formulations.

14.1.3 Chitosan nanogel preparation methods Nanogels can be formed via different formation methods. Cross-linking can be achieved via physical cross-linking and chemical cross-linking ways [9]. While physical cross-linking is formed with the help of electrostatic and hydrophobic interactions, noncovalent bonds, and hydrogen bonds, chemical cross-linking is achieved through covalent bonds between functional groups of polymer matrix and cross-linker agent [10]. Because of the strongness and weakness of links, physically cross-linked formed networks can be easily broken, and chemically formed networks are more durable structures. Cross-linking is an important parameter for nanogels that are used in drug delivery systems because they prevent the dissolution of networks in water-based systems [11]. Apart from protection from dissolution, covalent and noncovalent bonds between drugs and polymeric networks affect the release rate from nanogels in drug delivery systems [12]. Behind the cross-linking method, there might be several factors that affect the drug delivery systems, such as the molecular weight of the polymer, cross-linking quantity, and interaction of drugs within the polymer chain [12]. For example, chitosan nanogel has been fabricated from chitosan nanoparticles and nanocomposite form laponite by ionic gelation method in the presence of tripolyphosphate (TPP) for the delivery of honey as a model drug [13]. The one-pot free radical polymerization method has been used to formulate pH-sensitive chitosan and kappa-carrageenan-based nanogel to deliver rivastigmine as a model drug [14]. Then graft polymerization of the acrylamide/sodium acrylate monomers on chitosan and kappacarrageenan backbone has been achieved in the presence of N, N’-methylenebisacrylamide as a cross-linking agent, and nitrogen-doped carbon dots and ammonium persulfate as initiator. The effect of the monomer concentration, crosslinker concentration, and initiator concentration on the swelling behavior of the chitosan and kappa-carrageenan has been evaluated. Instead of the usage of one kind of preparation method, a combination of them might be used to eliminate the drawbacks of each method and produce more stable nanogels. For instance, a nanogel has been produced from phenolic hydroxyl-modified chitosan for the delivery studies of the 5-Fu as a model drug by using ionic and enzymatic cross-linking methods together [15]. For the production of multiple cross-linked chitosan nanogel, firstly modification of chitosan with phloretic acid has been achieved in the presence of 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride (EDCHCL) and N-hydroxysuccinimide (NHS). Then phenolic hydroxyl-modified chitosan has been used for the nanogel formation by using sodium TPP and horseradish peroxidase for ionic gelation and enzymatically cross-linking, respectively. And also, the ionic gelation contribution of sodium TPP has been compared with sodium molybdate. The multiple crosslinked phenolic hydroxyl-modified chitosan nanogel has shown 35.01% drug loading and 66.82% encapsulation effectiveness. It has been aimed that producing chitosan nanogel by photocross-linking of poor solvent-induced nanoaggregates as a simple approach (Fig. 14.2) to overcome side effects of chemical initiator and cross-linking agents and required complicated conditions for the chemical methods [16]. For this purpose, firstly, modification of carboxymethyl chitosan has been achieved by using o-nitrobenzyl alcohol and then self-assembled of modified carboxymethyl chitosan has occurred in ethanol-water mixture solution containing uncross-linked nanoaggregates under 365 nm light irradiation for a few minutes because of UV-induced cyclization between primary amine and o-nitrobenzyl alcohol. Photoclick produced modified carboxymethyl chitosan nanogel has been decorated with lactobionic acid for loading of doxorubicin. The

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FIGURE 14.2 Schematic illustration of chitosan-based nanogel formation by photoclick reaction for the delivery of tumor-targeting drug [16]. permission will be taken from Lu DQ, Liu D, Liu J, Li WX, Ai Y, Wang J, et al. Facile synthesis of chitosan-based nanogels through photo-crosslinking for doxorubicin delivery. Int J Biol Macromol 2022;218:335e45. https://doi.org/10.1016/j.ijbiomac.2022.07.112.

evaluation of the nanogel has confirmed that the production of nanogel with photocross-linking method is an alternative way to chemical methods with drug delivery ability.

14.2 Chitosan nanogels in drug delivery Nanogels have been favored in various application areas like food, nutrition, pharmaceutical, biomedicine, biotechnology, and so on. Among these areas, most investigated and popular ones are drug delivery systems and regenerative medicine. Nanodrug delivery system is preferred in new drug formulations to eliminate the limitations of common drug delivery systems. Because of a lot of inherent features of nanogels, they are an excellent candidate for drug delivery and regenerative medicine. Nanogels are accepted as next-generation drug delivery systems because of their outstanding properties like uncomplicated preparation, biocompatibility, tunable size, swelling ability, hydrophilicity, and stimuli responsiveness to environmental changes [17], high stability, and drug-loading capacities [18]. For example, due to the large surface area of nanogels, they can penetrate easily through tissue barriers [19]. Also, nanogels can be modified chemically by incorporating functional molecules to improve their carrier properties in different delivery systems, such as targeted drug delivery, stimuli response drug release, and to formulate composite nanogels with other polymers [12]. The presence of hydrophilic functional groups like amino, carboxylic, sulfonic, hydroxyl, and so on in the structure of nanogels enhances their stability and colloidal dispersion [8]. Because of the rapid responsiveness of nanogels, they are effective materials for biomedical applications [19] like drug delivery and regenerative medicine. For example, they can shrink or swell based on the change in external environment conditions [18]. Cross-linking structure of nanogels can be used for the encapsulation of hydrophobic drugs to enhance their stability during storage or when circulating to reach the target system, and additionally small size of the nanogels presents a high specific surface area for bioconjugation of targeting molecules [18]. Besides the abovementioned properties, special thermodynamic flexibility, good solubility, strong viscosity, and competence in sterilization-like features make nanogels satisfactory carriers for drug delivery systems [12]. Additionally, carbon and water molecules that are present in the structure of nanogels provide compatibility with biological tissue and improve the biocompatibility and biodegradability properties of nanogels [12]. Besides general reasons for the usage of nanogel

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systems in drug delivery and regenerative medicine, some unique features make chitosan favorable for nanogel construction of drug delivery and regenerative medicine. For instance, the cationic nature of chitosan is the best choice for the controlled release of anionic drugs, and also because of its cationic nature, it shows mucoadhesive properties [20]. Chitosan has in situ gelling, transfection-enhancing, and permeation-enhancing properties [20], and these characteristics make chitosan attractive for drug delivery systems. Moreover, the modification opportunity of the chitosan improves the solubility of the chitosan other than in the acidic environment and also contribute the formation of stimuli-responsive chitosan [21]. Also, biocompatibility, biodegradability, and bacteriostatic properties are outstanding features of the chitosan to be preferred as a polymer in drug delivery systems [21]. Classification of nanogels might be made by considering several properties like polymer type used like a native, blend, modified, functionalized, and so on, or linkage type used in a formulation like covalent or noncovalent, type of responsiveness like pH, temperature, magnetic, structure like multilayered, core-shell, hollow, hairy, and so on [17]. The number and the diversity of the classification criteria can be increased according to the subject. In this chapter, the recent chiton nanogels studies in drug delivery are evaluated by considering chitosan type: native chitosan, the derivative of chitosan, functionalized chitosan, and blend of chitosan with other polymers. Other than the type of chitosan, stimuli responsiveness, temperature, pH, and dual responsiveness of chitosan nanogels in drug delivery are evaluated. Additionally, drug delivery systems of the chitosan nanogels are examined according to administrative routes. In the case of regenerative medicine, the chapter focused on chitosan nanogel studies formulated for wound healing and dentistry applications.

14.2.1 Chitosan-based nanogels Chitosan nanogel has been prepared to encapsulate Artemisia scoparia extract to enhance the cytotoxicity of the extract against the hepatocarcinoma cell line (Huh-7) [1]. Nanogel formation has been achieved between free amine groups of chitosan and carboxylic groups of myristic acid (MA) which has been used as a cross-linker, and the reaction has been mediated by 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC). The anticancer activity and controlled release of A. scoparia-encapsulated CS-MA spherical nanogel with a dimension less than 100 nm have been evaluated. The results presented that extract-loaded nanogel has led to apoptosis in the Huh-7 cell line, and it has been concluded that A. scoparia-encapsulated CS-MA nanogel could be an alternative for cancer treatment. Chitosan-based nanogel has been formed from high molecular chitosan (HCS) by using sodium TPP as a cross-linker to encapsulate resveratrol (RSV) which is a nonflavonoid polyphenol [6]. The formed HCS nanogels were in nearly 140 nm in size and had 32
2 mV Ƶ-potential value. While CS nanogel showed 24
10% encapsulation efficiency of RSV, the HCS nanogels presented 59
1% RSV encapsulation efficiency. Additionally, cytotoxicity studies in human retinal pigment epithelial cells (ARPE-19) concluded that HCS nanogels were biocompatible with ocular drug delivery systems. Another chitosan nanogel formation has been achieved with the aim of encapsulation of polyoxometalates (POMs) as an antitumor agent for breast cancer therapy [22]. For this purpose, WellseDawson-type [P2Mo18O62]6- phosphomolybdate has been encapsulated chitosan nanogel formulated by reverse microemulsion technique in the presence of poly(ethylene glycol) bis(carboxymethyl) ether (PEG(COOH)2). Cross-linking has been achieved between eCOOH groups of (PEG(COOH)2) and amine groups of chitosan. It has been shown that chitosan nanogel presented nearly 90% loading efficiency of POM, whereas the release of POM from chitosan nanogel was 30%.

14.2.2 Chitosan derivative-based nanogels Despite the unique chemical structure of chitosan still, some limitations restrict the usage of chitosan in biomedical applications such as drug delivery and regenerative medicine. The most critical disadvantage of chitosan is poor solubility in neutral aqueous solutions and organic solvents. While chitosan is soluble in weakly acidic solutions, it is insoluble in water. Additionally, the solubility of chitosan can be affected by the molecular weight of chitosan [23]. Therefore, chemical modification of chitosan is required to improve deficiencies by producing derivatives of chitosan. Derivatives of chitosan (Fig. 14.3) are produced by modifying reactive functional groups, like amino groups, of chitosan, and in contrast to native chitosan, these derivatives are water-soluble and have low molecular weight [23]. For instance, trimethyl chitosan (TMC) is an example of a chitosan derivative that is soluble in neutral and alkaline solutions, carboxymethyl chitosan is an example of a pH-dependent water-soluble chitosan derivative, thiolated chitosan (TC) is another chitosan derivative obtained by the attachment of thiol groups (-SH) to chitosan structure, and glycated chitosan which is a water-soluble derivative is produced by the reaction between galacturonic acid and chitosan [24]. Besides native chitosan, these derivatives with enhanced properties can be used to formulate nanogel for drug delivery and regenerative medicine. The modification of the chitosan with hydrophilic or hydrophobic possessing compounds

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FIGURE 14.3 Structure of chitosan derivates obtained after chitosan modification. revised from Iacob AT, Lupascu FG, Apotrosoaei M, Vasincu IM, Tauser RG, Lupascu D, et al. Recent biomedical approaches for chitosan based materials as drug delivery nanocarriers. Pharmaceutics 2021;13:1e36. https://doi.org/10.3390/pharmaceutics13040587 and Permission will be taken from Iacob AT, Lupascu FG, Apotrosoaei M, Vasincu IM, Tauser RG, Lupascu D, et al. Recent biomedical approaches for chitosan based materials as drug delivery nanocarriers. Pharmaceutics 2021;13:1e36. https://doi. org/10.3390/pharmaceutics13040587.

facilitates the entrance of the drug to the system via different administration routes and improves controlled release behavior and bioavailability of the drug [7]. For example, carboxymethyl chitosan has been used with chitosan for the formulation of nanogel to deliver rifaximin (RFX) which is an antibacterial agent against both gram-positive and gramnegative anaerobic and aerobic bacteria [26]. The formation of nanogel structure has been achieved by ionic gelation and electrostatic interaction between NHþ 3 groups of neat chitosan and COO groups of carboxymethyl chitosan. Characterization of the RFX-encapsulated chitosan nanogel showed that they were in 170 nm size with 0.17e0.42 PDI range. The produced nanogel had higher swelling ratio at acidic conditions (pH 5.5, 130.71%) than at alkaline pH (pH 7.4, 81%). The swelling behavior of the RFX-loaded chitosan nanogel was parallel to in vitro release behavior, which was 76.37% and 53.80% within 72 h for pH 5.5 and 7.4, respectively. Also, mucoadhesive strength and hemocompatibility results of the RFX-chitosan nanogel presented that chitosan nanogel was appropriate for the delivery of RFX. The sulfonated chitosan (SCS) has been obtained by introducing sulfate ions to the main chain of CS to enhance the adsorption of Agþ [27]. Nanogel has been constructed between SCS, Agþ, and chitosan by electrostatic interaction and sodium borohydride reduction method [27]. The formulated AgNPs@CS/SCS has been evaluated in terms of the release of Agþ as antibacterial agent and outcomes noticed that AgNPs@CS/SCS present long-acting and slow release of Agþ. Nanogel of TC has been formed by water-in-oil (W/O) emulsion technique in the presence of poly(ethylene glycol) bis(carboxymethyl) ether (PEGBCOOH) as a cross-linker agent [19]. After, the formation of nanogels labeling of nanogel with folate by in situ functionalization has been conducted to click folate moiety to chitosan with the aid of thiol/maleimide reaction. The aim of labeling CsSH/PEGBCOOH nanogels with folate was to increase the recognition of drugs by only tumor cells. It has been reported the incorporation of thiol moieties enhanced the mucoadhesive properties of nanogels and also enabled the

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labeling of CsSH/PEGBCOOH nanogels with folate to use in target drug delivery studies. An example of a chitosan derivative, carboxymethyl chitosan-based nanogel has been engineered as a drug delivery system by checking the impact of two different cross-linker agents which were acid-labile cyclic ortho ester compound with dual epoxy end groups and ethylene glycol diglycidyl ether (EDGE) on nanogel structure [28]. The nanogel formulation has been achieved by the emulsion solvent evaporation method, and doxorubicin hydrochloride has been used as a model drug. While the acid-labile cyclic ortho ester linker was used to produce pH-sensitive nanogels and EGDE was used for the formulation of nonsensitive nanogels. It was stated that nanogels that linked with acid labile ortho ester linkers were more effective to deliver DOX into the tumor-like multicellular spheroids (MCTSs). Also, acid-degradable nanogels have been found that could be degradable at mildly acidic conditions but stable under physiological conditions. Like previous examples, carboxymethyl chitosan has been used for the formulation of valproate-D-nanogel by the means of copolymerization of carboxymethyl chitosan with diallyl disulfide (D-nanogel) and then grafting with valproate (valproate-D-nanogel) [29]. The research aimed to overcome cisplatin resistance in cancer treatment and improve early apoptosis. It has been found that cisplatin-loaded valproate-D-nanogel was effective to decrease the systemic toxicity of cisplatin and inhibit cisplatinresistant cancer. A recent study has been interested in the formulation of carboxymethyl cellulose-based chitosan nanogel formulation for the delivery of doxorubicin [30]. First, the modification of carboxymethyl chitosan has been achieved via glycidyl methacrylate, then polymerization of the double bond on the side group of carboxymethyl chitosan has been done, and then carboxymethyl chitosan cross-linked with NeN’-bis(acryloyl) cysteamine (BAC) and modified with folic acid. The nanogel has been fabricated from this folic acid-modified carboxymethyl chitosan and has been encapsulated with doxorubicin. The produced nanogel has been evaluated against stimuli of pH and glutathione which are crucial cancer cell-stimulating factors in most tumors. It has been demonstrated that formulated nanogel was effective for doxorubicin encapsulation and had high pH and glutathione responsiveness.

14.2.3 Functionalized chitosan-based nanogels The grafting of functional compounds on chitosan molecules enhances its properties and increases its functionality. For instance, a targeted drug delivery study has been designed for the controlled release of 5-fluorouracil (5-FU) which is an anticancer drug [8]. It was stated that in the core-shell system, the core part acts as a magnetic heater, whereas the shell acts as a drug-loaded part [8]. The produced magnetic nanogelic core-shell consisted of Fe3O4 core and chitosan-polyacrylic acid-based nanogelic shell. The formation of a nanogel core-shell has been done via in situ polymerization of acrylic acid on Fe3O4-chitosan in the presence of glutaraldehyde as a cross-linker. The results of the study indicate that Fe3O4 core structure enhanced the release of 5-FU five times under tumor tissue conditions (pH 4.5) than in physiological conditions (pH 7.4). A similar core-shell nanogel has been produced from gold nanoparticles and chitosan for the loading of curcumin [31]. It has been stated that the conjugation of gold nanoparticles on therapeutic agents improves the characteristics of drug delivery system that are already owned [31]. The formed gold-chitosan core-shell nanogels were in 36.85 nm particle size with 0.27 PDI value. The drug-loading and encapsulating efficiency of curcumin to gold-chitosan core-shell nanogel were 11% and 78.66%, respectively. Also, it has been indicated that the release of curcumin under an acidic pH environment (5.3) than neutral pH (7.4) was obvious. Another functionalized chitosan nanogel has been produced via folic acid, and fabricated nanogel has been evaluated for the codelivery of paclitaxel and curcumin [32]. For the construction of folic acidconjugated chitosan nanogel, initially pluronic P123 has been activated by p-nitrophenyl chloroformate (NPC), and then produced compound has been partially substituted with 3-amino-1-propanol and then conjugation with chitosan has been achieved to obtain chitosan-pluronic P123 copolymer. Finally, the chitosan-pluronic copolymer has been conjugated with folic acid, and nanogel formulation has been achieved by self-assembled method. The folic acid-conjugated chitosanpluronic P123 nanogel has shown 98.63
0.42 and 97.82
0.48 loading capacity for paclitaxel and curcumin, respectively. The anticancer activity of the drugs loaded folic acid-conjugated chitosan-pluronic P123 nanogel has been evaluated for paclitaxel and a combination of paclitaxel and curcumin individually. Outcomes of the anticancer activity on human breast cancer showed that nanogel loaded with paclitaxel and curcumin has been more effective because of the synergistic effect of drugs. Additionally, similar folic acid-conjugated chitosan nanogel has been constructed to be used as cargo for the fluorescent nanoparticles, which are imaging agents for targeting tumor cells [33]. First of all, fluorescent green carbon dots have been produced from grapefruit extract via hydrothermal reaction and used as a cargo system for antimiR-21 to image targeted SKOV3 cells, and then the encapsulation of these green carbon dots by folic acid-conjugated chitosan has been performed. Similar to the previous study, the nanogel has been formulated from conjugated chitosan to load nanoparticles for imaging cancer cells [34]. Firstly, chitosan modification has been done with polyacrylic acid, and then modified chitosan has been conjugated with cystine-functionalized gold nanoparticles to form nanogel, and lastly, doxorubicin has been loaded as a model drug.

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14.2.4 Chitosan nanogels with other polymers There have been numerous studies on a blend of chitosan with other biopolymers for nanogel formation. For example, CS and sodium alginate (SA)-based nanogels have been designed as a nanocarrier for the delivery of timolol maleate (TM) to the cornea to treat glaucoma which causes irreversible blindness [35]. CS-SA nanogels have been produced by the pregelation method in spherical form with 80e100 nm size range. The TM-loading and encapsulation efficiency of CS-SA nanogels were 93.96% and 41.66%, respectively. The outcome of ex vivo permeation studies revealed that the permeation of TM from CS-SA nanogels more than eye drops was high, and also in vivo results support this outcome. It has been stated that research was first for the ocular delivery of TM to the cornea, and it has been found that suitable for this aim. Another blend of nanogel was formed between carboxymethyl starch (CMS) and chitosan hydrochloride (CHC) for the encapsulation and delivery of curcumin [9]. Cross-linking of nanogel was achieved through amide linkages between carboxyl groups of water-soluble CMS and amino groups of water-soluble CHC in the presence of EDC as a cross-linker agent. The optimal CHC-CMS nanogel was performed at 2:1 ratio (CHC:CMS) with 378.2 nm diameter size and 40.23 mV zeta potential value. The results of the research demonstrated the high encapsulation efficiency of curcumin, which was between 89.49% and 94.01% and high sustained release in gastrointestinal fluid. Another curcumin nanocarrier was produced from cross-linking of CS and carboxymethyl konjac glucomannan (CMKGM) in the presence of and EDC/NHS [36] (Fig. 14.4). Cross-linked nanogel was formed between negative groups of CMKGM and positive groups of CS. The presence of EDC/NHS as a cross-linking agent did not affect the particle size of nanogel, but it led to a decrease in zeta potential value. Additionally, cross-linked CMKGM/CS nanogel enhanced the stability and encapsulation efficiency of curcumin and performed controlled and pH-dependent release behavior under simulated gastrointestinal conditions. A nanogel consisting of CS and hyaluronan (HA) has been designed as a nanocarrier of methotrexate (MTX) and 5aminolevulinic acid (ALA) for the treatment of psoriasis, which is a chronic skin disease [37]. It has been aimed that coloading of MTX and ALA to decrease their limitations, which were systemic toxicity and poor permeation, respectively. It has been stated that CS/HA nanogel that loaded with both MTX and ALA has been not been studied before, and results of the research indicated MTX-ALA nanogel was a suitable candidate for psoriasis treatment. An electrostatic interaction and cross-linking by genipin have facilitated the formation of a nanogel comprising glycol chitosan (GC) with hydrophilic ethylene glycol branches and fucoidan (Fu), an anionic sulfated polysaccharide rich in fucose. [38]. The formulated GC/Fu nanogel has been used as a carrier for the antiinflammatory peptide KAFAK. The KAFAK-loaded GC/Fu nanogel exhibited 14.0
2 mV zeta potential value and encapsulation efficiency, and drug-loading values of GC/Fu nanogel were 80.99
0.49% and 9.53
0.06%, respectively. It has been stated that KAFAK-loaded GC/Fu nanogel was found effective in the down-regulation of the secretion level of inflammatory factors (IL-6 and TNF-a) and expression of chondrogenic markers. The results of the study supported that KAFAK-loaded GC/Fu nanogel might be a therapeutic agent for osteoarthritis (OA) intervention. The double-walled poly lactic-co-glycolic acid (PLGA) and chitosan nanogel (PCNGL) have been formed by solvent evaporation technique to decrease the disadvantages of PLGA-like side effects and cell-specific interactions [2]. The designed PCNGL has been assessed for controlled delivery of temozolomide, and in vitro study results showed that 85% transdermal release has been achieved in the environment that mimics skin microenvironment.

14.2.5 Stimuli-responsive chitosan nanogels Stimuli-responsive polymers are novel types of polymers and are recognized as smart or intelligent polymers because of their advantageous ability to give a response to environmental change [39]. These engineered polymers are designed to give response to environmental signals or stimuli by varying their physical and/or chemical features [39]. Drug delivery systems can be designed by using these kinds of stimuli-responsive polymers that respond to environmental alterations such as pH, temperature, light, magnetic field, and so on. Chitosan nanogels are advantageous systems for stimuliresponsive drug delivery systems because of promising characteristics like low toxicity, biocompatibility, and modification ability, and there have been numerous investigations. Stimuli-responsive nanogels can be formulated based on the alteration of either one type of change like pH, temperature, redox, magnetic field, or light or based on dual change like pHredox, temperature-pH, magnetic-light, and so on. In the below section, the recent chitosan nanogel-based stimuliresponsive drug delivery system researches were classified by the stimuli types.

14.2.5.1 Temperature-responsive chitosan nanogels A thermo-responsive chitosan nanogel has been designed to deliver curcumin to decrease insolubility, instability during storage, and absorption limitations of it [40]. In the study, poly-(N-isopropylacrylamide) (pNIPAM) was used due to its

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FIGURE 14.4 Illustration of CMKGM/CS nanogel formation. revised from Wu C, Sun J, Jiang H, Li Y, Pang J. Construction of carboxymethyl konjac glucomannan/chitosan complex nanogels as potential delivery vehicles for curcumin. Food Chem 2021;362:130242. https://doi.org/10.1016/j.foodchem. 2021.130242 and permission will be taken from Wu C, Sun J, Jiang H, Li Y, Pang J. Construction of carboxymethyl konjac glucomannan/chitosan complex nanogels as potential delivery vehicles for curcumin. Food Chem 2021;362:130242. https://doi.org/10.1016/j.foodchem.2021.130242.

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thermo-responsive and low critical solution temperature (LCST) of  C properties that are favorable for biomedical studies like drug delivery and regenerative. Chitosan was used as a natural polymer and backbone both to graft pNIPAM on it and to enhance the biodegradability of the formed nanogel. Grafting of pNIPAM on chitosan was achieved by EDC/NHS coupling reaction, and nanogel form was developed by sonication technique. The production of pNIPAM-grafted chitosan nanogels (CS-g-pN) in the spherical form in 200e300 nm size was accomplished. The uptake and localization studies of curcumin in NIH-3T3 and HeLa cells revealed that nanogels are compatible with these cell lines. Cytotoxicity of the curcumin-loaded nanogels was observed on MDA-231, Caco-2, HepG2, and HT-29 cancer cell lines. In the case of thermo-responsiveness, sensitivity was affected by grafted pNIPAM length and density. Another thermo-responsive chitosan nanogel has been fabricated from the low molecular weight, which is named oligochitosan (short-chain chitosan, w5e15 kDa molecular weight) to bind and release ibuprofen [41]. The formed oligochitosan nanogel has been induced with b-glycerophosphate, and thermo-responsive behaviors of the nanogel have been evaluated with isothermal titration calorimetry and high-sensitivity differential scanning calorimetry (DSC). It has been concluded that the oligochitosan nanogel loaded with ibuprofen below the transition temperature demonstrated binding and slow-release patterns at body temperature.

14.2.5.2 pH-responsive chitosan nanogels Apart from temperature responsiveness, nanogel can be formulated to give response to changes in pH of the environment. For this aim, pH-responsive chitosan-based nanogel loaded with 5-FU has been generated to check its effectiveness against melanoma [42]. The formulation of the nanogel was done by the ionic gelation method in the presence of sodium TPP as a cross-linker agent. It was concluded that chitosan-based nanogels formed in 100e180 nm range size and the release of the 5-FU was induced at a slightly acidic pH. Additionally, the effectiveness of the 5-FU was achieved even at the low dose of 0.2% w/v. The pH-responsive chitosan nanogel drug delivery system has been performed for enhanced tumor drug delivery and augmented chemotherapy [43]. For this purpose, nanogel has been prepared from polypyrrole-grafted chitosan in the presence of glutaraldehyde as a cross-linking agent via the miniemulsion technique. After the preparation of polypyrrolegrafted chitosan nanogel, it has been treated with an alkaline solution to get ultrafast charge reversible nanogel. Alkaline treatment aims to obtain a nanogel that acquires a negative charge at physiological conditions, whereas possesses a positive charge at slightly acidic conditions. Then, pH stimuli-responsive nanogels have been loaded with doxorubicin as an anticancer drug. The evaluation of nanogels to be able to prolong the blood circulation time, enhance penetration, and encourage cellular internalization for augmented chemotherapy of ovarian carcinoma demonstrated that self-adaptive charge alteration of nanogels to be positively charged induces the achievement of the goals of the study. A pHstimulated chitosan nanogel has been formulated for transdermal delivery of capecitabine as a targeting drug for the evaluation of skin cancer [44]. The nanogel has been performed by the ionic gelation method, and pluronic F 127 has been used as copolymer and transcutol has been used to decorate the surface for nonionic penetration (Fig. 14.5). The pH responsiveness of the capecitabine encapsulated chitosan nanogel has been similar to skin cancer environment at slightly acidic conditions (pH 5 and 6), and sustained drug release has been observed at these pH conditions, but at pH 7, drug release has been insufficient. The significant pH-responsive release of capecitabine has been detected at acidic (pH 4) and slightly acidic (pH 5 and 6) in 12e24 h. The same research group recently evaluated chitosan nanogel fabricated in the same manner for the encapsulation of temozolomide for targeting tumor therapy [45]. In the case of the pH stimuli release of temozolomide, it has been stated that pH-triggered and sustained drug release has been observed under mildly acidic (pH 6) conditions. Lastly, rivastigmine release from chitosan and kappa-carrageenan-based nanogel has been observed under different pH conditions and results showed that pH responsiveness has been observed under simulated gastric conditions [14].

14.2.5.3 Dual-responsive chitosan nanogels A dual-responsive, thermo-pH-responsive, chitosan-based nanogel drug carrier system has been formulated by grafting PNIPAM which is a temperature and pH-responsive polymer to chitosan to delivery of curcumin as a therapeutic agent to increase its solubility [46]. In this study, chitosan was thiolated with L-cysteine and the connection of gold nanoparticles (AuNPs) to nanogels was achieved by eSH groups for synchronized drug delivery and photothermal therapy (PTT). The formation of TC-PNIPAM nanogel was performed via the free radical surfactant-free emulsion polymerization method. The results demonstrated that hydrodynamic size of nanogel was nearly 167 nm, and thermo-pH responsive curcumin release from nanogel and AuNP@Ng reached equilibrium at 80% after 72 h at 37 C and acidic pH conditions. The pHresponsive release of curcumin was higher in an acidic environment when compared with physiological pH. The study also indicated that curcumin loading led to an increase in the size of AuNP@Ng from 215.16 to 226 nm, and this nonsignificant

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FIGURE 14.5 Schematic illustration of capecitabine-encapsulated chitosan nanogel formation and transdermal delivery of drug release [44]. permission will be taken from Sahu P, Kashaw SK, Sau S, Kushwah V, Jain S, Agrawal RK, et al. pH triggered and charge attracted nanogel for simultaneous evaluation of penetration and toxicity against skin cancer: in-vitro and exvivo study. Int J Biol Macromol 2019;128:740e51. https://doi.org/ 10.1016/j.ijbiomac.2019.01.147.

change in particle size was attributed to the high loading capacity of the nanogel because of the flexibility and softness of the plasmonic nanogel. Lastly, cytocompatibility evaluation of the plasmonic nanogel on MDA-MB-231 human breast cancer and nontumorigenic MCF 10A cell lines showed that thermo-pH responsive nanogel was cytocompatible for drug delivery. Another dual-responsive chitosan nanogel has been created based on pH and redox responsiveness to increase the solubility of doxorubicin (DOX) and sustain controlled delivery of DOX in cancer cells [47]. For the preparation of pH and redox-responsive chitosan nanogel, first, CS was functionalized with N-phthaloyl (CTS) and then polymerization of poly(hydroxy ethyl methacrylate) (HEMA) was achieved and CTS-g-PHEMA was obtained. Third, maleic acid (Mac) grafting on CTS-g-PHEMA was provided via a reaction between hydroxyl groups of HEMA and maleic anhydride, and CTS-g-PHEMA-MAc was formed. Last, grafting of N, N’ BAC was completed using the double bonds of Mac. As a result, CTS-g-PHEMA-Mac-g-MBA (Fig. 14.6) was prepared as pH & redox responsive nanogel for the triggered release of DOX in cancer cells. Whereas the pH sensitivity of the formulated chitosan-based dual-responsive nanogel comes from carboxylic groups of MAc, redox sensitivity comes from BAC. The dual-responsive CTS-g-PHEMA-Mac-g-MBA nanogel was assessed in terms of DOX-loading capacity and release, biocompatibility, and cytotoxicity. Briefly, the outputs of the study demonstrated that drug-loading capacity and efficiency were found 14% and 76%, respectively. On the other hand, drug release was found more efficient at pH 5 than at pH 7.4. The cytotoxicity analysis results note that DOX-loaded nanogels were more efficient than free DOX. Inorganic compounds such as gold and magnetic particles can be used in nanogel formulations to create sensitive drug carrier materials to environmental changes like optical, electrical, and magnetic [48]. In this regard, researchers have formulated a dual-responsive chitosan nanogel based on magnetic and light responsiveness. The nanogel was made up of chitosan and pNIPAM by radical polymerization, and then modification of the nanogels was done by the incorporation of gold and magnetic nanoparticles to get a magnetic and light-sensitive drug carrier. The gold nanoparticles and magnetic

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FIGURE 14.6 Schematic illustration of CTS-g-PHEMA-Mac-g-MBA nanogel formation for DOX release. permission will be taken from Mahmoodzadeh F, Ghorbani M, Jannat B. Glutathione and pH-responsive chitosan-based nanogel as an efficient nanoplatform for controlled delivery of doxorubicin. J Drug Deliv Sci Technol 2019;54:101315. https://doi.org/10.1016/j.jddst.2019.101315.

particles acted in the nanogel as photothermal transducer and magnetic controller. It was stated by the researchers under green light irradiation the presence of gold nanoparticles in nanogels led to local heat due to their surface plasmon resonance of them and the existence of PNIPAM caused to shrink of nanogels. The gold nanoparticles caused both damage to cancer cells and trigger of drug release. Another dual responsive chitosan-based nanogel has been formulated from multifunctional magnetite mesoporous silica nanoparticles coated with pH-responsive chitosan hydrogel for chemotherapy and hyperthermia [49]. The drug release which was tamoxifen as a model drug has been evaluated at different pH and temperatures for 72 h to observe responsiveness. The results of the study have revealed that release whereas tamoxifen at physiological conditions (pH: 7, T: 37 C) has been 15% and at conditions simulating endosomes/hyperthermia (pH: 5, T: 43 C) has been 70%.

14.2.6 Chitosan nanogels with different delivery approaches Besides the classification of drug delivery systems based on the used chitosan type, chitosan, chitosan derivatives, and functionalized chitosan, it might be classified in terms of delivery approaches. For instance, multiresponsive chitosan nanogel has been formulated for oral delivery of doxorubicin [50], an oral intelligent-responsive core-shell nanogel has been produced from chitosan-oligosaccharide and SA for the treatment of the intestinal infection by Escherichia coli with enrofloxacin [51]. Similarly, an interpenetrating polymer network nanogel has been produced from beta-cyclodextrin and chitosan to improve the solubility and release of rosuvastatin during oral administration [52]. Additionally, a nanogel has been fabricated from chitosan and poly(2-ethyl-oxazoline). After the construction of the nanogel, chlorogenic acid has been encapsulated in it to use for in oral delivery in collagen-induced arthritis [53]. A gastrointestinal delivery-aimed nanogel has been made from succinylated GC and succinyl prednisolone conjugate to treat ulcerative colitis [54], and similar to this study, florfenicol-loaded carboxymethyl-loaded chitosan-gelatin shell nanogel has been formulated for the treatment of E. coli in mice via gastrointestinal delivery [55]. A genipin-cross-linked carboxymethyl chitosan nanogel has been produced and loaded with isoniazid and rifampin for lung-targeted treatment of tuberculosis [56].

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Green tea extract which is rich in antioxidant content has been loaded into chitosan nanogel to increase topical drug delivery opportunities of it [57]. MTX which is a vital therapeutic agent in head and neck carcinomas, and primary brain lymphomas containing chitosan nanogel have been formulated [58]. It has been administered intravenously to understand the effect of drug-loaded chitosan nanogels versus simple drug solutions in vivo. An interpenetrating polymer network nanogel has been fabricated from polyvinyl lcoholl and chitosan for transdermal delivery of ondansetron, which is a steroid [59]. Like this study, a chitosan-SA-based nanogel has been constructed and used for the transdermal delivery of pirfenidone used for the treatment of idiopathic pulmonary fibrosis to eliminate drawbacks caused by oral delivery [60]. A recent study has focused on the production of chitosan-based nanogel for OA drug delivery [61]. For this purpose, chitosan and sodium hyaluronate-based nanogel has been formulated by the ionic gelation method and cell studies of the produced nanogels have been conducted on human cells obtained from osteoarthritic patients and zebrafish embryos. It has been stated that the nontoxicity of up to 100 mg/mL and physicochemical properties of the chitosan-based nanogels were suitable for OA drug delivery. An ocular treatment aimed at chitosan nanogel has been fabricated with the ionic gelation method and loaded with previously formed nanovesicles containing acetazolamide, which is used orally for intraocular pressure [62]. It has been aimed that reduced the side effects of acetazolamide caused by oral usage by introducing it to mucoadhesive chitosan nanogel formulation to be applied topical delivery. A chitosan-based nanogel drug delivery system has been formulated for the treatment of hyperlipidemia, which is known excess amount of lipids or lipoproteins in the blood [63]. For this purpose, initially, pravastatin, which is used as a cholesterol-lowering drug containing chitosan nanoparticles, has been produced by ionic gelation technique and then drugloaded nanoparticles have been used for the synthesis of nanogel in the presence of poloxamer.

14.3 Chitosan nanogels in regenerative medicine Nanogel formulations of chitosan can be used in biomedical applications for regenerative purposes. These applications might aim to repair damaged tissues, for instance, repair skin after burns or wounds, or protection of tissues against possible damages, for instance, antimicrobial infection, and so on.

14.3.1 Chitosan nanogels applications in wound healing The nanogels can be used as directly as a carrier for drugs or as support of the polymer matrix, as in recent research [64]. The nanogel that has been produced from maleoyl-chitosan and poly(aspartic acid) (MAC5/PAS) to incorporate into the hydrogel of thiolated hyaluronic acid (HASH) hydrogel as both carrier of amoxicillin (Amox) and building blocks of hydrogel. The prepared nanocomposite hydrogel was aimed to use as a therapeutic scaffold that mimics extracellular matrix and carrier antimicrobial agent against gram-negative wound infections. The outcomes of the research presented that filling the HASH hydrogels with nanogels led to a decrease in swelling capacity of hydrogels and a controlled degradation rate in contrast to hydrogels without filling of nanogels. Also, in vitro and in vivo assay results showed that rapid release of Amox was observed at pH 5.4, and hydrogels were biocompatible for their usage of them as therapeutic scaffolds in dressing applications. In a recent study, chitosan nanogel has been formulated to indicate the wound-healing effect of probiotics [65]. The chitosan nanogel prepared with the ionic gelation method has loaded with Lactobacillus reuteri, Lactobacillus fermentum, and Bacillus subtilis sp. Natto probiotic strains supernatants. The wound-healing effects of chitosan nanogels with probiotic lysate have been tested on SpragueeDawley rats as an animal model, and chitosan nanogel without probiotic lysate has been used as a control. It has been indicated that although chitosan has antimicrobial activity, the wound-healing activity has been observed only for probiotic-loaded nanogels. The wound-healing activity of probiotics has presented differently, for instance, while Lactobacilli reuteri and L. fermentum have shown an effect on the rate of wound closure, B. subtilis sp. natto has been active on wound-healing quality and slow wound closure rate it has been attributed to the presence of nattokinase enzyme that possesses anticoagulant activity. It has been concluded that a chitosan nanogel formulation loaded with probiotic lysate might be a promising candidate for wound-healing treatment with antibiotic, swelling, and rheological properties. The macrophages have significant role in wound-healing treatments, while classically activated microphages (M1) are important in terms of proinflammatory and start of immune cell recruitment, alternatively activated macrophages (M2) are critical for resolution of inflammation phase and reformation of damaged tissues, and their balance is significant to be effective on wound healing [66]. Therefore, the research focuses on the improvement of immune response in wound healing by forming a covalently bounded chitosandaloe vera nanogel. While the aim of the practice of chitosan in wound healing has been attributed to its immune-modulatory properties and

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aloe vera has been preferred, particularly its antiinflammatory and antimicrobial properties. The antimicrobial and antifungal effects of nanogels were tested on various bacterial and fungal strains, and the results demonstrated that it had an antimicrobial effect on Staphylococcus aureus and Salmonella typhi and a fungicidal effect on Kluyveromyces marxianus and Candida albicans at different chitosan to aloe vera ratios. On the other hand, wound-healing studies have been performed on Wistar rats as model animals. The results of the research presented that the action of aloe vera in wound healing has been observed as generally antiinflammatory, and the blend of the aloe vera and chitosan (1 and 3:1 v/w) led to the modulation of M1-M2 responses as that hypothesized by the researchers. In a similar study of chitosan, nanogel formulation has been performed to achieve treatment of dermatophytosis which is a disease resulting from infection of pathogenic fungi Microsporum canis and mostly affects keratinized structures like hair, nails, and skin [67]. For this purpose, modification of chitosan has been achieved with cinnamic acid to improve its chemical properties of it, and then nanogel form has been produced. The produced nanogels also have been used to encapsulate essential oils of Syzygium aromaticum and Cinnamomum ssp. To detect antifungal activities of them against M. canis. It has been achieved that cinnamic acid-grafted chitosan nanogel with 176.0 and 263.0 nm average size, respectively, encapsulated with essential oils of Cinnamomum and S. aromaticum. On the other, the average size of cinnamic acid-grafted chitosan nanogel without encapsulation of essential oils has been 335.4 nm. And also encapsulated formulation of cinnamic acid-grafted chitosan nanogel has been more effective to inhibit M. canis than without essential oil formulation, and their effectiveness has been 74% and 89% for S. aromaticum and Cinnamomum, respectively. It has been demonstrated that modification of chitosan provided that affinity of essential oils and encapsulated forms of cinnamic acid-grafted chitosan nanogel might be an alternative application for the treatment of dermatophytosis. Protection of damaged tissues like skin from infections is important. Skin tissues are damaged by burning to tend sepsis that ended with mortality, and the protection requires intensive care. The research has aimed to prepare a magnetic nanogel to facilitate from antimicrobial properties of magnetic nanoparticles, chitosan, and gelation and to improve the release of chlorhexidine gluconate used for burn care [68]. The nanogel has been produced from a mixture of chitosan and gelation with an average size of 39.75 nm, and the previously produced magnetic nanoparticles of cobalt ferrites introduced the nanogel, and lastly inoculation of magnetic nanogel with chlorhexidine gluconate has been done. Characterization and electrochemical studies of the magnetic nanogel indicate that nanogel gives response to pH changes, and release of the drug can be achieved by triggering the pH from 6 to 7. And also, it has been concluded that besides the release-enhancing properties, chitosan-gelatin magnetic nanogel can be used for the detection of sepsis via magnetic particles. Another current research has been focused on wound healing of skin burns by designing a chitosan-poloxamer-apigenin nanogel as a skin dressing [69]. In this formulation, chitosan and poloxamer were chosen to produce polymer structure and apigenin, which is a natural flavonoid preferred because of its antimicrobial, antiinflammatory-like benefits. The antibacterial activity of chitosan-poloxamer-apigenin nanogel has been tested against S. aureus, E. coli, and Pseudomonas aeruginosa. It has been concluded that the designed wound dressings had low toxicity, and it had bacteriostatic activity especially higher for gram-negative bacteria. And also, the bacteriostatic activity of this wound dressing contributed to the nature of the chitosan. Lastly, chitosan-carboxymethyl nanogel has been designed to benefit from the antimicrobial, antiinflammatory properties of the essential oil of Nigella sativa and second-generation synthetic statin, atorvastatin in a transdermal application for wound healing [70]. First, chitosan-carboxymethyl nanogel has been produced with the solvent-evaporation emulsification method, then produced nanogel firstly loaded with essential oil of N. sativa to obtain oil nanogel (ONG) and then ONG has been loaded with atorvastatin to obtain atorvastatin-oil nanogel (ATONG). ONG and ATONG have been produced with 172 and 193 nm particle size, respectively. After the evaluation of nanogels toward in vitro drug release, cytotoxicity, wound closure by fibroblasts, in vitro rabbit ear skin permeation of ONG and ATONG, and antimicrobial activity, it has been noted that ATONG was successful in terms of drug loading, drug release, and stability and also release of atorvastatin from nanogel is intracellularly in fibroblasts in a safe manner. Additionally, the results of the scratch test and the antimicrobial test showed that ATONG might be an alternative treatment for wound-healing therapy.

14.3.2 Chitosan nanogels applications in dentistry The formulation of chitosan-based nanogel loaded with Mentha piperita essential oil (MPEO) has been done to inhibit the biofilm of Streptococcus mutants on dental surfaces [71]. The formulation of chitosan nanogel was performed by solegel technique in the presence of TPP as a cross-linker agent. The release results of the research have demonstrated that MPEOloaded chitosan-TPP nanogel was appropriate for the controlled release of MPEO for the protection of dental surfaces from the biofilm of S. mutants [71]. Biofilm formation on the tooth surface causes dental caries and periodontal disease and is generally caused by S. aureus, Klebsiella pneumoniae, and P. aeruginosa [72]. For the treatment of biofilm, a chitosanbased nanogel has been formulated with the ionic gelation method in the presence of calcium phosphate, and an

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antimicrobial agent which is chlorogenic acid has been loaded into in chitosan nanogel to be used for degradation of biofilm [72]. The antimicrobial activities of chlorogenic acid-loaded chitosan nanogel have been tested against S. aureus, K. pneumoniae, and P. aeruginosa at different chlorogenic acid concentrations. The zone of inhibition results of the chlorogenic acid-loaded chitosan nanogel showed that S. aureus has caused to 18.7 mm zone of inhibition. And the biofilm degradation achievement of the chlorogenic acid-loaded chitosan nanogel was 68%. It has been stated that chlorogenic acid-loaded chitosan nanogel with positive zeta potential value was more effective against gram-positive than gramnegative bacteria by disrupting the bacterial cell membrane. And also, it has been concluded that the negative charge of the phosphate ions and the positive charge of the chitosan might be responsible for the neutralization of bacterial growth. A chitosan-based nanogel system has been developed for the treatment of periodontitis, which is a dental inflammatory disease that caused damage to soft tissue and bone that surround and supports the tooth [73]. The chitosan nanogel has been formulated to load both antiinflammatory and antimicrobial agents for the treatment of periodontitis. For this purpose, triclosan has been selected as an antimicrobial agent and flurbiprofen has been used as an antiinflammatory agent to understand their dual activities of them on periodontitis. The nanogel formulation has been produced from different stages. First, nanoparticles of triclosan have been produced via the solvent displacement method, and then chitosan hydrogel with flurbiprofen has been produced, and at the end step, nanogel formulation has been achieved by adding triclosan nanoparticles in the flurbiprofen-chitosan hydrogel. The antimicrobial evaluation of dual effective chitosan nanogel on E. coli and S. aureus has shown an inhibition effect by producing zone of inhibition as 10 and 13 mm, respectively. On the other hand, the evaluation of the antiinflammatory effect of produced nanogels has been done in vivo on experimental periodontitis (EP) rats by considering GI and PI values which are responsible for gingival inflammation and accumulated plaque, respectively. The results of the study have presented that because of the resolution of inflammation and reduction of plaque, nanogels possessed an antiinflammatory effect. The structure of dentinal tubes in the tooth looks like a honeycomb structure and includes dentinal fluid, and in the case of opening of the dentinal tubules, dentin hypersensitivity can occur and it can cause painful or uncomfortable sensations [74]. Therefore, to reduce the effect of dentin hypersensitivity, dentinal tubules must be occluded. Recent research has aimed to produce material from chitosan-based nanogel for occlusion of the dentinal tubule. The formation of the nanogel was accomplished through the utilization of lysozyme and carboxymethyl chitosan. Subsequently, the nanogel was employed for encapsulating amorphous calcium phosphate, which serves as an occlusive material for dentinal tubules. The outputs of the study have confirmed that the calcium and phosphate ions that encapsulated by carboxymethyl chitosan and lysozyme nanogel had ability to cover dentinal tubules and might be used for dentinal remineralization. The following study of the research group has aimed the utilization of the same amorphous calcium phosphate-encapsulated carboxymethyl chitosan and lysozyme nanogel for controlled delivery of calcium and phosphate ions and formation of aprismatic enamel-like layer on the demineralized enamel surface [75]. They have concluded that the formed nanogel has been suitable for the prevention and treatment of early enamel caries and required that more nanogel application.

14.4 Characterization methods of chitosan nanogels in drug delivery The chitosan nanogels must be characterized efficiently to understand both their physicochemical properties and suitability for drug delivery and regenerative medicine applications. Because characteristics of the nanogels determine the efficiency of nanogels in usage. The physicochemical characterization of nanogels can be made with different measurement techniques to analyze different properties. For instance, dynamic light scattering (DLS) is an important tool to characterize three important features of nanogels. DLS measurement allows for the determination of key parameters in assessing the particle homogeneity of nanogels, such as the polydispersity index (PDI), particle diameter, and zeta potential values. The PDI serves as an indicator of particle uniformity, while the particle diameter and zeta potential values provide insights into the magnitude of particle charges responsible for dispersion, aggregation, or flocculation processes [26,76] can be made by DLS measurement. The relative surface charge should be considered during the process of nanogel because electrostatic repulsive forces are effective on the physiochemical stability of the polymers and so aggregation mechanism. The characterization of nanogels by X-ray diffraction (XRD) is made to get information about the amorphous or crystalline properties of the polymers used in nanogel formation, formed nanogel, and drugs that are loaded in or onto nanogels. To verify the successful formulation of the nanogel, Fourier transform infrared spectroscopy (FTIR) analysis should be conducted. This analysis allows for the examination of the physical state and molecular interactions among the components used in the nanogel formulation, including the polymer, monomer, cross-linking agent, and the drug intended for loading. By assessing these interactions, it can be determined whether the desired formulation has been achieved or not. The morphological properties of the nanogels are evaluated by scanning electron microscope (SEM) by taking surface and cross-section images of nanogels. Transmission electron microscope (TEM) analysis is made to analyze the size, shape,

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and morphology of nanogels. Thermal stability and changes in physical properties of nanogels with respect to temperature are evaluated through thermal gravimetric analysis (TGA) and differential scanning calorimetry (DSC). TGA is utilized to measure the thermal stability of nanogels, while DSC is employed to monitor alterations in their physical characteristics over time when exposed to elevated temperatures. For the TGA analysis change in weight is considered, and for the DSC analysis, heat generation is considered. Last, tensile properties like tensile at the break, young’s modulus, and elongation of the nanogels can be done to get information about the structure of the nanogels. After the physicochemical characterization studies of the nanogels, their suitability for drug delivery and regenerative medicine applications on biological systems should be evaluated using various assays. For instance, drug-loading and release patterns of nanogels can be assessed by UV-vis spectrometer. Cytotoxicity, biodegradability, inflammatory response, hemocompatibility, cellular uptake, intracellular fate, and in vivo biodistribution are some important key parameters [11] that should be considered in the usage of nanogels in biological systems like drug delivery and regenerative medicine.

14.5 Conclusion Chitosan has a unique chemical structure because of its hydroxyl and amino groups that simplify chemical reactions and functionalization of it. Hence, chitosan has the opportunity to be modified to enhance its properties like solubility. The polyelectrolyte nature, mucoadhesive property, nontoxicity, in situ gelling ability, and biocompatibility features increase the utilization capacity of chitosan in different areas and different formulations. For example, its polyelectrolyte nature led to easy interaction of it with various biomolecules like a drug, DNA, proteins, and so on. Therefore, it is a suitable compound for biomedicine applications. Recently there has been various research that focused on the formation of chitosan nanogel for drug delivery and regenerative medicine. In this chapter, recent chitosan-based nanogels systems for drug delivery and regenerative medicine have been mentioned. It tried to classify drug delivery studies based on the chitosan type used in nanogel formulation, stimuli-responsive formulation, and types of drug delivery approaches. On the other hand, chitosan nanogels for regenerative medicine have been summarized based on wound healing and dentistry applications.

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Int J Biol Macromol 2018;118:1319e25. https://doi.org/10.1016/j.ijbiomac.2018.04.147. [61] Manivong S, Garcia AA, Patten SA, Fernandes JC, Benderdour M, Banquy X, et al. Chitosan-based nanogels: synthesis and toxicity profile for drug delivery to articular joints. Nanomaterials 2022;12. https://doi.org/10.3390/nano12081337. [62] Abdel-Rashid RS, Helal DA, Omar MM, El Sisi AM. Nanogel loaded with surfactant based nanovesicles for enhanced ocular delivery of acetazolamide. Int J Nanomed 2019;14:2973e83. https://doi.org/10.2147/IJN.S201891. [63] Saraogi GK, Tholiya S, Mishra Y, Mishra V, Albutti A, Nayak P, et al. Formulation development and evaluation of pravastatin-loaded nanogel for hyperlipidemia management. Gels 2022;8:1e15. https://doi.org/10.3390/gels8020081. [64] Rusu AG, Chiriac AP, Nita LE, Ghilan A, Rusu D, Simionescu N, et al. Nanostructured hyaluronic acid-based hydrogels encapsulating synthetic/ natural hybrid nanogels as promising wound dressings. Biochem Eng J 2022;179:108341. https://doi.org/10.1016/j.bej.2022.108341. [65] Ashoori Y, Mohkam M, Heidari R, Abootalebi SN, Mousavi SM, Hashemi SA, et al. Development and in vivo characterization of probiotic lysatetreated chitosan nanogel as a novel biocompatible formulation for wound healing. Biomed Res Int 2020 2020. https://doi.org/10.1155/2020/ 8868618. [66] Ashouri F, Beyranvand F, Beigi Boroujeni N, Tavafi M, Sheikhian A, Varzi AM, et al. Macrophage polarization in wound healing: role of aloe vera/ chitosan nanohydrogel. Drug Deliv Transl Res 2019;9:1027e42. https://doi.org/10.1007/s13346-019-00643-0.

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[67] de Carvalho SYB, Almeida RR, Pinto NAR, de Mayrinck C, Vieira SS, Haddad JF, et al. Encapsulation of essential oils using cinnamic acid grafted chitosan nanogel: preparation, characterization and antifungal activity. Int J Biol Macromol 2021;166:902e12. https://doi.org/10.1016/ j.ijbiomac.2020.10.247. [68] Khan MA, Mujahid M, Gul IH, Hussain A. Electrochemical study of magnetic nanogel designed for controlled release of chlorhexidine gluconate. Electrochim Acta 2019;295:113e23. https://doi.org/10.1016/j.electacta.2018.10.098. [69] Zou J, Lv L, Qi X. Design of chitosan-poloxamer nanogel and evaluation of its therapeutic effect on skin burn. Mater Express 2021;11:1800e7. https://doi.org/10.1166/mex.2021.2099. [70] Bagheri F, Darakhshan S, Mazloomi S, Shiri Varnamkhasti B, Tahvilian R. Dual loading of Nigella sativa oil-atorvastatin in chitosane carboxymethyl cellulose nanogel as a transdermal delivery system. Drug Dev Ind Pharm 2021;47:569e78. https://doi.org/10.1080/ 03639045.2021.1892742. [71] Ashrafi B, Rashidipour M, Marzban A, Soroush S, Azadpour M, Delfani S, et al. Mentha piperita essential oils loaded in a chitosan nanogel with inhibitory effect on biofilm formation against S. mutans on the dental surface. Carbohydr Polym 2019;212:142e9. https://doi.org/10.1016/ j.carbpol.2019.02.018. [72] Palaniraj S, Murugesan R, Narayan S. Chlorogenic acid- loaded calcium phosphate chitosan nanogel as biofilm degradative materials. Int J Biochem Cell Biol 2019;114:105566. https://doi.org/10.1016/j.biocel.2019.105566. [73] Aminu N, Chan SY, Yam MF, Toh SM. A dual-action chitosan-based nanogel system of triclosan and flurbiprofen for localised treatment of periodontitis. Int J Pharm 2019;570. https://doi.org/10.1016/j.ijpharm.2019.118659. [74] Song J, Wang H, Yang Y, Xiao Z, Lin H, Jin L, et al. Nanogels of carboxymethyl chitosan and lysozyme encapsulated amorphous calcium phosphate to occlude dentinal tubules. J Mater Sci Mater Med 2018;29. https://doi.org/10.1007/s10856-018-6094-9. [75] Song J, Li T, Gao J, Li C, Jiang S, Zhang X. Building an aprismatic enamel-like layer on a demineralized enamel surface by using carboxymethyl chitosan and lysozyme-encapsulated amorphous calcium phosphate nanogels. J Dent 2021;107. https://doi.org/10.1016/j.jdent.2021.103599. [76] Dickinson E. Hydrocolloids as emulsifiers and emulsion stabilizers. Food Hydrocolloids 2009;23:1473e82. https://doi.org/10.1016/ j.foodhyd.2008.08.005.

Chapter 15

Heparin-based nanocomposite hydrogels Amrita Thakur1, Vinay Sagar Verma2, Jyoti Ahirwar3, Sandeep Kumar Sonkar4 and Hemant Ramachandra Badwaik5 1

Department of Pharmaceutics, School of Pharmacy, Vishwakarma University, Pune, Maharashtra, India; 2Shri Shankaracharya Technical Campus,

Faculty of Pharmaceutical Sciences, Bhilai, Chhattisgarh, India; 3Daksh Institute of Pharmaceutical Sciences, Chhatarpur, Madhya Pradesh, India; 4

Rungta College of Pharmaceutical Sciences and Research, Raipur, Chhattisgarh, India; 5Shri Shankaracharya Institute of Pharmaceutical Sciences

and Research, Bhilai, Chhattisgarh, India

15.1 Introduction In recent decades, merging nanotechnology with other scientific disciplines has garnered significant interest. Various attempts have been made to combine nanoscale procedures with traditional techniques to produce better materials. Nanocomposite (NC) hydrogels (HGs) are an illustration of such a partnership between nanotechnology and biomaterial research (HGs). HGs are made up of interconnected, synthetic, or natural polymer chains that are cross-linked to create a hydrophilic material with a macromolecular structure resembling a gel. They can expand to so many times their weight (W/W) and contain up to 99% water or biological fluids [1]. This highly hydrated, three-dimensional (3D), porous network may mimic the environment of real tissue and is frequently widely employed to store, release, or gather components [2]. Depending on the production process, HGs can either be chemical (thermosetting) or physical (thermoplastic) in nature. Chemical gels are produced by covalent cross-linking using procedures like polymerization with a cross-linker or other methods of cross-linking already-existing polymers (heat, ultrasonic, ultraviolet [UV], irradiation, etc.). A physical gel is made up of an amorphous network of hydrophilic polymers that are bound together by noncovalent interactions like hydrogen bonds and van der Waals forces. Chemical gels expand but do not dissolve in water, whereas physical gels finally dissolve in water and may be melted with heat [1]. When choosing materials to create an HG, one should consider swelling, mechanical characteristics, diffusion rates, and chemical functionality. These features are influenced by the crosslinking density, intercross-link distance, macromolecular structure, and residual compounds in the gel (monomers, initiators, etc.). When used as scaffolds during tissue engineering or in any other application where high mechanical strength, robust compression tolerance, and good elasticity are required (such as cartilage tissues), HGs’ inadequate mechanical strength is a disadvantage [3]. They are difficult to transfer and load onto different body components due to their weak mechanical properties. The potential biological uses of HGs have been investigated, and contemporary research is focusing on ways to enhance their mechanical and chemical characteristics. The high surface area-to-volume ratios of nanoparticles (NPs) and nanolayers make them perfect for utilization in polymer networks [4]. By physically and chemically crosslinking polymeric chains with various nanoscaled designs, NC HGs are networks with unique, desired features. New properties and behaviors manifest once fillers are nanoscale distributed throughout the composite. The addition of such a dispersion can also improve several properties of the empty matrices [5]. Due to their openness to environmental stimuli including light, heat, and electric/magnetic fields, HGs gain from this inclusion. The network’s cells can expand, release, or absorb water, and discharge any accumulated contents depending on the stimulation. Smaller particle size HGs react more quickly [3]. This property enables the use of HGs in biosensors and biodevices. The introduction of NPs makes the polymer network more complicated, enhancing mechanical toughness. The polymer chains are connected by NPs. The ideal mechanical strength of biomaterials is crucial for their use in contact lenses and wound dressings for regenerative medicine. HGs are a potential material for wound dressing because of their strong adherence to surfaces, particularly skin or soft tissues, and mobility. NCs maintain their targets because the rough surfaces they create by incorporating NPs into the network interact with the flaws of the surface. The capacity to create microfabricated structures is necessary for manufacturing at the nano- and microscale. 2D surfaces are most frequently Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine. https://doi.org/10.1016/B978-0-323-95351-1.00011-9 Copyright © 2024 Elsevier Inc. All rights reserved.

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FIGURE 15.1 Overview of nanocomposite hydrogels.

employed for cell culture. These traditional development conditions greatly differ from the typical 3D tissue environment and may lead to aberrant cell behavior. Therefore, it is crucial to develop 3D cell culture platforms that can faithfully mimic the conditions of real tissue. Microfabrication techniques must be used to create biomaterials that are secure for cells. Making visible components that use the peculiar physical and mechanical features of the little things they contain is tricky when creating NCs [4]. Fig. 15.1 provided a general overview of NC HGs.

15.2 Advantages of nanocomposite hydrogel over hydrogel The field of bioengineering, which includes specialties like tissue engineering and medication delivery, has led the way in developing novel HG applications during the previous 20 years. However, new applications demand enhanced HG characteristics. It may be possible to get around some of the present restrictions by combining HGs (both natural and synthetic) with precisely specified nanomaterials that can now be synthesized using both organic and inorganic components. Examples of such materials include NC HGs, which can be engineered to mimic natural tissues’ conditions more closely or provide individualized drug release profiles. These biomaterials can potentially improve our understanding of critical biological processes and the optimization of many other characteristics necessary for biomedical use. There are many benefits of using NC HGs instead of the regular HG. In biomedicine, NC HGs represent cutting edge of new material developments. The optical, sensor, actuator, tissue engineering, and many other industries can benefit from using these mechanically robust and better HGs. There are many advantages of these HGs over the more common types. They’re incredibly stable, stiff, and have excellent conductivity and strength. The NC HGs outperform regular polymer HGs in every measurable way. On the basis of the integration of the NCs into the polymer system, NC HGs can perform and enhance mechanical, electrical, or biological capabilities [6,7]. The NC HG’s composition improves its inherent qualities. It has been shown, for instance, that surface functionalization may concurrently change the mechanical, chemical, and biological properties of NPs by enhancing their interactions. There are a number of advantages that NC HG has over conventional HG.

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15.2.1 Mechanical The control of cellular function and the determination of a material’s biological applicability are greatly influenced by its mechanical characteristics [8]. For instance, HGs must adjust to the mechanical conditions of the body at implantation locations [9]. To enhance mechanical performance, characteristics of HGs, the cross-linking density is often increased [10]. An HG becomes stiffer, less porous, and less hygroscopic as its cross-linking density rises. Alternately, HGs’ mechanical characteristics can be enhanced by NPs, which will increase their stability and biocompatibility.

15.2.2 Electrical The discovery of electrically conductive materials could have significant implications in several fields of biomedicine. Muscles, neurons, and heart tissue, to name a few, all need to be able to conduct electricity. Tissue engineers could benefit from biomaterials that more closely resemble their native tissue microenvironments if they learned to imitate these processes. In applications such as medication delivery, biosensors, and actuators, conductive biomaterials may also play a significant part in controlling cellular activity [11]. The creation of electrically conductive HGs may be made possible by including a variety of conductive NPs through into polymeric network. The material’s electrical conductivity wouldn’t be harmed in this way, not even by HG networks. New forms of electrically conducting NC HG were synthesized out of a diversity of polymeric (both natural and synthetic) and nanomaterial (both organic and inorganic) sources, and they have a number of interesting biological applications.

15.2.3 Biological The biological qualities of a material are crucial in determining its viability and potential for biomedical usage. HGs generally fail in vivo due to a lack of cell/matrix connections, despite having superior mechanical and electrical properties [12]. Essential biological features of biomaterials that need to be characterized in depth include toxicity and biomaterial/ cell interactions. Since it reflects a biomaterial’s response to the biological environment (positive or negative), evaluating biocompatibility is crucial [13e15]. Bioactivity and HG breakdown are two more crucial factors to consider. HGs’ biological characteristics can be substantially changed by the addition of NPs. The biological compatibility, stability, and properties of NC HGs are briefly discussed in Ref. [2]. Numerous physiologically beneficial characteristics, including cell adhesion, proliferation and migration, breakdown, and immunological control, can be enhanced by NC HGs. These techniques frequently concentrate on nanomaterials with a variety of functionalities.

15.2.4 Magnetic Adapting to new conditions is a crucial part of a cell’s normal behavior, which is often regulated through two-way, feedback-based communication. HGs must be responsive to stimuli for dynamic processes beyond structural support if they are to serve these purposes [15,16]. Receptivity to stimuli can be induced by chemical and physical changes that destroy or produce secondary forces (also including electrostatic interactions, hydrogen bonds, etc.), or even by straightforward acid-base reactions [17,18]. The HG structure’s sensitivity to changes in secondary stresses is significantly improved by the use of NPs. Utilizing magnetic NPs enables a certain type of remotely controlled reaction to stimuli. These magnetically sensitive NC HGs have been used to create better drug delivery systems, magnetic sensors, and diagnostic imaging tools [19,20].

15.3 Types of nanocomposite hydrogels In order to create NC HGs, various bases are combined with various kinds of nanomaterials. Carbon-based materials, such as carbon nanotubes or CNTs, graphene, and nanodiamonds, as well as polymeric NPs (such as dendrimers, liposomes, and nanogels, etc.), inorganic/ceramic NPs (such as hydroxyapatite, nano clays, glass ceramics, wollastonite, and calcium phosphate), and metal/metal oxide NPs are examples of nanomaterials (such as gold, silver, iron oxide, zinc oxide, and copper oxide) (Fig. 15.2).

15.4 Heparin Heparin (HP) is a biocompatible and water-soluble polysaccharide. It is produced in the granules of mast cells and basophils, a vital part of animal immunological response. HP is primarily a linear sulfated glycosaminoglycan in terms of its

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FIGURE 15.2 Types of nanocomposite hydrogels.

structure. As shown in Fig. 15.3, repeating units of D-glucosamine (2-amino-2-deoxyglucopyranose) and uronic acid make up the polysaccharide (pyranosyluronic acid). Like other polysaccharides, it is a polydisperse mixture of chains with various molecular weights [21,22]. Molecules are typically 15 kDa in size. A trisulfated disaccharide is the main structural element of HP. It also features a few more disaccharide structures. A disaccharide that makes up the majority of HP’s main repeating unit is composed of N- and 6-O-sulfated-D-glucosamine, -L-iduronic acid, and to a smaller amount, -D-glucuronic acid, and -N-acetyl glucosamine (10%). The structural dissimilarity at the disaccharide level that exists in HP leads to sequence microheterogeneity.

15.4.1 Heparin’s role in biological systems HP is a famous and very efficient anticoagulant and antithrombotic medication with use in other therapeutic disciplines [23e25]. By combining HP with NPs, the biological effects of HP may be enhanced. However, due to the synergistic properties that are shown by this kind of composite, the combination of nanomaterials with biomolecules is of great interest. Additionally, HP improves the biocompatibility, and performance of NPs has a wide range biological application, such as imaging, biosensors, drug delivery, tissue engineering, etc. [26]. The polyelectrolyte HP has the highest negative charge density and tallest packed carboxyl and sulfo groups of any well-known natural macromolecule. As a result of its broad negatively charged structure, HP is better able to interact electrostatically with various proteins in the plasma and outer cellular matrix, such as chemokines, proteases, and enzymes like antithrombin III that control cell adhesion, motility, augmentations, and segregation [27]. These interactions typically lead to protein stability or help cell receptor kinship. It also has a high affinity for a number of growth enhancers via electrostatic engagement. Materials like biosensors and scaffolds are the frameworks for limited discharge made using fibroblast and vascular endothelial growth factor to stabilize HP [28].

FIGURE 15.3 Heparin’s structural unit.

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15.5 Preparation of heparin-based nanocomposites HP may associate with many proteins that include HP-fastening peptide domains with a wide variety of proteins that have HP-fastening peptide domains. Working with biomolecules created for diverse biological uses is made simpler for HP by these relationships [29]. It is also likely to electrostatically associate with some other positively charged compounds due to the relatively significant negative charge present, which can lead to the creation of NPs. By adding HP to functionalized nanomaterials, polymers, and other materials, many kinds of NC materials have been developed [30]. HP has been added to functionalized nanomaterials, polymers, and other materials to form several types of NC materials. Nearly often, NPs are added to NCs to either add additional qualities or act as a strong support. Numerous techniques might be used to make polysaccharide NCs (Fig. 15.4). Systematically reviewing the numerous methods for preparing HP NCs for tissue engineering was done by Barik, Sharma, and Rath [31].

15.5.1 Electrospinning A key method for creating fibers with diverse variety of dimensions, from nanometers to micrometers, is electrospinning. These fibers have tiny pore diameters and a high ratio of surface to volume in their raw form. A typical electrospinning system consists of a high-voltage source that generates a powerful electric field, a fiber collector, and a syringe pump that can contain both naturally occurring and synthetic polymeric solutions. The structure of the as-obtained fibers is influenced by the surface tension, molecular weight, conductivity, and viscosity of the polymer solution, as well as by the applied voltage. A few of these variables are temperatures, flow velocity, humidity, and the separation between the collectors and the siring pump [32]. Nanoscale fibers are significantly more suitable with the microstructures of the extracellular matrix (ECM) than micrometer-sized fibers. Electrospun nanofibrous designs outperform microfibrous architectures in several ways, including being tightly tethered to ECMs, having a greater surface area than microfibers, resulting in noticeable protein adsorption, and improving the movement of nutrients between rapidly growing cells and their surrounding live tissue. Despite the fact that the electrospinning technique is well established, it is still uncommon to use electrospun fibers for tissue engineering in the present [33e36]. Currently, this method is widely utilized in the production of HP-based NCs. Kwon and Matsuda, for instance, functionalized elastomeric nanofibers using the electrospinning method. According to this procedure, an equimolar ratio of poly(L-lactide-co-caprolactone) (PLCL) was initially copolymerized. Afterward, 1,1,1,3,3,3-hexafluoro-2-propanol (HFIP) was used to solubilize the copolymerized PLCL, the HP-TBA salt, or type I collagen in order to make them electrospinning-ready. As a consequence, nanofibers are created, with a mean diameter ranging from 120 to 520 nm [37]. By using the electrospinning method, Volpato et al. created a novel matrix-imitating nanoassembly [38]. The already-made chitosan (CS) electrospun fibers were then combined with polar oppositely charged polyelectrolyte complex nanoparticles (PCNs), which eventually serve as fibroblast growth factors. The use of artificial materials in tissue engineering has drawn attention to two significant challenges, however: the cytocompatibility of materials and the biomechanics of some acellular matrix. In order to address these problems, poly(e-caprolactone) (PCL) nanofibers are electrospun outside the decellularized rat aorta artery. They are then utilized to create a hybrid tissueengineered vascular (HTEV). After that, exogenous HP was injected into the acellular artery’s interstitium to prevent

FIGURE 15.4 Representation of cells and its external environment. Adopted with permission from the Barik S, Sharma RK, Rath C. Heparin-based nanocomposites for tissue engineering. Polysaccharide-based nanocomposites for gene delivery and tissue engineering. Woodhead Publishing; 2021. p. 81e101.

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thrombosis and lessen platelet aggregation [39]. Fibrous membranes created utilizing this electrospinning technology have been discovered to hold great promise for the simultaneous regeneration of hard and soft tissues. Using a two-step electrospinning method, Liu et al. created a unique bilayer fibrous membrane that allows for targeted tissue regeneration (GTR) [40].

15.5.2 Layer-by-layer assembly This is an additional simple and flexible way for creating micronanostructured bone scaffolds and bioactive materials/NCs. This process involves depositing different layers of components with various charges onto a substrate. Electrostatic interactions between various charged components, which stabilize the various components as a result, influence how different layers are assembled [41e45]. The main advantage of this approach would be that materials can be produced just at nanoscale level that used a range of assembly related elements, allowing the properties of the final product to be modified as needed. A number of solid and colloidal material can also be employed as a substrate to produce the layers, regardless of size or structure [46e49]. Due to the minimal processing requirements, the preparation method is also well fitted to its physiological environment. This technology offers additional advantages in addition to those discussed, including such cost efficiency and simple, easy assemblage for production, which makes it simpler to prepare a range of biomaterials with particular features on a particular substrate. Even though biomaterials are stacked with other biomaterials, such as CS (as cationic surface), HP (an anionic surface), etc., to improve biocompatibility [50,51]. For instance, Zhang et al. created polyelectrolyte multilayered vascular patches using layer-by-layer self-assembly (PEM). A polyurethane-coated decellularized scaffolding (PU/DCS) is covered with alternate layers of HP and CS, two substances having essentially opposing charges [52]. Hyaluronic acid and HP were used in the layer-by-layer assembly approach to effectively coat stainless steel (SUS316L) [53].

15.5.3 Covalent functionalization A carboxylic group and an amino group constitute typical HP moieties. Covalent bonding with the carboxylic group makes it simple to immobilize it on the surface of the NCs as prepared. There are numerous reports that demonstrate functionalization of HP. HP was initially immobilized on the surface using (CaeP)/poly (hydroxybutyrate cohydroxyvalerate) (PHBV) nanocomposite scaffolds, which were produced first. CaeP NPs and PHBV were combined utilizing simply a solid-in-oil-in-water (S/O/W) emulsion solvent evaporation technique to create a CaeP/PHBV microsphere. The resulting composite was utilized to fabricate porous 3D scaffolds employing selective laser sintering techniques. The scaffolds’ surfaces were later modified in two processes. Initially, gelatine was entrapped on the corpora [54]. Alehosseini et al. developed tricalcium phosphate (Ca3(PO3)2) and other such calcium phosphate (CaeP)-based ceramics containing NCs. Silica was first utilized in the solegel method of production to stabilize TCP NPs. The as-prepared NPs were mixed with varying quantities of PCL in an electrospinning procedure to produce fibrous NCs. The bioactivity of NCs was subsequently increased by applying HP in varying amounts to the fibrous surface [55]. An alternate technique of manufacture involves covalently coupling bone morphogenetic protein-2 (BMP-2) to HP and plasminogen-free fibrinogen (BMP-2). In this process, HP and fibrinogen were covalently linked utilizing carbodiimide chemical properties. Thrombin and HPconjugated fibrin (HCF) gel were coupled to provide an injectable method for long-term BMP-2 administration [56].

15.5.4 Cosynthesis method It is common and effective to combine two polymers to create a new material with characteristics that neither of the component parts could have on their own. Additionally, BC (bacterial cellulose) may be used to make hybrid fibers. The development of NCs containing additional beneficial components is aided by BC, a naturally occurring polysaccharide produced by Acetobacter xylinum. The ultrafine microfibril form, that resemble the collagen network, explains this. Using a cosynthesis process, hybrid nanofiber was produced. In this technique, BC and HP solutions were combined in a growth medium where bacteria simultaneously produced both polymers, resulting in the development of a new nanofibrous scaffold [57].

15.5.5 Physical self-assembly method The constituent parts of the system, which could be substances, colloid substances, polymeric materials, or macroscopic molecules, tried to arrange into a sequence or an array of fully functional structure as a result of brief local interactions among both the constituent parts without having any external direction in a local process known as self-assembly. For

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example, Tan et al. produced scaffolds from bovine jugular veins from buffaloes using a sequence of detergent-enzymatic decellularization and dye-arbitrated photooxidation procedures. Additionally, they produced HP/CS NPs by a physical selfassembly procedure. The NPs were then placed onto the developed scaffolds to increase their loading capacity while retaining the scaffold’s biological activity [58].

15.5.6 Spontaneous emulsion solvent diffusion method In these approaches, the solvent is a dual mixture of an organic solvent which is water-miscible and an organic solvent that’s also water-immiscible. Here, that after solvent has evaporated, the NPs are prepared via emulsification process. Chung et al. used this methodology to produce DL-poly(glycolide-co-lactide) (PLGA) NPs functionalized with HP. In this procedure, a DMSO solution of PLGA was combined with a water solution of Pluronic F-127 to completely entrap HP on the surface of the NPs. These recently developed HP-functionalized PLGA NPs were composed of a hydrophobic DLPLGA core, a hydrophilic Pluronic F-127 outer surface, and HP being entrapped on the particles’ surface [59].

15.5.7 Coprecipitation and solvothermal method Iron oxide magnetic NPs coated by HP are frequently made using these synthesis techniques. FeC12-4H2O and FeCl3.6H2O being simultaneously added to a round bottom flask with stirring in the coprecipitation method. Iron oxide NPs were formed as a result of adding NH4OH dropwise here to solution to modulate the pH. HP was added to the synthesized NPs as an additional embellishment to enhance their bioactivity [60]. For the first creation of Fe3O4 magnetic NPs, another group used the solvothermal approach (magnetic nanoparticle [MNPs]). Then, to create MNPs@APTES, asprepared Fe3O4 MGNP was coupled with (3- aminopropyl)-triethoxysilane. CS and HP were used to further create such MNPs H APTES [61]. The gold NPs were stabilized by coating the generated NPs using glycol CS and Tween 80. Tween 80 and glycol were the surfactants used in this case. Similar to it though, Yuk et al. synthesized iron oxide NPs that use the coprecipitation method. To create iron oxide NPs containing gold deposits, citrate-stabilized gold NPs were first coated onto the as-prepared NPs because of the CS. HP was dissolved in water and then added to a colloidal solution to produce NC particles [62].

15.5.8 Other methods NPs cannot be employed directly for biological purposes, especially those made of metal or carbon. However, adding more bioactive substances might improve their biocompatibility and antibiofouling properties. Multiwalled carbon nanotubes and HP coupled to metallic NPs (such Au, Ag), to name just a few instances, are examples. Before HP binds to metal NPs on the surface during manufacturing, HP molecules are chemically activated with other groups, such as thiol, 2,6diaminopyridine, DOPA, and other physiologically active compounds. As a result, HP can attach to the NPs’ surface more readily [63e67]. Another study used CS that had been positively modified and negatively charged hyaluronic acid to make mucoadhesive nanocarriers by using the ionotropic gelation process. Following that, unfractionated or low molecular-weight HP was added to CS-HA NPs. Low molecular-weight HP-loaded NPs had a size of 152 nm, whereas unfractionated HP-loaded NPs were determined to be 162 and 217 nm in size [68].

15.6 Application of heparin-based nanocomposites HP is a biocompatible, natural polymer, which has been used as an anticoagulant for a very long period of time. The superiority of novel NC HG composites functionalized using HP is currently recognized as a result of their enormous improvement in nanobiotechnology [27,69]. In biomedical field including imaging, biosensing, drug administration, biomedical engineering, etc., such materials are frequently employed. They incorporate HP in certain way.

15.6.1 Drug delivery Although taking HP orally would be ideal, HP is poorly absorbed by the digestive system and has a short half-life [70]. The low bioavailability of HP is attributed to its large molecular weight, negative charge, attraction to negatively charged mucosal and epithelial cells, and difficulties in crossing membranes [71]. In order to enhance the bioavailability and half-life of this essential medication, many strategies have been investigated. The use of NPs as HP carriers is one area that is being studied. The oral bioavailability of HP conjugated to positively charged nonbiodegradable polymethacrylates, Eudragit RS and RL, as

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well as polycaprolactone and poly(lactic-co-glycolic acid) (PLGA), was studied by Jiao et al. These multiemulsion polymeric NPs, with diameters in the range from 260 to 300, were used to entrap HP. The above NP composites were evaluated for their anticoagulant efficacy both in vitro and in vivo. Comparing orally delivered free HP to HP-loaded NPs, the researchers found that the latter better maintained their anticoagulant activity. They came to the conclusion that the HP was protected from deterioration by the NPs and that it was progressively released from the particles in an undamaged state. With the use of a resonant mirror system and Caco-2 cells, Lamprecht et al. looked into the mechanisms behind HP NP interactions with mucin and epithelial cells. The nonbiodegradable polymer Eudragit RS and the biodegradable PLGA were two NPs on which HP was adsorbed that were similar to those Jiao used. NPs may initially interact with the negatively charged mucosal layer on epithelial cells, according to a two-step procedure Lamprecht et al. deduced from adhesion experiments [72]. The HP medicine spent more time in contact with the epithelial barrier due to the NPs’ mucoadhesive properties. A negatively charged glycoprotein called mucin may then push HP off the particle’s surface, enabling the drug to enter the body. In order to increase HP absorption with oral administration, it may be necessary to forego frequent invasive injections by discovering the mechanism of adhesion that accounts for the improved bioavailability. In 2015 alone, cancer claimed 8.8 million lives, making it the leading cause of death worldwide [73]. Cancer is a common disease nowadays. As a result, there is an increasing need to find novel therapies that go beyond the ones that are now available (surgical, chemotherapeutic, radiotherapy, targeting, photothermal therapy, and immunotherapeutic). Additionally, laser therapy, blood transfusion and donation, hyperthermia, photodynamic and photothermal therapy, stem cell transplants, and photodynamic and photothermal therapy are used. Cancer treatment and cancer diagnostics both employ HP-based NCs [74]. There have been attempts to use HP as a component of a multifaceted nanosystem for the cancer treatment and diagnosis because it is biocompatible and is effectively absorbed into the body as a result of its biological interactions with proteins, growth factors, chemokines, cytokines, enzymes, and lipoproteins, which are all engaged in various biological processes. There have been prior investigations on the effects of HP alone in the treatment of cancer [75]. The anticancer medications doxorubicin, docetaxel (Taxotere), paclitaxel (Taxol), and sorafenib are a few of them and have all been utilized in nanosystems intended for therapeutic purpose [76].

15.6.2 Tissue engineering In the multidisciplinary field of tissue engineering, components of damaged or poorly functioning human organs or tissues caused by diseases or trauma are advanced, intensified, and restored using concepts from several scientific disciplines such as chemistry, biology, engineering sciences, etc. In these treatments, a scaffold is implanted together with the appropriate bioactive substances and cells to promote the formation of new tissues and organs. The creation, adhesion, and demarcation of cells are supported by scaffolds, which are 3D structures. For tissue regeneration, it is essential. A scaffold should be bioactive, compatible, and degradable by nature. Additionally, the by-product of its dilapidation ought to be nonflammable, nonhazardous, and simple to remove from the spot of implantation. As a result, there is a lot of interest in biomaterials that can replicate the biological, mechanical, and structural traits of actual tissues. Notable advancements have been made in the production of new materials as well as in the process of supporting healthy cell activity [77,78].

15.6.2.1 Role of nanocomposites based on heparin in tissue engineering To fully comprehend tissue engineering, one needs to be familiar with the fundamental cellular milieu architecture. Cellular activities, in general, and tissue healing need the presence of cellular milieus like receptors, growth hormones, heparan sulfate, ECM, cellecell interaction, adhesion protein, and integrin. Cells use the ECM to access the plethora of signals coming from the outside environment during tissue healing. Protein fibers like fibrillar collagen and elastin, whose sizes range from 10 to several hundred nanometers, make up the intricately woven meshes that make up the ECMs. Laminin and fibronectin, two nanoscale adhesion proteins, are present in the ECM fibers and are used to bind cells at particular places. In addition, these adhesion proteins communicate with the cells through integrin. Cell migration, differentiation, growth, and shape are all governed by these ECM structural components. The connection between cells and the extracellular microenvironment (EME), that enables cells to adapt to the demands of the environment, is crucially mediated by ECMs [79]. The EME typically consists of three types of effectors: (a) soluble macromolecules (growth factors, for example); (b) soluble macromolecules (glycosaminoglycan side chains); and (c) proteins located in the EME. The cytoskeleton and the ECM are tightly connected by these effectors. Cells interact with the mechanical qualities of their surroundings in order to receive mechanical impulses, and they then respond with chemical messages, as shown in Fig. 15.4. The signals produced by the ECMs’ effectors make the final choice for processes involving cells, including decisions on whether they need to differentiate, multiply, migrate, undergo apoptosis, or perform other specific tasks (Fig. 15.5A). These cell

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FIGURE 15.5 Representation of (A) cells process and (B) tissue formation. Adopted with permission from the Barik S, Sharma RK, Rath C. Heparinbased nanocomposites for tissue engineering. Polysaccharide-based nanocomposites for gene delivery and tissue engineering. Woodhead Publishing; 2021. p. 81e101.

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functions are crucial for tissue repair or development (Fig. 15.5B). It is essential to construct synthesized scaffolds for bioengineering such that they function similarly to that same natural ECM. It is also important to point out that only a limited number of solid materials, which result in the construction of the structural scaffold, are needed to build mechanically robust structures of ECMs. Resistant to tensile and compression stresses is made possible by the fibrils and hydrated networks of the structural scaffolds and interstitial fluid, respectively [80]. Thus, these biophysical characteristics regulate a number of crucial cellular processes, including as adhesion and migration. Researchers are urged to replicate natural ECM activities and create synthetic ECM for bioengineering because of the beneficial properties of ECM. In synthetic scaffolds, the biophysicochemical operations of the natural ECM have been modeled. The primary difficulty, however, is that the ECM’s composition and spatial structure differ depending on the makeup and function of the tissues [40,52,81e84]. Growth factors are essential for engineering tissues that include creation or regeneration. For instance, during skin regeneration, keratinocyte growth factor, fibroblast growth factor, vascular endothelial growth factor, and interleukin 1 synthesis drives cell proliferation, macrophage activity, and angiogenesis. Contrarily, BMP-2 is a growth factor that stimulates macrophages and endothelial factor (VEGF) and is involved in the migration and regeneration of bone and cartilage. The biological and chemical properties of HP have lately been used to further alter the bioactivity of scaffolds created in order to optimize its hemocompatibility with locally damaged tissues. But in addition to serving as a carrier for various growth factors, HP may also be employed to provide growth factors for a long time at the site of repair. Sulfation is thought to enhance HP’s interaction with proteins and cells. As a result, growth hormones including BMP-2, VEGF, and FGF have been introduced to HP-based NCs for tissue engineering [33,39,55,60,85e88]. Growth factors are combined with HP to enhance their capacity to interact with growth factor receptors and protect them against chemical and proteolytic inactivation. Additionally, HP and VEGF cooperate to promote angiogenesis. Additionally necessary for supplying oxygen and nutrients during scaffold regeneration are blood vessels. In this regard, VEGF functions as a growth factor in the development of vascular tissues. However, a large dosage is necessary due to its short half-life (50 min). HP is utilized to increase its shelf life and local distribution. In order to rebuild scaffolds, VEGF promotes the growth of endothelial cells as well as the maturation and creation of blood vessels at the location of the wound. In this way, tissue engineering applications typically employ HP-based NCs. HP attaches to BMP-2 similarly to how it does in bone tissue engineering. A bone regeneration growth factor called BMP-2 that promotes bone repair and has exceptional clinical effectiveness. BMPs encourage mesenchymal stem cells to migrate to locations where new bone is being formed during bone regeneration. These mesenchymal stem cells undergo additional differentiation to become osteoblast lineage, which helps to regenerate bone. Since normal fibrin generates BMP-2 for local usage, BMP-2 has a limited half-life even in this case. As a result, a significant dosage is administered during clinical therapy. For a longer period before BMP-2 is released, HP is utilized to bind regular fibrin [34,57,59,89,90]. NCs based on HP are commonly used in tissue engineering. Based on a detailed analysis of the previous research, HPbased NCs are used primarily for the engineering of vascular and bone structures. Fig. 15.6 provides a schematic representation of skeletal and vascular tissue restoration. Table 15.1 provides some illustrations of tissue engineering applications using HP-based NCs. To create an electrospinning elastomeric nanofiber fabric, Kwon and Matsuda combined PLCL, salt of TBA-HP, copolyester, and other materials. According to their joint assessment of the fabric’s capacity to encourage cell proliferation, the fabric they created is appropriate to be employed as a scaffold in vascular tissue engineering. Chung and his coworkers fabricated HP-PLGA NPs by utilizing a solvent diffusion approach without undergoing any chemical modifications to the component components. They were able to detect changes in the mean size and surface charge of the NPs by altering the HP% during the manufacturing process. Due to its linear and full liberation profile free of any early burst, they went on to study the growth factor release from HP-PLGA NPs and concluded that the material may be employed as a growth factor release component in tissue engineering scaffolds. Using PLGA microsphere synthesis, Na et al. created encrusted biocompatible HP/poly(L-lysine) NPs. The degree of immobilization of the generated NPs on to PLGA has been evaluated by confocal laser microscopic examination. They also discovered that combining human mesenchymal stem cells with PLGA microspheres causes in vivo cell proliferation, differentiation, and proliferation of cells, as well as the synthesis of the ECM. These findings suggest that the information generated can deal with issues. They also looked at the growth factor release from HP-PLGA NPs, which exhibits a linear, full liberation profile devoid of any early burst, demonstrating the material’s appropriateness as a growth factor release component in tissue engineering scaffolds [59]. In vivo cell division, tissue formation, and the synthesis of ECM were also shown when human mesenchymal stem cells and PLGA microspheres were combined, according to findings from a cell culture study. A layer-by-layer manufacturing method was created by Na et al. [91] that depends on the potent ionic charges of the different polymers and chemicals utilized. Their preliminary study used positively charged PLL and HP to produce NPs. A coating of poly(ethyleneimine) (PEI), which

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FIGURE 15.6 Graphical representation of (A) vascular tissue repairing mechanism and (B) bone tissue repairing mechanism. Adopted with permission from the Barik S, Sharma RK, Rath C. Heparin-based nanocomposites for tissue engineering. Polysaccharide-based nanocomposites for gene delivery and tissue engineering. Woodhead Publishing; 2021. p. 81e101.

could interact electrostatically with the HP-PLL NPs, was subsequently applied to the PLGA microspheres. The HP-PLL NPs are coated on the microsphere’s surface to a 70% extent. Rai et al. used nanohydroxyapatite (n-HA) as the support for stacking HP, then poly(L-histidine) (PH), followed by another layer of HP in a layer-by-layer construction process for load-bearing bone applications [92]. PLGA, polyethylene glycol, growth factor, and HP were the materials employed by Park and his colleagues to create 3D nanostructured scaffolds (L-lysine). Following physical attachment of HP/poly(L-lysine) NPs with growth factor incorporation to the positively charged surface of PLGA microspheres, low MW PEI coating was applied layer by layer. The produced NPs were said to be evenly distributed across a substrate and to have a biocompatible composition. Due to the biochemical activity, they display in cells, they may also help with a variety of problems associated with tissue regeneration [93]. Duan and Wang developed a tricalcium-phosphorus (hydroxybutyrate-co-hydroxy valerate) nanosheets scaffold using the selective laser sintering technique that has a distinctive structural layout and restricted linked porosity. They investigated the cytocompatibility of the NCs. HP was then immobilized inside the trapped gelatine after they changed the material’s surface by trapping it inside its porous structure. On this changed surface, there is a distinct binding site for growth factor and conjugated HP. The material created has shown promise in bone tissue engineering applications [54]. In bone healing, bone morphogenetic proteins (BMPs) typically perform as the best osteoinductive growth factors. Yang et al. produced an injectable device that covalently conjugated HP to fibrinogen and released BMP-2 over time. It was shown that the BMP-2 release from the produced chemical was superior to fibrinogen when mixed with thrombin to form HCF gel. Furthermore, it was shown that HCF’s release of BMP-2 significantly increases the level of alkaline phosphatase activity in cultured osteoblasts. This demonstrates that long-term liberation is superior to short-term liberation for bone regeneration and that BMP-2 released from HCF is physiologically active. For the engineering of bone tissue, the material performs well [56]. NPs are used as fillers to enhance the mechanical qualities of current bone-engineering scaffolds. However, NPs have a propensity to aggregate, which leads to the loss of nanoscale properties. In a model system, immobilizing PH (representing growth factors or adhesion proteins) and HP (for stability and localization) enhanced n-HA dispersion and stability. The nHA supports appear to have absorbed the negative charges from the HP molecules. After the initial layer of HP was immobilized, the composites’ zeta potential showed that Ca2þ was released, proving that specific carboxyl and sulfo groups, rather than general electrostatic interactions, were how HP interacted with the n-HA. The negatively charged HP layer was quickly absorbed by the positively charged PH molecules, and the subsequent layers became more electrostatically linked. The addition of HP induced charge repulsion, which made the particles stable for 23 days. For NPs in bone-engineered scaffolds to have better mechanical properties, this stability is essential. These n-HA particles may have

244

Composite preparation method

Application

References

Elastomeric nanofiber fabrics comprising copolymer, PLCL, along with TBA salt of HP

Coelectrospinning

Vascular tissue engineering

[37]

2

HP-PLGA NPs

Solvent diffusion

Tissue engineering

[59]

3

HP/poly(L-lysine) NPs-coated PLGA microsphere

Layer-by-layer assembly

Suitable for tissue regeneration

[91]

4

Bionanohydroxyapatite particles

Layer-by-layer assembly

Bone engineering scaffolds

[92]

5

Growth factor-loaded HP/poly(L-lysine) NPs with polyethyleneimine-precoated PLGA microspheres

Layer-by-layer assembly techniques

Suitable for tissue regeneration

[93]

6

Three-dimensional calcium phosphate/PHBV NC microspheres

Selective laser sintering

Bone tissue engineering

[55]

S.No.

Name of the system

1

7

HP-conjugated fibrin gel

Coprecipitation

Bone tissue engineering

[56]

8

HP nanoparticle-immobilized decellularized bovine jugular vein scaffold

Physical self-assembly

Cardiovascular tissue engineering

[58]

9

HP-bacterial cellulose hybrid nanofibrous scaffold

Cosynthesis process

Vascular tissue engineering

[57]

10

HP-conjugated poly(ε-caprolactone) tubular scaffolds

Electrospinning

Vascular tissue engineering

[94]

11

Nanoassembly comprising CS electrospun fibers, HP containing PCNs, and HP containing PEMs

Dip coating

Tissue engineering

[38]

12

HP-succinylated gelatin NPs

Microemulsion technique

Bone tissue engineering

[87]

13

HP laden PCL-a-TCP fibrous NC membranes

Electrospinning

Bone tissue engineering

[55]

14

Multistructured vascular patches of PEM by deposition of HP and CS alternately onto PU/DCS

Layer-by-layer assembly techniques

Vascular tissue engineering

[52]

15

HP-grafted PCL/gel-PCL/gel/n-HA bilayered fibrous membrane

Electrospinning

Tissue engineering

[40]

Adopted with permission from Barik S, Sharma RK, Rath C. Heparin-based nanocomposites for tissue engineering. Polysaccharide-based nanocomposites for gene delivery and tissue engineering. Woodhead Publishing; 2021. p. 81e101.

Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine

TABLE 15.1 An HP-based systems with applications in tissue engineering.

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osteoinductive properties by substituting PH with other positively charged growth factors or adhesion proteins, making them helpful in different orthopedic implants and bone applications [95].

15.7 Conclusion and prospective The application of NCs based on HP is an interesting subject because even though the material in question was found several years ago, it has been upgraded, and it is still being researched in order to reveal its capabilities and its structure. Additionally, in addition to its antiinflammatory and antithrombotic properties, research is being done on potential novel applications for it. The therapy of diseases like cancer is by far the most significant and relevant use, followed by tissue engineering, biosensing, and detection as the next most important and pertinent application. As a result, the range and complexity of possible applications are impressive and enticing. HP is a natural or synthetically acquired biocompatible substance from which HP-based NCs can be prepared. Its variants are also helpful since they were developed in order to increase their applications in the medical field or to get around any drawbacks or hazards. When it is used for the synthesis of NPs, it can be chemically modified in several different ways, such as through conjugation and cross-linking, to produce nanobiomaterials. These nanobiomaterials have the potential to be used for a variety of applications because they have certain functionality that is tailored to a particular goal. In addition, HP can be employed in the production of NPs without undergoing any modification at all. Some in vitro and in vivo investigations for medical applications have noted that in the near future, HP-based solutions could be a bright choice for the development of additional alternatives to treat, detect, and prevent diseases in humans. These studies have been carried out for medical applications. On the other hand, in terms of imaging, detecting, and biosensing, in spite of the fact that there are not many studies on these topics, therefore, these areas have the potential to undergo more research and development in order to provide highly helpful materials. Moreover, for the production of nanobiomaterials, HP-based NPs could be incorporated into already present materials to improve their antibacterial and antifungal activities. This is especially important in light of the fact that numerous microorganisms have established a high level of resistance to the common antibiotics that are currently in use. Additionally, the possibility of improvement in the techniques like detection and sensing is possible by the use of HP. This is because of the structure of HP, that even at very low concentration, it can interact with materials, like dyes and biomolecules. This is because the structure of HP allows it to interact with other materials. The conjugated and cross-linked HP-based NPs are the ones that are most relevant because, although they are typically synthesized by incorporating a variety of materials, each of these materials plays a vital role in the specific application.

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Chapter 16

Hydrogels based on chondroitin sulfate nanocomposites Leena Kumari1, Kalyani Sakure2 and Hemant Ramachandra Badwaik3 1

School of Pharmacy, Techno India University, Kolkata, West Bengal, India; 2Rungta College of Pharmaceutical Sciences and Research, Bhilai,

Chhattisgarh, India; 3Shri Shankaracharya Institute of Pharmaceutical Sciences and Research, Bhilai, Chhattisgarh, India

16.1 Introduction Multicomponent systems made up of 3D polymeric networks are known as hydrogels (HGs). There are two major compartments in this network: polymer chains and water molecules, which are found in the intervals between polymer chains. HGs have been extensively employed in drug delivery and biomedical applications, including nanomedicine and tissue engineering owing to their unique features [1,2]. HGs are porous 3D structures that are chemically or physically crosslinked. Because of their thermodynamic affinity for solvents and porous structure, they possess the properties of absorbing water. Another intriguing aspect of HGs in terms of drug delivery systems is their capacity to respond to various internal and external stimuli such as pH, ionic strength, electromagnetic field, temperature, electrostimulated, and so on. Smart materials are a type of materials whose biological applicability as drug delivery systems is directly influenced by the type of stimuli they respond to. Refs. [3,4]. In addition to these advantages, HGs also have some disadvantages. For instance, many HGs limited capacity for drug loading and potential for premature disintegration or carried away from the targeted site are both caused by their poor tensile mechanical strength. The only drugs that might be facilely incorporated into gels are those that are hydrophilic; however, loading hydrophobic drugs might cause issues including inadequate drug loading, poor drug homogeneity, and so forth. The initial quick release of drugs from specific HGs is another aspect to take into account. The large pore sizes and high water content of the majority of HGs promote the burst release of drugs. In addition, the size of HGs makes them difficult to administer as injectables in some situations, necessitating surgical implantation [5]. The injectability mechanism in HGs with shear thinning properties can be demonstrated. These ex vivo-generated injectable gels have the ability to flow when subjected to shear stress and recover their gel structure following relaxing. Returning to an unharmed primary form within a physiological site is the process of self-healing. The essential requirements for a therapeutic application can be challenging to fulfill in a single biomaterial (for instance, a basic polymer). A potential solution to the issue is the inclusion of fillers in the hydrogel. The behavior of the final composite is a combination of the filler and matrix characteristics, or perhaps recently found synergistic characteristics. Typically, the matrix is composed of metal, ceramic, or polymer. A fiber, particle, or laminate can be used as filler, also known as reinforcement or reinforcing agent [6]. However, composites reinforced with laminates are not appropriate for injectable HGs. The composite is referred to as a nanocomposite (NC) when the filler is a nanosized or nanostructured material, such as nanoparticles (NPs) or nanofibers [7]. The use of nanoscale compounds provides important size-related benefits, such as a high surface-to-volume ratio, which improves matrixefiller interactions that result in improved attributes like mechanical strength. The fast release of hydrophobic drugs and limited loading ability are the drawbacks of HGs as drug carriers [8,9]. NPs have been focused a lot by the researchers for loading insoluble pharmaceuticals; as a result, NC HGs were designed to unite the advantages of both NPs and HGs. The properties and applications of various NC HGs based on natural and synthetic polymers are shown in Fig. 16.1. The choice of polymers to be employed in the synthesis of hydrogel is critical in terms of their composition, as biocompatibility is a critical aspect for living organisms. Polysaccharides are one of the most prevalent natural polymers employed in the manufacture of HGs, and they can be found in a variety of places, including microorganisms, animals, and plants [10]. Because of their high biocompatibility, numerous functional groups, and capacity to immobilize biomolecules, Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine. https://doi.org/10.1016/B978-0-323-95351-1.00006-5 Copyright © 2024 Elsevier Inc. All rights reserved.

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FIGURE 16.1 The properties and applications of NC HGs based on natural and synthetic polymers. Reprinted from Wahid F, Zhao XJ, Jia SR, Bai H, Zhong C. Nanocomposite hydrogels as multifunctional systems for biomedical applications: current state and perspectives. Compos B Eng 2020;200:108208, Copyright (2020), with permission from Elsevier.

polysaccharide-based HGs are potential biomaterials for applications in wound healing and tissue regeneration [11]. Owing to their high water content and soft, rubbery consistency, which resembles living tissue, polysaccharides are typically nontoxic, have a high biocompatibility and biodegradability, and possess several distinct physicochemical properties that make them ideal for a variety of drug delivery applications [12]. The NC HGs made from polysaccharides are nontoxic, biodegradable, and cytocompatible. They can also be mixed with a variety of materials (organic and inorganic) to form hybrid HGs [13,14]. Although chitosan [15] and alginate [16] are the most commonly used polysaccharides in the production of HGs, chondroitin sulfate (CS), an animal-derived polysaccharide found mostly on the cell surface and in cartilages, is also suitable for these applications. When opposed to synthetic polymers, CS has a lot of advantages. In connective tissues and cartilages, it has a crucial structural role. It gives connective tissues compressive strength by managing their water content, and it has properties that make it excellent for bioapplications, such as biodegradability, multifunctionality, and high water absorption [17,18]. CS cannot be employed as a solid-state drug delivery system because it is water-soluble. As a result, it is often mixed with other polymers like collagen, hyaluronan, gelatin, chitosan, poly(vinyl alcohol), and poly(lactic-co-glycolic acid), to develop materials with more stable characteristics, as described in various research [17,19]. The numerous research works related to HGs are explored extensively. The current review examines the primary methods for making CS-based NC HGs, as well as the applications of these HGs in drug delivery and other biomedical fields. As a result, we shed light on present trends and future prospects, with a focus on drug delivery. This book chapter is expected to lead to new directions in the design, production, and use of CS-based NC HGs.

16.2 Structure and bioactivities of chondroitin sulfate CS is a glycosaminoglycan (GAG) found in the extracellular matrix (ECM) of human bone that is noncollagenous. The carbohydrate chain of CS is made up of repeated disaccharide units of b-1, 3-linked N-acetyl galactosamine (GalNAc) and b-1, 4-linked D-glucuronic acid (GlcA) [20,21]. The structure of CS is shown in Fig. 16.2. It is a combination of acid mucopolysaccharides derived from pig, shark, chicken, cattle, and cartilage. According to the sulfide location and degree, CS is commonly categorized into CS-O, CS-A, CS-C, CS-D, and CS-E with acetyl GalNAc and GlcA replacing one or more sulfate groups. D-GlcA and 2-hydroperoxy group-deoxy-D-galactose are found in equal proportions in both CS-A and CS-C, as well as acetyl and sulfuric acid residue. The only variation between them is the sulfuric acid ester’s location on the hexosamine residues. Aside from sulfate group diversity, the biological activities and pharmacological characteristics

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FIGURE 16.2

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Structure of CS. Chondroitin-4-sulfate: R1 ¼ H; R2 ¼ SO3H; R3 ¼ H. Chondroitin-6-sulfate: R1 ¼ SO3H; R2, R3 ¼ H.

of these several forms of CS are extremely diverse [22]. The bioactivities of different sulfate groups are vastly different. Proteoglycans high in CS-D and CS-E, for instance, have been shown to aid in neuro-repair, but the converse is true for CS-A and CS-C. The sulfate group discrepancy is mostly influenced by origin, such as mammalian, marine, and bone resource [23]. For example, CS obtained from porcine has a larger ratio of CS-A/CS-C than CS derived from bovine, and CS derived from marine organisms has a higher proportion of disulfide group. Furthermore, the sea cucumber’s source is much liable to produce fucosylated CSs. CS is a complex macromolecular sulfate polymer derived from natural resources. As a significant element of the ECM, it maintains and protects the extracellular microenvironment. It is also a multifaceted signal molecule and regulator that affects cell signaling directly or indirectly and is engaged in a variety of physiological functions [24]. Anticoagulant, antiinflammatory, antitumor, antioxidant, and immune-enhancing properties are some of the bioactivities of CS. As a result, CS has a diverge spectrum of therapeutic uses. For example, by decreasing the release of TNF-a and CXCL8 in cultured human mast cells activated by IL-33, CS can be utilized for the treatment of allergy or inflammation [25]. In addition, it has been used to treat cancer, angiocardiopathy, and osteoarthritis [23,26,27]. It has been discovered that CS is linked to schizophrenia and axonal regeneration [28,29]. At the same time, due to its nontoxicity, biodegradability, biocompatibility, no side effects, and anionic characteristics, its material use has become a research hotspot.

16.3 Method of preparation of nanocomposite hydrogel systems 16.3.1 Crosslinking method One of the most popular ways to develop hybrid NC HGs is to combine hydrogel precursors with colloidal nanoparticle suspensions before the preparation of hydrogel. This technology frequently offers relatively uniform integration compared to other assembly approaches, including nanoparticle inwards diffusion that may produce poorly defined nanoparticle gradients [30]. Alternatively, it is feasible to manipulate nanoparticle location during crosslinking of hydrogel in some cases (e.g., by applying magnetic fields), allowing bioinstructive gradients to be designed along the 3D constructions [31]. Because NPs can be preassembled prior to hydrogel crosslinking, this technology permits a large range of NPs to be incorporated in the HGs. Conversely, this flexibility allows researchers to create complex and advanced NPs without jeopardizing the formation of hydrogel, as long as the crosslinking chemistry and circumstances used do not have a substantial impact on nanoparticle stability and physicochemical qualities. The presence of the specified nanomaterials, on the other hand, should not interfere with the crosslinking of the hydrogel matrix. Another significant benefit of this technology is that it may be used to construct cell-laden NC HGs employing cytocompatible biomaterials and appropriate crosslinking conditions. Increases in nanoparticle content, which may alter hydrogel network development and increase the possibility of particle-particle aggregation, are a significant disadvantage of this method. Standardizing production and generating repeatable reactions to input stimuli necessitate the development of homogeneous NC HGs.

16.3.2 In situ method By producing in situ nanoparticle synthesis after hydrogel crosslinking, undesirable particleeparticle aggregation can be prevented and hybrid platforms with high nanocarrier content could be constructed. The restriction is that the circumstances necessary for hydrogel crosslinking shouldn’t have any impact on the generation of NPs. This method is especially

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beneficial for putting together electro-sensitive NC HGs. This assembly process has also been used in designing tissue adhesive, electrically conductive NC HGs that could be used to create paintable regenerative cardiac patches [32,33]. This design included a porcine gelatin hydrogel platform that was coupled with dual-functionalized dopamine and pyrroledependent hyperbranched polymer. Similar to the earlier study, Fe3þ caused the production of polypyrrole NPs in situ inside the template of the hydrogel, as well as the polymerization of polymer-tethered pyrrole and dopamine moieties, strengthening the hybrid network. Furthermore, by applying wet-adhesion catecholamine chemistry, these NC HGs could quickly attach to humid beating heart surfaces, enhancing electrical signal transmission, and revascularizing infarcted myocardium [34]. This assembly technique has thus been extensively employed to produce electrically conductive NC hydrogel platforms and is suitable for preventing nanoparticle aggregation. Despite this, the limitation on the amount of components that may be assembled into NPs under these circumstances is a drawback of this process for developing NC HGs (e.g., pyrrole, calcium NPs). Such restrictions limit the types of NPs that can be included, limiting their application. This method drastically restricts the encapsulation of viable cells and exposes living cells to potentially harmful monomers before the network is formed. Furthermore, it may be challenging to achieve homogenous nanoparticle production throughout the whole build volume when building thicker NC hydrogel platforms due to insufficient catalyst dispersion or penetration into the hydrogel matrix [35].

16.3.3 Inwards diffusion method Another method involves loading preformed NPs into hydrogel networks after crosslinking, which takes advantage of the porous characteristics of hydrogel networks and their capacity to spontaneously include particles of size less than the matrix pores [36]. The conventional approach to achieve the integration is swelling of hydrogel in concentrated solution of nanoparticles. One of the main advantages of this method is the separation of hydrogel and nanoparticle production, which prevents aggregation brought on by the existence of catalysts, stimuli, or specific precursors during hydrogel formation as well as interaction between nanoparticle species and hydrogel crosslinking [37]. The frequent generation of nanoparticle concentration gradients from surface to core sections, with diffusion dependent on hydrogel pore size, NPs species properties, and other physicochemical characteristics, is among the major downsides of this method (i.e., hydrophilicity and overall charge).

16.3.4 Codependency and crosslinking method In addition to working as network filler agents in distinct NC HGs, NPs are ideally suited to be used as crosslinking elements in such three-dimensional frameworks [38]. Technological advancements in biomaterial design, bioconjugation, and polymer chemistry have made it possible to tailor nanoparticle surfaces with specific functional moieties to enhance interactions with hydrogel precursors and target compounds, resulting in NC HGs via a codependent assembly approach. In addition, such interactions between nanoparticles and hydrogel are excessively variable on account of the durability or flexibility that might be retained in the resulting hybrid mesh, due to the nature and majority of covalent, noncovalent, or dynamic covalent connections that are produced. Codependent NC hydrogel crosslinking density and the release profile of bioactive substances contained in the hydrogel or NPs are also changed when the nanoparticle content is changed. A connection of this kind is particularly desirable for expanding the therapeutic range and pharmacokinetics. In order to improve the mechanical stability and stimulus sensitivity of NC hydrogel networks with many conjoined connections, NPs may also contain multivalent binding motifs on their surfaces [39].

16.4 Drug delivery, biomedical, and other applications of chondroitin sulfate nanocomposite hydrogel 16.4.1 Drug delivery applications The chondroitin sulfate NC HGs have extensive uses in the area of drug delivery. The CS-based NC HGs have also been used to administer a variety of drugs under regulated conditions. For instance, hybrid HGs including NCs for the regulated delivery of two drugs, L-dopa and vitamin B12, were made using HGs composed of 50% CS, 50% casein, and varying concentrations of SiO2. The addition of 5% SiO2 into the materials allowed for the best arrangement, distribution of the pores, which are in charge of giving a more regulated release. Additionally, the matrices released L-dopa well but were inefficient at releasing vitamin B12. L-dopa is less polar and interacts more with the CAS/CS matrix than vitamin B12

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because it is more hydrophilic and interacts with the medium [40]. In another research work, an NC system with a hydrogel shell and a prodrug core loaded with two drugs at the same time was developed to improve cancer sensitivity to chemotherapeutics. A self-assembled prodrug of PTX coupled with CS was used to incorporate sunitinib and PTX at the same time. Both the free drugs and drug-loaded micelles were encapsulated in a CS hydrogel shell, and a glutathione Stransferase-mediated disulfide link was established in the shell, allowing the shell to degrade in an acidic medium of the tumor. When the shell was ruptured, the two drugs contained therein were immediately released, having a synergistic antitumor effect. It was followed by the slow release of the pharmaceuticals contained therein, which kept drug levels at therapeutic levels in treatment-resistant tumors. The CS released from the destroyed shell suppressed Bcl-XL, making the tumor more susceptible to treatment. An appropriate way to distribute drug to tumors that are resistant to therapy is with a programmable NC system with a heterogeneous release profile and tumor sensitization augmented by CS [41]. In another study, natural NCs made using bacterial cellulose for functional materials were described. Fermentation alterations in gel bacterial cellulose with CS and hyaluronic acid were accomplished, and their cell behavior is given in order to develop scaffolds with drug delivery capabilities, porous structure, and enhanced cell adhesion. For the first time, a viability and cytotoxicity investigation with stem cells is provided employing gel bacterial cellulose scaffolds for regenerative medicine. MTT (3-(4,5-dimethylthiazolyl-2)-2,5-diphenyltetrazolium bromide) viability studies demonstrate that both gel bacterial cellulose samples have higher cell adhesion with time, while LDH (lactate dehydrogenase) assays show that both samples have low cytotoxicity [42]. In another investigation by Simao et al. [43], the polymeric materials were created by chemically crosslinking silica (SiO2) nanospheres, casein (CAS), and CS, resulting in a highly crosslinked network (Fig. 16.3). By adjusting the polymeric proportion, the hydrogel release profile may be customized. The addition of 5% silica nanospheres in mass to all CAS/CS matrices resulted in a more regulated and prolonged release of L-dopa, with the exception of the matrix containing 5% silica, 30% CS, 70% CAS, and whose L-dopa was released up to 87 h. HGs are also cytocompatible. These novel HGs are highly appealing materials that could be employed for regulated and retarded drug release, and also for scaffolds and wound dressing. By dissolving the CS in deionized water to create hydrogels, sodium alginate (SA) was employed to coat the CS. This was done to make sure that CS would transit via the stomach and get released in the intestinal tract, increasing its potency as an anticancer immunopotentiator. To create CS/SA/chitosan composite nanoparticles, chitosan was then electrostatically deposited outside the hydrogels. To develop nanoparticles having a size range less than 200 nm, it is vital to consider how the component content and process factors affect the size and swelling tendency of the composite micro/nanoparticles. The results showed that at pH 6.8, as compared to pH 1.2, the cumulative release rate of these composite nanoparticles was greater with a weight ratio of SA to CS of 2:1. The CS/SA/ chitosan composite particles were developed by optimizing the process parameters, allowing the majority of the CS to pass through the stomach and release in the gastrointestinal (GI) region [44].

16.4.2 Tissue engineering Biopolymer-based HGs are known for their biocompatibility and potential to mimic ECM structure to promote cellular function. These HGs, on the other hand, have poor mechanical characteristics, uncontrolled breakdown, and minimal

FIGURE 16.3 Illustration of the synthesis of SiO2 nanoparticles-incorporated hydrogel based on chondroitin sulfate. Reprinted from Simão AR, Fragal VH, de Oliveira Lima AM, Pellá MC, Garcia FP, Nakamura CV, et al. pH-responsive hybrid hydrogels: chondroitin sulfate/casein trapped silica nanospheres for controlled drug release. Int J Biol Macromol 2020;148:302e15, Copyright (2020), with permission from Elsevier.

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osteogenic activity, limiting their use in bone regeneration. As a result, bioactive fillers for bone regeneration were produced using hybrid gelatin (gel)/oxidized CS (OCS) HGs incorporating mesoporous bioactive glass nanoparticles (MBGNs). The addition of MBGN into the HGs had no effect on their injectability. The osteogenic differentiation and proliferation of rat bone marrow mesenchymal stem cells in vitro, as well as the restoration of rat cranial defects in vivo, were also enhanced by the HGs in the presence of MBGNs. The hybrid gel-OCS/MBGN HGs showed great osteogenic properties and mechanical characteristics, as well as predictable gelation and degradation activity, making them an excellent choice for scaffolds or injectable biomaterials for bone regeneration/repair applications [45]. Bioinks based on self-healing and shear-thinning HGs have gained increasing attention because they can be used to create sophisticated three-dimensional physiological microenvironments. While protecting cells from shear stresses throughout extrusion-based bioprinting, it is challenging to provide excellent structural stability and clarity of printed components. HGs made up of laponite (LA), silicate nanomaterials, and CS-based glycosaminoglycan NPs (GAGNPs) for bioprinting applications that are shear-thinning and printable. NC HGs (GLgels) were quickly formed as a result of the reversible interactions among the edges of LA and the negatively charged groups of GAGNPs (Fig. 16.4). During bioprinting, the hydrogel’s shear-thinning characteristic shielded encapsulated cells from strong shear forces. The bioinks might be easily printed into free-standing, shape-persistent structures with large facet ratios. The potential of GLgels to stimulate cell growth and proliferation was validated in vitro tests. Alkaline phosphatase (ALP) activity and calcium deposition were used to establish in vitro osteogenic development of preosteoblasts murine bone marrow stromal cells entrapped within the GLgels. The GLgel’s in vivo biocompatibility and biodegradability were validated in rats after subcutaneous implantation. The osteoinductive properties of the designed shear-thinning hydrogel can be employed as a novel bioink for 3D printing constructions for bone tissue engineering [46]. Scaffolds made of bioconjugated HGs are appealing for tissue engineering because they may replicate the features of human tissue to some extent. They can, for example, boost their bioactivity with cells. However, most HGs have processability issues, which limit their usage in 3D printing to create custom scaffolds. Through a microextrusion procedure, bioconjugated hydrogel NC inks were produced for 3D-printed scaffold creation with better processability and biocompatibility. The hydrogel is made of photocrosslinkable alginate that has been bioconjugated with gelatin and CS to

FIGURE 16.4 Schematic representation of the formation of glycosaminoglycan NPs (GAGNPs) and the formulation of GLgel. Reprinted from Zandi N, Sani ES, Mostafavi E, Ibrahim DM, Saleh B, Shokrgozar MA, et al. Nanoengineered shear-thinning and bioprintable hydrogel as a versatile platform for biomedical applications. Biomaterials 2021;267:120476, 120476, Copyright (2021), with permission from Elsevier.

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resemble cartilage ECM, and the nanofiller is made of graphene oxide to improve cell proliferation and printability. The inclusion of graphene oxide into hydrogel inks significantly enhanced the resolution and shape fidelity of threedimensional printed scaffolds, owing to a rapid recovery of viscosity following extrusion of the ink, according to the findings. Furthermore, due to the template formation of the graphene oxide liquid crystal, the NC inks form anisotropic threads after the three-dimensional printing process. Bioconjugated scaffolds have stronger cell proliferation than pure alginate in an in vitro proliferation assay of human adipose tissue-derived mesenchymal stem cells (hADMSCs), with NCs exhibiting the highest values over lengthy periods of time. At Day 7, the Live/Dead assay shows that the hADMSCs attached to the various scaffolds are fully viable. The scaffolds made using NC hydrogel inks had ability to promote cell proliferation in the direction of the three-dimensional printed threads, which was noteworthy. Furthermore, immunostaining after 28 days of culture revealed that the bioconjugate alginate hydrogel matrix triggered chondrogenic differentiation without the use of exogenous pro-chondrogenesis agents. This NC hydrogel inks are a good choice for cartilage tissue engineering based on three-dimensional printing because of their strong cytocompatibility and chondroinductive impact toward hADMSCs, as well as their increased printability and anisotropic architectures [47].

16.4.3 Wound dressing CS-based NC sponges have also been investigated as a potential wound dressing material. For example, a chitosane hyaluronan composite sponge with CS nanoparticle (nCS) was produced. Simple ionic crosslinking with EDC was used to make the hydrogel, which was then lyophilized to make the composite sponge (Fig. 16.5). The size range of the nCS suspension was 100e150 nm. The manufactured sponges were further tested for platelet activation, blood clotting, biodegradation, swelling, and porosity. NCs with a porosity of 67% exhibited improved blood clotting and swelling properties. Human dermal fibroblast (HDF) cells were used to test the sponges’ cell adhesion and cytocompatibility, and the NC sponges demonstrated more than 90% survivability. Within two days of the investigation, NC sponges demonstrated increased HDF cell growth. These findings suggested that NC sponges could be a viable wound dressing option [48].

FIGURE 16.5 Synthesis of chitosan/HYA-nCS composite sponge. Reprinted from Anisha BS, Sankar D, Mohandas A, Chennazhi KP, Nair SV, Jayakumar R. Chitosanehyaluronan/nano chondroitin sulfate ternary composite sponges for medical use. Carbohydr Polym 2013;92:1470e76, Copyright (2013), with permission from Elsevier.

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16.4.4 Other applications CS-based NC hydrogel also finds promising employment in several other areas including food industry and other biomedical applications. For example, to aid in the stabilization of blueberry anthocyanins (ATCs), a unique nanocomplexembedded hydrogel system was developed. Copigmentation of ATCs and CS was used in the method, which was then incorporated into a kappa-carrageenan (KC) hydrogel (as shown in Fig. 16.6). Because of the high charge density of CS, the CS-ATCs nanocomplex exhibited a uniform and more regular structure. The impact of ATCs on nanocomplex stabilization varied depending on storage conditions like metal ions, pH, and temperature. When ATCs were exposed to low pH, the nanocomplex effectively protected them against degradation. This research could help the blueberry beverage sector enable deep processing and render a substantial contribution to the use of blueberry ATCs as a beneficial resource [49]. In order to enable the DielseAlder “click” reaction with a furan-modified pigskin gelatin, titanium dioxide (TiO2) NPs with clickable functional groups were produced. For TiO2 functionalization with the maleimide group, a bifunctional dopamine-maleimide linker was used. Functional NPs were then employed as crosslinkers for gelatin HGs, together with CS. The NC HGs’ swelling and rheological characteristics verified the formation of covalently connected heterogeneous networks. When maleimide-coated NPs were used, the storage moduli values increased. Subsequently, the network’s swelling was substantially diminished, demonstrating that more crosslinked networks were forming. The DielseAlder cycloaddition confirmed the contribution of the surface-attached maleimide functional moieties. Electrostatic force microscopy revealed that HGs responded to electrostatic forces [50]. Nanostructured iron (III) compounds have a high degree of solubility, bioavailability, and redox inertia, making them suitable candidates for iron fortification applications. Assuring the availability of iron in polarized human intestinal epithelial (Caco-2) cells, Feng et al. [51] formulated ferric oxyhydroxide NPs (FeONPs) composed of CS and its NCs with CS or protamine sulfate (PS) in neutral aqueous solution at ambient conditions. The schematic representation of the synthesis of FeONPs is demonstrated in Fig. 16.7. In polarized Caco-2 cells, the calcein fluorescence-quenching assay revealed strong iron uptake from CS-FeONPs, PS/CSFeONPs, and CS/CS-FeONPs, primarily by endocytosis, with greater iron absorption from PS/CSFeONPs and CS/CS-FeONPs. Another investigation used a novel technique to make bioNC HGs by combining a furan-modified gelatin with CS and surface-modified AgNPs as multifunctional crosslinkers. In combination with further amide coupling, DielseAlder cycloaddition of benzotriazole maleimide (BTM) functionalized FIGURE 16.6 Schematic representation of the potential crosslinking in KC (A), free-ATCs-KC (B), and the CS-ATCs-KC (C), respectively. Reprinted from Xie C, Wang Q, Ying R, Wang Y, Wang Z, Huang M. Binding a chondroitin sulfate-based nanocomplex with kappa-carrageenan to enhance the stability of anthocyanins. Food Hydrocoll 2020;100:105448, Copyright (2020), with permission from Elsevier.

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FIGURE 16.7 Schematic representation of the preparation of FeONPs by chondoitin sulfate. Reprinted from Feng G, Tang S, Gang Y, Zeng M, Wu H. Chondroitin sulfate and its nanocomposites with protamine or chitosan stabilize and deliver available nanosized iron. Int J Biol Macromol 2020;150:501e08, Copyright (2020), with permission from Elsevier.

Ag NPs and furan containing gelatin resulted in a stable and biocompatible hybrid NC. The hydrogel’s storage moduli are roughly higher (wthree fold) than those of a blank hydrogel devoid of NPs, showing that the covalently attached NPs have a stabilizing role. Furthermore, the materials’ swelling and drug release capabilities, as well as toxicity and biocompatibility testing, indicate that this type of material has biomedical significance [52].

16.5 Future directions and conclusion HGs are transparent, soft materials consisting primarily of water that can have both solid and liquid properties. The constraints in characteristics and applications that had previously been deemed difficult to overcome have been removed due to the recent advancement in NC HGs. CS-based NC HGs are novel delivery vehicles that combine the properties of NPs with those of a soft matter gel. As a result, a variety of studies into the synthetic, biological, and physicochemical aspects of NC HGs have been conducted. Despite the high level of interest in generating NC HGs with specific functions, we must acknowledge that NC hydrogel development is yet in the early stages. For their final uses in important sectors like as tissue engineering and drug administration, various fine-tuning is required. Methods for designing various NC HGs, for example, are always being developed, and the approach is still working to find the best way to build NC HGs with more specificity. Furthermore, controlling the stimuli responsiveness, biodegradation, and cytotoxicity of NC HGs, that has to be resolved in the future, is still a difficulty. As a result, the introduction of innovative NCs with smart functionalization procedures may be the greatest way to accomplish future progress. Furthermore, industrial-scale NC manufacture remains a challenge. NC HGs with specified structures are now possible owing to advancements in processing technology such as 3D printing and microfluidic reactors. As a result, we anticipate the development of newer NC HGs with organized structures and innovative characteristics for enhanced biological applications.

References [1] Chai Q, Jiao Y, Yu X. Hydrogels for biomedical applications: their characteristics and the mechanisms behind them. Gels 2017;3:6. [2] Ahmed EM. Hydrogel: preparation, characterization, and applications: a review. J Adv Res 2015;6:105e21. [3] Mishra B, Upadhyay M, Reddy Adena SK, Vasant BG, Muthu MS. Hydrogels: an introduction to a controlled drug delivery device, synthesis and application in drug delivery and tissue engineering. Austin J Biomed Eng 2017;4:1037e49. [4] Ho TC, Chang CC, Chan HP, Chung TW, Shu CW, Chuang KP, et al. Hydrogels: properties and applications in biomedicine. Molecules 2022;27:2902. [5] Gu D, O’Connor AJ, GH Qiao G, Ladewig K. Hydrogels with smart systems for delivery of hydrophobic drugs. Expet Opin Drug Deliv 2017;14:879e95. [6] Utech S, Boccaccini AR. A review of hydrogel-based composites for biomedical applications: enhancement of hydrogel properties by addition of rigid inorganic fillers. J Mater Sci 2016;51:271e310. [7] de Oliveira AD, Beatrice CA. Polymer nanocomposites with different types of nanofiller. In: Nanocomposites-recent evolutions; 2018. p. 103e4. [8] Ghasemiyeh P, Mohammadi-Samani S. Hydrogels as drug delivery systems; pros and cons. Trends Pharmacol Sci 2019;5:7e24. [9] Larrañeta E, Stewart S, Ervine M, Al-Kasasbeh R, Donnelly RF. Hydrogels for hydrophobic drug delivery. Classification, synthesis and applications. J Funct Biomater 2018;9:13. [10] Yang Q, Peng J, Xiao H, Xu X, Qian Z. Polysaccharide hydrogels: functionalization, construction and served as scaffold for tissue engineering. Carbohydr Polym 2022;278:118952.

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[11] Zhu T, Mao J, Cheng Y, Liu H, Lv L, Ge M, et al. Recent progress of polysaccharide-based hydrogel interfaces for wound healing and tissue engineering. Adv Mater Interfac 2019;6:1900761. [12] Mohammed AS, Naveed M, Jost N. Polysaccharides; classification, chemical properties, and future perspective applications in fields of pharmacology and biological medicine (A review of current applications and upcoming potentialities). J Polym Environ 2021;29:2359e71. [13] Deng H, Yu Z, Chen S, Fei L, Sha Q, Zhou N, et al. Facile and eco-friendly fabrication of polysaccharides-based nanocomposite hydrogel for photothermal treatment of wound infection. Carbohydr Polym 2020;230:115565. [14] Liu Y, Zhu M, Meng M, Wang Q, Wang Y, Lei Y, et al. A dual-responsive hyaluronic acid nanocomposite hydrogel drug delivery system for overcoming multiple drug resistance. Chin Chem Lett 2022 (in press). [15] Peers S, Montembault A, Ladavière C. Chitosan hydrogels incorporating colloids for sustained drug delivery. Carbohydr Polym 2022;275:118689. [16] Zheng D, Ramos-Sebastian A, Jung WS, Kim SH. Fabrication and preliminary evaluation of alginate hydrogel-based magnetic springs with actively targeted heating and drug release mechanisms for cancer therapy. Compos B Eng 2022;230:109551. [17] Alinejad Y, Adoungotchodo A, Hui E, Zehtabi F, Lerouge S. An injectable chitosan/chondroitin sulfate hydrogel with tunable mechanical properties for cell therapy/tissue engineering. Int J Biol Macromol 2018;113:132e41. [18] Sharma R, Kuche K, Thakor P, Bhavana V, Srivastava S, Mehra NK, et al. Chondroitin Sulfate: emerging biomaterial for biopharmaceutical purpose and tissue engineering. Carbohydr Polym 2022;286:119305. [19] Kwon HJ, Han Y. Chondroitin sulfate-based biomaterials for tissue engineering. Turk J Biol 2016;40:290e9. [20] Zhu W, Ji Y, Wang Y, He D, Yan Y, Su N, et al. Structural characterization and in vitro antioxidant activities of chondroitin sulfate purified from Andrias davidianus cartilage. Carbohydr Polym 2018;196:398e404. [21] Volpi N. Chondroitin sulfate safety and quality. Molecules 2019;24:1447. [22] Krichen F, Bougatef H, Capitani F, Amor IB, Koubaa I, Gargouri J, et al. Purification and structural elucidation of chondroitin sulfate/dermatan sulfate from Atlantic bluefin tuna (Thunnus thynnus) skins and their anticoagulant and ACE inhibitory activities. RSC Adv 2018;8:37965e75. [23] Krichen F, Bougatef H, Sayari N, Capitani F, Amor IB, Koubaa I, et al. Isolation, purification and structural characterestics of chondroitin sulfate from smooth hound cartilage: in vitro anticoagulant and antiproliferative properties. Carbohydr Polym 2018;197:451e9. [24] Ustyuzhanina NE, Bilan MI, Dmitrenok AS, Nifantiev NE, Usov AI. Fucosylated chondroitin sulfates from the sea cucumbers Holothuria tubulosa and Holothuria stellati. Carbohydr Polym 2018;200:1e5. [25] Gross AR, Theoharides TC. Chondroitin sulfate inhibits secretion of TNF and CXCL8 from human mast cells stimulated by IL-33. Biofactors 2019;45:49e61. [26] Sun Y, Zhang G, Liu Q, Liu X, Wang L, Wang J, et al. Chondroitin sulfate from sturgeon bone ameliorates pain of osteoarthritis induced by monosodium iodoacetate in rats. Int J Biol Macromol 2018;117:95e101. [27] Yang L, Wang Y, Yang S, Lv Z. Separation, purification, structures and anticoagulant activities of fucosylated chondroitin sulfates from Holothuria scabra. Int J Biol Macromol 2018;108:710e8. [28] Yukawa T, Iwakura Y, Takei N, Saito M, Watanabe Y, Toyooka K, et al. Pathological alterations of chondroitin sulfate moiety in postmortem hippocampus of patients with schizophrenia. Psychiatr Res 2018;270:940e6. [29] Takeda A, Okada S, Funakoshi K. Chondroitin sulfates do not impede axonal regeneration in goldfish spinal cord. Brain Res 2017;1673:23e9. [30] Gao W, Zhang Y, Zhang Q, Zhang L. Nanoparticle-hydrogel: a hybrid biomaterial system for localized drug delivery. Ann Biomed Eng 2016;44:2049e61. [31] Gomez-Florit M, Pardo A, Domingues RM, Graça AL, Babo PS, Reis RL, et al. Natural-based hydrogels for tissue engineering applications. Molecules 2020;25:5858. [32] Alcântara MT, Lincopan N, Santos PM, Ramirez PA, Brant AJ, Riella HG, et al. Simultaneous hydrogel crosslinking and silver nanoparticle formation by using ionizing radiation to obtain antimicrobial hydrogels. Radiat Phys Chem 2020;169:108777. [33] Bi X, Liang A. In situ-forming cross-linking hydrogel systems: chemistry and biomedical applications. In: Emerging concepts in analysis and applications of hydrogels. 86; 2016. p. 541e7. [34] Liang S, Zhang Y, Wang H, Xu Z, Chen J, Bao R, et al. Paintable and rapidly bondable conductive hydrogels as therapeutic cardiac patches. Adv Mater 2018;30:1704235. [35] Kuang L, Ma X, Ma Y, Yao Y, Tariq M, Yuan Y, et al. Self-assembled injectable nanocomposite hydrogels coordinated by in situ generated CaP nanoparticles for bone regeneration. ACS Appl Mater Interfaces 2019;11:17234e46. [36] Thoniyot P, Tan MJ, Karim AA, Young DJ, Loh XJ. Nanoparticleehydrogel composites: concept, design, and applications of these promising, multi-functional materials. Adv Sci 2015;2:1400010. [37] Thomas PC, Cipriano BH, Raghavan SR. Nanoparticle-crosslinked hydrogels as a class of efficient materials for separation and ion exchange. Soft Matter 2011;7:8192e7. [38] Gaspar VM, Lavrador P, Borges J, Oliveira MB, Mano JF. Advanced bottom-up engineering of living architectures. Adv Mater 2020;32:1903975. [39] Lavrador P, Esteves MR, Gaspar VM, Mano JF. Stimuli-responsive nanocomposite hydrogels for biomedical applications. Adv Funct Mater 2021;31:2005941. [40] Simão AR, Fragal VH, Pellá MC, Garcia FP, Nakamura CV, Silva R, et al. Drug polarity effect over the controlled release in casein and chondroitin sulfate-based hydrogels. Int J Biol Macromol 2020;158:116e26. [41] Zhang H, Xu J, Xing L, Ji J, Yu A, Zhai G. Self-assembled micelles based on Chondroitin sulfate/poly (d, l-lactideco-glycolide) block copolymers for doxorubicin delivery. J Colloid Interface Sci 2017;492:101e11.

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[42] de Olyveira GM, Costa LM, Basmaji P, de Cerqueira Daltro G, Guastaldi AC. Hydrogel bacterial cellulose behavior with stem cells. Adv Sci Eng Med 2015;7:393e7. [43] Simão AR, Fragal VH, de Oliveira Lima AM, Pellá MC, Garcia FP, Nakamura CV, et al. pH-responsive hybrid hydrogels: chondroitin sulfate/casein trapped silica nanospheres for controlled drug release. Int J Biol Macromol 2020;148:302e15. [44] Yin, et al. Preparation and characterization of sodium alginate/chitosan composite nanoparticles loaded wit hchondroitin sulfate. 2021. [45] Zhou L, Fan L, Zhang FM, Jiang Y, Cai M, Dai C, et al. Hybrid gelatin/oxidized chondroitin sulfate hydrogels incorporating bioactive glass nanoparticles with enhanced mechanical properties, mineralization, and osteogenic differentiation. Bioact Mater 2021;6:890e904. [46] Zandi N, Sani ES, Mostafavi E, Ibrahim DM, Saleh B, Shokrgozar MA, et al. Nanoengineered shear-thinning and bioprintable hydrogel as a versatile platform for biomedical applications. Biomaterials 2021;267:120476. [47] Olate-Moya F, Arens L, Wilhelm M, Mateos-Timoneda MA, Engel E, Palza H. Chondroinductive alginate-based hydrogels having graphene oxide for 3D printed scaffold fabrication. ACS Appl Mater Interfaces 2020;12:4343e57. [48] Anisha BS, Sankar D, Mohandas A, Chennazhi KP, Nair SV, Jayakumar R. Chitosanehyaluronan/nano chondroitin sulfate ternary composite sponges for medical use. Carbohydr Polym 2013;92:1470e6. [49] Xie C, Wang Q, Ying R, Wang Y, Wang Z, Huang M. Binding a chondroitin sulfate-based nanocomplex with kappa-carrageenan to enhance the stability of anthocyanins. Food Hydrocoll 2020;100:105448. [50] García-Astrain C, Miljevic M, Ahmed I, Martin L, Eceiza A, Fruk L, et al. Designing hydrogel nanocomposites using TiO2 as clickable crosslinkers. J Mater Sci 2016;51:5073e81. [51] Feng G, Tang S, Gang Y, Zeng M, Wu H. Chondroitin sulfate and its nanocomposites with protamine or chitosan stabilize and deliver available nanosized iron. Int J Biol Macromol 2020;150:501e8. [52] García-Astrain C, Chen C, Burón M, Palomares T, Eceiza A, Fruk L, et al. Biocompatible hydrogel nanocomposite with covalently embedded silver nanoparticles. Biomacromolecules 2015;16:1301e10.

Further reading [1] Wahid F, Zhao XJ, Jia SR, Bai H, Zhong C. Nanocomposite hydrogels as multifunctional systems for biomedical applications: current state and perspectives. Compos B Eng 2020;200:108208.

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Chapter 17

In situ gel based on gellan gum Jieyu Zhu1, Yijun Pan1, Haizhou Peng2, Jinzhang Fang2, Guoxin Du2, Akshaya Tatke1 and Bo Liang1, 2 1

iVIEW Therapeutics Inc., Cranbury, NJ, United States; 2iVIEW Therapeutics (Zhuhai) Co. Ltd., Zhuhai, Guangdong, China

17.1 Introduction As a novel delivery system, in situ gels will undergo sol-to-gel transition when contacting with body fluids [1]. This drug delivery system has become a new trend in the design of pharmaceutical formulations, because its sustained and prolonged drug release characteristics help to increase the bioadhesion and bioavailability of drugs and improve the compliance and comfort of patients during use [2]. Various polymers, including natural polymers (e.g., gellan gum [GG], chitosan, pectin, xanthan gum, etc.) [3e6] and synthetic polymers (e.g., poloxamer, carbopol, aliphatic polyesters, etc.) [2,7,8], can have gel-forming properties and can be used in different pharmaceutical formulations according to the delivery route. The formation of in situ gel is triggered by physical or chemical mechanisms. Physical mechanisms include swelling [9], diffusion [10], or other temperature-triggered gelation [11,12]. Chemical mechanisms consist of ionic cross-linking [13,14], pH changes-triggered cross-linking [15,16], enzymatic cross-linking, and photopolymerization [1,17]. GG is an anionic polysaccharide produced by bacteria Sphingomonas elodea (formerly produced by Pseudomonas elodea as extracellular polysaccharides), which has a repeating unit consisting of two b-D- glucose, one L-rhamnose, and one Dglucuronic acid [18,19]. Gellan gum is classified into low acyl (also called deacylated) and high acyl gellan gum based on the amount of acetate groups attached to its backbone. Two kinds of gellan gum can both be used as gelling agents. Rigid, nonelastic, and brittle gels are formed by low acyl gellan gum, while soft, elastic, and nonbrittle gels are formed by high acyl gellan gum [20]. The desired texture can be obtained by adjusting the ratio of these two gellan gums. Low acyl gellan gum products can be commercially obtained from CP Kelco. KELCOG Gellan Gum (food grade or household grade) is widely used as a gelling agent, which forms gels when it contacts with monovalent, divalent, and multivalent ions. USP grade LA gellan gum (Gelrite) is also commercially available and has been used in approved drugs (for example, TimopticXE, 0.25% and 0.5% timolol maleate ophthalmic gel-forming solution from Bausch & Lomb Americas Inc.). Recently, iVIEW-1201 1.0% povidone iodine ophthalmic gel-forming sterile solution [21] was developed by iVIEW Therapeutics Inc. in New Jersey, USA, using gellan gum as in situ gel-forming agent for the treatment of viral and bacterial conjunctivitis, which is now in global Phase II clinical trials.

17.2 Mechanisms of gellan gum-based in situ gels Gellan gum exists in two forms: one is high acyl gellan gum (HA gellan gum, also called natural gellan gum), and the other is low acyl gellan gum (LA gellan gum, as also called deacylated gellan gum). On average, each repeating unit of natural gellan gum has one acyl group connected to the glucose molecule. Due to the existence of the acyl group, the gel formed is soft, elastic, and has strong adhesion. After heat treatment with alkali (pH > 10), the molecular acyl groups in the structure were partially or completely removed, which significantly weakens the molecular steric hindrance and enhances the gelforming ability. Deacetylated gellan gum (Gelrite or Kelcogel) is commercially available and commonly applied in food, cosmetic, and pharmaceutical industries [22]. Gellan gum is available in an odorless dry powder form with beige color. Gellan gum powder has good stability against heat, acid, and enzyme. Gellan gum is not soluble in cold water nor nonpolar organic solvents. It can be dispersed in hot water and dissolved during stirring after cooling, and it will form a transparent solid gel. In the presence of integrating agent, it can be dissolved in hot deionized water or solution with low ionic strength. The pH of aqueous solution is neutral.

Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine. https://doi.org/10.1016/B978-0-323-95351-1.00009-0 Copyright © 2024 Elsevier Inc. All rights reserved.

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As a new generation of gelling agent, gellan gum has the unique advantages such as low dosage and good gelling property. Gellan gum is both pH-sensitive and ion-sensitive. It is an anionic polysaccharide that shows in situ gel-forming property with the presence of mono (Naþ, Kþ) or divalent cations (Ca2þ, Mg2þ). The gelling concentration of gellan gum can be even lower if the metal cations in the solution reach a certain amount. With the addition of 10 mM 0.04 g/L CaCl2, only 0.02% gellan gum is needed to form a structured network with solid-like behavior [23]. Compared with traditional gelling polysaccharides (such as xanthan gum, carrageenan, agar, etc.), much lower concentration (1%) [24]. Currently, gellan gum has gradually replaced traditional gelling agents (agar and carrageenan) in industry. The gelling formation of gellan gum requires the presence of salt ions, and the gelling strength is affected by the type and concentration of salt ions. Although monovalent cations (Naþ, Kþ) and divalent cations (Ca2þ, Mg2þ) can both initiate gelation of gellan gum, but the influence of divalent cations is much greater. The ion concentration required for monovalent cations is about 25 times higher than that of calcium ions to form a gel at the same gellan gum concentration [25]. The maximum gel hardness and modulus are also generated at very low divalent cation concentration. This is because the double helix hydrogen bonds formed by Ca2þ and Kþ ions are different. Ca2þ forms a carboxylate-Ca2þ-carboxylate structure, while Kþ forms a carboxylate-Kþ-H2O-Kþ-carboxylate structure [26].

17.3 Compatibility of gellan gum in different formulations Gellan gum can be easily combined with other gums/polymers. LA gellan gum, in particular, is compatible at neutral pH with milk proteins, soy, and egg albumen [27]. The usage of gellan gum requires an understanding of its compatibility with different drug delivery platforms. Gellan gum has been widely used in numerous formulations to increase viscosity or impart in situ functionality. Gellan gum is currently used in many kinds of dosage forms, including gels [28], tablets [29], beads [30], films [31], microspheres [32] and microcapsules [33], hydrogels [34], nanoparticles [35], etc. This section discusses the various applications of gellan gum in aqueous and lipid-based formulations.

17.3.1 Gellan gum used in aqueous-based formulations Gellan gum was one of the most widely used polysaccharides in the in situ gel technologies. Researchers had prepared an in situ gel formulation based on gellan gum, polyvinyl alcohol, and chitosan for continuous intraocular delivery of besifloxacin. The results showed that the formulation was nonirritating and showed long-term antibacterial efficacy. Among the currently used polymer combinations, gellan gum exhibits unique ion-induced gelatinization, which may also result in better gel strength and quick response [36]. In suspension formulation system, local suspension of 3% 2-methyl yam alcohol (2-MS) was prepared with gellan gum (0.5% W/V). The results showed that no cortical vacuole formation occurred after 21 days of galactose feeding in rats treated with topical suspension containing gellan gum. Gellan gum can increase the viscosity of the suspension to increase the bioadhesion on the cornea and to increase the bioavailability of the 2-MS [37]. A transmucosal ophthalmic nanosuspension for Posaconazole delivery was developed by Khare et al. Gellan gum was able to prolong the localization time of the drug at the corneal site for up to 24 h [38]. An aqueous suspension of gellan gum-based microcapsules was used to deliver hydrophobic molecules, in which gellan gum serves as a stabilizer and a gelling agent. Because of its thermoresponsive nature, it forms a physical cross-link hydrogel in the presence of cations and is stable and less permeable when compared to alginate-based hydrogels. This work opens a wide range application for the water-based microcapsules coupled with hydrogel, to incorporate various poorly water-soluble compounds and optimizing a manufacturing production method [39]. In order to decrease the mucociliary clearance and facilitate drug uptakes, budesonide, another poorly watersoluble drug, was incorporated together with povidone iodine in gellan gum-based in situ gels, to treat chronic rhinosinusitis (CRS) [40]. Gellan gum has been also used in combination with cyclodextrins as an in situ gelling agent to deliver poor watersoluble drugs via ocular delivery. Senjoti et al. incorporated Flurbiprofen (FBP), a poorly soluble drug into cyclodextrin complexes and further combined the system with in situ gelling system. This system greatly improved the gelation with a combination effect and also the drug permeation through ocular tissues [41]. In another recent study, interpenetrated cross-linked polymer network-based microbeads were prepared by combining gellan gum with carboxymethyl tamarind gum, which also displayed a more sustained release profile of diclofenac sodium [42]. As a natural source and biocompatible polymer, gellan gum-based DTX nanovesicle hydrogel showed instant in vitro in situ gelling effect, good injectability, lower toxicity, and a better tumor reduction compared to the conventional commercial DTX injectable formulation [43]. Aqueous-based gellan gum nanoparticles were used to deliver Tacrilomus

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for dry eye disease. The drug release from gel nanoparticles was prolonged throughout the 12-hour study, and precorneal retention was higher compared with tacrolimus solution [44]. Besides nanoparticles, gellan gum is also compatible with nanomicelles. It was reported that curcumin-loaded micelles were mixed with gellan gum at the ratio of 3:1 and 1:1 to form a transparent curcumin-mixed micelle, which had a weak irritation toward ocular tissue. This mixed formulation of micelles and in situ gels with gellan gum as the main excipient can increase the application of curcumin as a low-solubility drug in ophthalmology [45]. Similarly, in situ gel-incorporated nanomicelles were also used to deliver other poorly water-soluble drugs, such as cyclosporine to treat dry eye disease [46] and difluprednate to treat postsurgery ocular inflammation [47].

17.3.2 Gellan gum used in lipid-based formulations The combination of nanoemulsion and in situ gel technology has obvious sustained release effect compared with the nanoemulsion alone. Acetazolamide based on HPMC/gellan gum and xanthan gum/gellan gum, respectively, has higher efficacy and longer antihypertensive time than the commercially available acetazolamide eye drops and oral tablets. Gellan gum showed optimum gelation effect at a concentration of 0.3% when mixed with artificial tears [48]. In a recent study, gellan gum in combination with carbopol was used to prepare teriflunomide-loaded nanolipid formulation in situ gel to examine its effect on the brain and spine tumors, compared to teriflunomide-lipid nanoparticles. The mucoadhesive function of gellan gum in the teriflunomide NLC gel prolonged the drug retention time in the nasal cavity, thereby showing an enhanced permeation [49]. There has also been a significant increase in the delivery of drugs to the brain via the nasal route. Fahmy et al. developed flibanserin-NLC combined with gellan gum as an in situ gel, which helped improve drug bioavailability to the brain [50]. Likewise, repaglinide-loaded solid lipid nanoparticle (SLN) formulation delivered through the nasal route has been developed by Elkarray et al. In comparison with the oral route, nasal administration showed 2.5 times decrease in the blood glucose level over the extended period of 48 h, thus, reducing the dosing frequency for repaglinide [51]. Similar to ocular drug delivery, gellan gum alters nasal drops gel due to the presence of cations present in nasal fluid which increases the formulation residence time and drug efficacy. A composite wound dressing of lipid nanoparticles containing gellan gum and alginate used for antimicrobial peptidelactostreptococcin was developed. The results showed that the mixture of lipid nanoparticles and gellan gum significantly slowed the release of Nisin compared with gellan gum alone. Gellan gum prolonged the antibacterial time and improved the antibacterial effect of Nisin [52]. Tatke et al. and Youssef et al. formulated gellan gum-based SLNs and nanolipid carriers (NLCs) for steroidal and antifungal ocular delivery, respectively. The transcorneal in vitro permeation of the drug molecules from the lipid nanoparticle in situ gel formulations (SLN and NLC) showed sustained release effect when comparing to nongel-based ones. Moreover, an in vivo ocular distribution performed by Tatke et al. concluded that in situ formulations delivered higher amounts of triamcinolone acetonide to the posterior ocular tissues in rabbits [53,54]. The existence of the gellan gum alters the ocular formulation solution to in situ gel due to the presence of cations present in the tear fluid, thus increasing the viscosity of the formulation and increasing the residence time and drug efficacy. The researchers used the thin-film hydration method to prepare liposomes containing gellan gum, trehalose, and borate. The results showed that the liposome-based in situ gels exhibited acceptable irritability, good resistance from tear drainage, and extended ocular retention time, which facilitates its use in topical ophthalmic products [55].

17.4 Applications of gellan gum as in situ gels The safety and nontoxicity of gellan gum are the prerequisites of its pharmaceutical and medical applications. The cationsensitive feature makes it possible to prepare in situ gel, improve the bioavailability, or achieve the purpose of local treatment, and can be used as the carrier of sustained and controlled release formulations to achieve constant or localized release [18,56]. Gellan gum in situ gel is applied to plentiful Active Pharmaceutical Ingredients (APIs) as therapeutic treatments for different indications. In these applications, gellan gum and APIs are formulated into in situ gelling delivery system and administrated by routes of ocular, intranasal, peroral, injectable, virginal, topical, and others. Via some of these routes, the use of gellan gum in situ gelling drug delivery system has been thriving in the emerging regenerative medicine [57,58]. By means of structural modification, property transformation or combination with delivery carriers such as micelles, nanoemulsions, nanosuspensions, inclusion complexes, liposomes, lipid nano- or microparticles, and so on [59,60], gellan gum performs excellent characteristics to improve the low drug bioavailability, reduce toxicity, and alleviate side effects. With tremendous research and development, gellan gum will have a broader application prospect in the field of drug discovery and medical treatment. In this book chapter, drug applications on different delivery routes like ophthalmic, intranasal, oral, injectable, transdermal, or topical pathways and other usual approaches are presented here, as summarized in Table 17.1.

TABLE 17.1 Applications of gellan gum in situ gels by various delivery routes. Formulation/delivery system

Application

References

Gatifloxacin

In situ gelling solution

Antibiotic

[86]

Moxifloxacin

In situ gelling solution

Antibiotic

[87]

Levofloxacin

In situ gelling solution of gellan gum and chitosan

Antibiotic

[64]

Ciprofloxacin HCl

In situ gel with gellan gum, modified xanthan gum, and hydroxy propyl methyl cellulose

Antibiotic

[88]

Povidone iodine

In situ gelling solution

Virucidal

[21]

Fluconazole

Gellan gum with cyclodextrin as in situ gelling solution

Antifungal

[89]

Posaconazole

In situ gelling nanosuspension

Antifungal

[38]

Natamycin

In situ gelling solution

Antifungal

[90]

Flurbiprofen axetil

Nanoemulsion in situ gel system

Antiinflammatory

[91]

Intranasal

Curcumin

In situ gelling nanomicellar solution

Antiinflammatory, antimicrobic

[45]

Difluprednate

In situ gelling nanomicellar solution

Antiinflammatory

[47]

Cyclosporine

In situ gelling nanomicellar solution

Antiinflammatory

[46]

Benzododecinium bromide

In situ gelling solution with carbopol 934P

Antimicrobic

[92]

Timolol maleate

Ion- and pH-activated eye drop

Antiglaucoma

[93]

Brimonidine tartrate

In situ gelling solution

Antiglaucoma

[94]

Pilocarpine HCl

Gellan gum in situ gelling solution with poloxamer 407

Antiglaucoma

[95]

Acetazolamide

In situ gelling nanoemulsions

Antiglaucoma

[48]

Connexin 43 antisense oligodeoxynucleotide

In situ gelling solution

Wound repair

[96]

Breviscapine

Nanosuspension in situ gelling system

Antioxidant; antiinflammatory

[97]

Povidone iodine/budesonide

In situ gel nasal spray

Antiinfective/antiinflammatory

[40]

Gastrodin

Solution; in situ gelification

Antiinflammatory

[98]]

Mometasone furoate

In situ nasal gel

Antiinflammatory

[99]

Metoclopramide

Microparticle solution; in situ gelification

Antiemetic

[100]

Dimenhydrinate

In situ nasal gel

Antiemetic

[101]

Granisetron HCl

Solution; in situ nasal droppable gel

Antiemetic

[102]

Lamotrigine

In situ nasal gel

Antiepileptic

[66]

Sumatriptan succinate

In situ nasal gel

Antimigraine

[68]

Lorazepam

In situ nasal microemulsion gel

Anxiolytic

[103]

Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine

API

Ophthalmic

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Delivery route

Antidepressant

[50]

Carvedilol

In situ nasal nanosuspension gel

Beta-blocker

[104]

Salbutamol sulfate

Solution, mucoadhesion in situ gelling

Bronchodilator

[105]

Rivastigmine

In situ nasal nanosuspension gel

Cholinesterase inhibitor

[106]

Rizatriptan benzoate

In situ gelling system

5-HT1 agonist

[107]

Povidone iodine

Solution; buccal gelification in situ

Antiseptic

[108]

Cefpodoxime proxetill

Solution; gastrointestinal tract in situ

Antibiotic

[109]

Clarithromycin

Floating in situ gelling system

Antibiotic

[72]

Amoxicillin

Floating in situ gelling system

Antibiotic

[74]

Ciclopriox olamine

Solution; buccal gelification in situ

Antifungal

[110]

Clotrimazole

Solution; gelification in situ

Antifungal

[111]

Fluconazole

Solution; oropharingeal gelification in situ

Antifungal

[112]

Ibuprofen

In situ gelling system

Analgesic and antiinflammatory

[113]

Carbamazepine

Liquid suppository; mucoadhesive rectal in situ gel

Antiepileptic

[75]

Ambroxol

Solution; gastrointestinal tract in situ gel

Antineural

[114]

Itopride HCl

Floating in situ gelling raft system

Antiemetic

[115]

Acetohydroxamic acid

Floating in situ gelling system

Antiinfectious

[116]

Paclitaxel

Urothelial liposome-in gel system; intravesical injection

Chemotherapy

[117]

Ornidazole

Solution; smart periodontal gel

Antibiotic

[118]

Vancomycin

Nanoparticle gel for skeletal injection

Antibiotic

[119]

Recombinant human bone morphogenetic protein-2

Liposomal in situ gel graft

Repair of alveolar bone clefts

[120]

3D hydroxyapatite scaffold model

Microparticle hydrogel

Build pore architecture 3D hydroxyapatite scaffolds

[121]

Acellular or cellular substitutes of the nucleus pulposus

Gellan gum-based microparticle hydrogel; intervertebral injection

Strategies for intervertebral disc degeneration

[122]

Cell delivery model

Implant/injection hydrogel scaffolds by low acyl and low/ high acyl gellan gum blends

Substitutions in musculoskeletal tissues

[123]

Secnidazole

Solution; in situ gel

Antibiotic

[124]

Clindamycin HCL

Solution; in situ gel

Antibiotic

[125]

Topical

Clotrimazole

In situ oral topical gel; oral candidiasis treatment

Antifungal

[111]

Intracanal

Chlorhexidine

Solution; in situ pH-sensitive solegel

Antibiotic

[126]

Oral

Injectable

Virginal

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In situ gel with nanostructured lipid carriers

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Flibanserin

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17.4.1 Ophthalmic in situ gelling drug delivery Topical installation is a desirable ophthalmic route for drugs due to the easy manufacturing process, convenient use, accurate delivery dosage, and high patient compliance. For decades, the widely used conventional eye drops still have poor intraocular bioavailability and need more frequent or increased drug dosage because of short action time, high dilution rates, and fast elimination induced by continuous lacrimal flow, blinking, and limited corneal permeability [61]. To figure out solutions to the problems, multiple advanced ophthalmic drug delivery systems were developed, such as nanoparticles, liposomes, implants, or in situ gels, among which, in situ gelling systems are desired to design formulations with prolonged drug release and higher loaded APIs [57,62,63]. In series of natural and synthetic polymers that can be used to form in situ gels for ophthalmic drugs, gellan gum gains much attention for its good tolerance, transparency, gelling properties, and quite low toxicity. In cation environment, in situ gelling gellan gum not only can form strong and clear gels from a drop of solution but also show no irritation to the eye as well as high patient compliance [18,64]. Once rapidly gelled, gellan gum in situ gels show resistance to the natural drainage process from the precorneal area, thus increasing the ocular retention time, reducing the drug adverse reaction, and improving bioabsorption of drugs. With these advantages, gellan gum-based in situ gelling ophthalmic delivery systems are primarily investigated for antibiotic, antiinflammatory, antiglaucoma, and some other drugs for ocular treatment (Table 17.1). Moreover, gellan gum and some polymeric materials are combined to act as ocular in situ gelling delivery systems in order to enhance their viscosity, hardness, and pseudoplasticity upon addition of monovalent and divalent cations, consequently reducing the influence of nasolacrimal drainage (Table 17.1). Topical ophthalmic formulations demonstrate large numbers of practical cases of commercially available gellan gum in drug delivery applications [65].

17.4.2 Intranasal in situ gelling drug delivery The human nasal mucosa is filled with w0.1 mL of tissue fluid that consists of multiple ions like sodium, potassium, calcium, chloride, and phosphate. Once the nasal fluid is fully in contact with gellan gum solution, it immediately forms an in situ gel, in which the drug can be retained with longer time and slower release. Thereby, gellan gum is particularly suitable for drugs requiring long-term intranasal administration [66]. Besides overcoming the drawbacks of short retention time and low bioavailability, in situ nasal gelling formulation with gellan gum overcomes the disadvantages of high viscosity, inconvenient use, and inaccurate dosage of some nasal bioadhesives [67e69]. Relying on the responsiveness to cations present in physiological conditions, gellan gum is applied for massive antiinflammatory, antiemetic, antimigraine, and antiepileptic drugs as in situ nasal gelling forms (Table 17.1). In these forms, gelling systems are loaded with nanosuspension, micro- or nanoemulsion, and liposomes to significantly enhance the drug bioavailability and mucoadhesiveness. In addition, the excipients of nasal preparations are generally not appropriate to have strong odor, while gellan gum solution is colorless and tasteless, which is suitable for use as a carrier of intranasal formulations.

17.4.3 Oral in situ gelling drug delivery As a safe, convenient, and preferable route, oral delivery administration achieves considerable efficacy; however, for conventional oral dosed formulations, local or systemic toxicity mainly generated from the short gastric retention time, rapid gastric release, or failure delivery to target sites, which usually leads to efficacy and application limitations of the drug. Novel oral delivery systems designed to be floating, swellable, mucoadhesive, high-density sinking, or colon-specific biodegradable can avoid drug damage from the upper and lower gastrointestinal tract (GIT) environment [70,71]. With peculiarities of solution viscosity, biocompatibility, and fast solution to gelation (solegel) transition in the presence of environmental cations or gastric acidic pH, gellan gum shows a wide range of investigations for in situ gelling oral liquid formulations [72]. Superior to common oral solutions, gellan gum in situ solution is liquid at room temperature and can transit into an in situ gel after oral administration with the help of structural modification or combination with polymeric or responsive materials, thus showing a sustained drug release and enhanced in vivo bioavailability [73e76]. A large proportion of examples displayed that gellan gum formulated antifungal, antibiotic, analgesic, and antiemetic drugs into gastrointestinal in situ gelling liquids, floating in situ gelling raft, and buccal gelification in situ solution (Table 17.1). Taken together, gellan gum is of great interest and particularly suited for such applications in oral fluids as it can freely load drugs with various properties in liquid state to produce better in vivo efficacy.

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17.4.4 Injectable in situ gelling drug delivery Formulated as minimally invasive injectable delivery systems, gellan gum generally transforms from liquid into in situ gel under physiological conditions [77]. Located at the site, the in situ gels contribute to drug repository and sustained release to acquire enhanced bioavailability. Some cases by intravesical, periodontal, and skeletal injection are shown in Table 17.1. In recent decades, gellan gum has exhibited conspicuous potency in tissue engineering and regenerative medicine attributable to the cationic sensitivity, good biocompatibility, and stability [78]. Via injection with the gellan gum-based suspension of transplanted cells, the in situ gels form supportive carriers to maintain increased survival of the cells and to further promote tissue regeneration for over couples of weeks or months. With high water content, gellan gum-based hydrogel benefits to nutrient delivering and cell signal transmission. In this way, it can promote the secretion of collagen from wounds and accelerate wound repair, producing no inflammatory reaction [79]. In order to supply a more viable injectable in situ gelling substrate for diverse human tissues such as brain, bone, cartilage, liver, and muscle, gellan gum is employed to develop moderate blending composites, physical or structural modification, which can produce more effective drug delivery and biomedical functions [80,81]. Take cartilage tissue engineering for example, gellan gum hydrogel networks with different structures (discs, membranes, fibers, particles, and scaffolds) can be prepared based on temperature and pH-dependent reactions through casting, extrusion, or lyophilization method, and the obtained gellan gum hydrogels displayed good encapsulation capacity of human nasal chondrocytes [80]. The bone tissue regeneration can be achieved by enzymatical mineralization of gellan gum hydrogels in mineralization media containing calcium and/or magnesium glycerophosphate [82]. Particularly for the skin tissue regeneration, gellan gum in situ gelling system has obtained more and more attention in drug delivery for wound healing and tissue repair as listed in Table 17.1. All these cases only give partial versatility of gellan gum-based in situ gels; for further research, it is expected to be explored with more potent and interesting outcomes in regenerative medicine delivery applications [83].

17.4.5 Other in situ gelling drug delivery routes Other dosing routes of in situ gels have been explored for gellan gum to deliver drugs, including intracanal, vaginal, and topical, as demonstrated in Table 17.1. Interestingly, there are some common senses among the reported works that the gellan gum-based in situ gelling delivery systems perform improved mucoadhesiveness, sustained drug release, and less administrated dose compared to conventional formulations [84,85]. Notably, it can be predicted that gellan gum in situ gel will have a broader application prospect in drug delivery or biomaterial production with further academic research and industrial development.

17.5 Conclusion Gellan gum offers a large number of properties such as biocompatibility, nontoxicity, mucoadhesion, thermal stability, good stability, good solubility, good water-holding capacity, high transparency, making it a proper candidate as a stabilizing, thickening, binding, and in situ gelling agent. In the development and use of drugs, gellan gum has been broadly used as a pharmaceutical excipient in the United States and other countries around the world. As an excellent ion-sensitive biopolymer, it has been used to prepare various bioadhesive in situ gels to regenerate biotissues, reduce the drug irritation, prolong the residence time, control release speed, and improve bioavailability and local therapeutic effects. More functional properties are expected to be obtained through structural modification, modification, or compatibility with other excipients. It is foreseeable that with the continuous deepening of research and development, gellan gum will have broader application prospects in the field of food, drug, and cosmetic industries.

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Chapter 18

Preclinical and clinical study of polysaccharide-based hydrogels Bijaya Ghosh1, Moumita Das Kirtania2 and Ranjan Kirtania2 1

NSHM College of Pharmaceutical Technology, NSHM Knowledge Campus, Kolkata Group of Institutions, Kolkata, West Bengal, India; 2School of

Pharmacy, Techno India University, Kolkata, West Bengal, India

18.1 Introduction Hydrogels are basically networks of hydrophilic polymers with capability of holding high amount of water in them [1]. The network structure gives them a rare flexibility comparable to the natural human tissue, and due to the presence of hydrophilic functional groups, they can carry water, without being dissolved in it. This is a highly sought-after property in biomaterials and opened their way of commercial application in a number of biomedical fields. In the last five decades, lots of research has been done on the pharmaceutical and biomedical application of hydrogels and some of them have made significant contribution in improving the quality of human life. Hydrogels can be obtained from both natural and synthetic sources. Hydrogels of synthetic origin allow molecular-scale control during preparation and can have customized physicochemical and biological properties [2]. However, the very nature of synthesis exposes them to initiators and cross-linkers, many of which have substantial toxicity. In this regard, hydrogels obtained from natural sources are considered safe. As a result, interest in the natural hydrogels has increased. In general, synthetic hydrogels score over the natural ones in terms of tensile strength and structural stability, whereas natural ones are better in terms of biocompatibility and biodegradability. Hybrid hydrogels that combine both natural and synthetic polymers can strike a balance to offer the benefits of both types. Development of hydrogel-based drug delivery systems is one of important focuses in current pharmaceutical research. Work is on to develop systems for small molecules [3e5], proteins [6,7], live cells [8], and some difficult to deliver biomolecules like DNA pieces [9]. Some of this research has already found their way in clinics. In the last five decades, the repertoire of drugs has been enriched by the invention of recombinant proteins and peptidebased therapeutic agents which are hydrophilic in nature [10]. Oral delivery of hydrophilic macromolecules is restricted by their size and lack of hydrophobicity. Environmentsensitive hybrid hydrogels have evolved as a good delivery media for these drugs. With greater understanding about the architecture of stem cells and opening of newer avenues of material engineering, repair and regeneration of complex tissues have become a reality [11]. Hydrogels play a prominent role in tissue engineering and regenerative techniques too. In our earlier book, we have devoted a chapter on commercially available hydrogel products used for biomedical purposes derived from plant and algal sources [12]. This chapter discusses the hydrogels that could pass through the stringent filters of safety and efficacy testing and finally reached the market to establish their utility in drug delivery and regenerative medicine with special emphasis on products derived from animal and bacterial sources. The content is framed to answer the following questions: (1). What are the diseases for which hydrogel-based delivery systems are mainly used? (2). What are their conventional modes of treatment? (3). How hydrogel-based delivery systems compare to the existing delivery systems? Some systems undergoing clinical trials are also included.

Polysaccharide Hydrogels for Drug Delivery and Regenerative Medicine. https://doi.org/10.1016/B978-0-323-95351-1.00001-6 Copyright © 2024 Elsevier Inc. All rights reserved.

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18.2 Market trend Recently Global Industry Analysts Inc. (GIA), a reputed market research company, has published their analysis on hydrogel market of post-COVID era. According to their analysis, global hydrogel market is likely to reach 16 billion by the year 2027 [13]. While consumption-wise, hygiene products remain to be the largest segment, there will be consistent growth in all the fields related to biomedical sectorsdwound care, drug delivery, and tissue engineering. In the drug delivery side, the use of hydrogel is most prevalent in the field of wound care. In 2016, wound care products accounted for 69.48% of the total consumption. Consumption of biodegradable, nonbiodegradable, and bioactive hydrogels is projected to increase in the fast-growing drug delivery sector. Tissue engineering sector is likely to experience a high demand of synthetic hydrogels. Sales-wise, the United States of America is the leader in hydrogel market followed by Europe and China. Among the US companies, Acelity and ConvaTec have the maximum presence, while Smith and Nephew Limited and DSM have the same status in the European market. According to a report by Grand View Research, the market for hydrogel-based drug delivery is likely to expand at a compound annual growth rate (CAGR) of 7.5% from 2021 to 2028. Though hydrogels made from synthetic sources have greater tunability, the natural segment has the largest revenue share (32.3% in 2020) [14].

18.3 Hydrogels in drug delivery system To exert therapeutic action, a drug must reach its site of action which is often deep inside the body protected by physiological barriers. Delivery systems help the drugs to reach this site. An ideal drug delivery system should guide drugs to its target site without any collateral damage. Hydrogel dosage forms or hydrophilic polymer network is found to be ideally adjusted to serve this purpose [15]. By imbibing water, hydrogels become wet and flexible and this property helps it to align itself with the surrounding material which improves its biological acceptability [16]. Having a consistency similar to the living tissue, hydrogels don’t have harmful influences on the metabolic processes of the humans. As these properties are highly desirable in the drug delivery systems, hydrogel-based products have been found suitable for administration through almost all the drug delivery routes. They can also be used as vehicles for delivering sensitive biologics (like plasma and sera) and immunological products like vaccines. The unique properties of hydrogels are due to the materials used in its preparation. By definition, hydrogels are networks of hydrophilic polymers. Hence hydrophilic polymers like polysaccharides and polypeptides are the main ingredients of hydrogels. These moieties have surface functional groups (CH3)2N, OH, CONH, SO3H, and COOH which can bind water to their structure. Both polysaccharides and polypeptides are abundant in nature and are widely used in the formation of hydrogels. However, properties of the natural polymers vary widely depending upon the source. So synthetic polymers have also been developed to fulfill the need of tailor-able physical and mechanical properties [1,17]. Alginate and cellulose derivatives are the most common ingredients from plant kingdom finding application in commercialized hydrogels. Collagen, fibrin, heparin, hyaluronic acid, and gelatin are major animal-derived ingredients represented in the drug delivery hydrogels. As hydrogels can mimic the biphasic natural environment of the body, they are widely used in tissue engineering too. Possibly, the two most important biomaterials used in tissue engineering are collagen and hyaluronic acid. Collagen, an extravascular protein, is the most abundant protein in the human body. Type 1 collagen, which presents 90% of all collagens, is made of three polypeptide chains and has a molecular weight of around 300 kD. At the neutral pH of water, collagen fibrils self-assemble to form bundles. The closeness of the fibers increases the possibility of cross-linking, and ultimately a hydrogel is formed [18]. Avoiding immune recognition is a challenge in the success of tissue-engineered products [19]. It is shown that coating the nanoparticles with collagen reduces its affinity for albumin and thereby reduces the chance of immune recognition [19]. Another molecule, hyaluronic acid, a natural ingredient of bone and skin, is present in almost all the commercially successful antiaging cosmetics as well as skin and bone regeneration products. It helps in hydrating collagen, retaining its structural flexibility and freshness. Oral dosage forms, though the cheapest option for drug delivery, are not suitable for every type of drugs. For successful delivery through the oral route, drugs should have low molecular weight, the right balance of hydrophilicity and lipophilicity as well as be stable in the harsh conditions of gastrointestinal tract (GIT). According to the Rule of 5 to be permeable, drugs should have a molecular weight of around 500, calculated logP of 5, and the number of H-bond donors and acceptors should be restricted to 5 and 10 only. Many modern drugs don’t comply with this rule. Some effective drugs have very low aqueous solubility, while some others have poor stability. Inclusion of hydrophilic moieties improves solubility but more often than not, at the expense of gastrointestinal (GI) permeability. The unique properties of hydrogels can be used to ease the administration of such drugs [20,21].

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Drugs having protein structures get digested in the GIT. Usually, injecting them to the systemic circulation is the only option for administration. Yet there are candidates which show serious toxicity when administered by this route. One example is interleukin 12, a molecule with the potential to downregulate inflammatory response in cancer [22]. A clinical trial conducted to establish its efficacy resulted in the death of a volunteer due to systemic toxicity [23]. Naturally for such drugs, the focus of pharmaceutical research shifted to the design of novel dosage forms that would result in low toxicity, high efficacy, and easier methods of administration [3]. Hydrogel products are especially suitable for such drugs.

18.3.1 Hydrogel products used in the diseases of oral cavity According to a survey conducted by World Health Organization (Global Burden of Disease Study 2019), approximately 3.5 billion people suffer from the diseases of the oral cavity, with cases of permanent teeth being the most common. The other oral diseases that are prevalent throughout the world include gum disease, infection of oral mucosa, and oral cancer. These diseases can be best treated by local application of drugs to the buccal cavity. As buccal mucosa has excellent accessibility as well as good blood supply, the site can be used for systemic effects too. Here, absorption happens through the internal jugular vein, which can avoid the first pass metabolism [24]. However, two challenges limit the administration of drugs through this route. These are insufficient mucoadhesion which results in low residence time and accidental swallowing. Buccal dosage forms need to adhere to the cheek for prolonged periods. Hence, adhesiveness is a required property for these dosage forms. The hydrogel-based buccal dosage forms dissolve slowly, letting the medication enter into the bloodstream without getting swallowed [25] (Table 18.1). Buccastem tablets contain prochlorperazine which is used to control nausea and vomiting in diagnosed migraine. Prochlorperazine is poorly water soluble (0.011 mg/mL), and oral bioavailability of prochlorperazine is approximately 12.5%. Buccastem tablets have a composite hydrogel base (made from xanthan and locust bean gum) and are administered through the buccal route. Used alone, locust bean gum makes hyper-tangled macromolecular solution, whereas xanthan gum makes a weak gel network. Together, they make a strong gel with synergistic effects [26]. Prochlorperazine availability from the buccal route was found to be better than the same from intramuscular and oral routes. It is shown that steady-state concentration of prochlorperazine obtained after 3 mg twice daily administration by buccal route was equivalent to 5 mg thrice daily administration by oral route. Gingivitis, a periodontal disease, is mainly caused by the microorganisms present in subgingival plaque. Gengigel containing 0.2% hyaluronic acid is applied on the gingival mucosa to treat gingivitis [27]. Hyaluronan is highly viscoelastic in nature. It can act as a bacteriostatic against Aggregatibacter actinomycetemcomitans, Prevotella oris, and Staphylococcus aureus strains, usually found in oral gingival lesions and periodontal wounds. Hyaluronic acid, a natural constituent of gingiva and periodontal ligament, accelerates the healing of mouth ulcers by slowing the penetration of viruses and bacteria [28]. Hyaluronic acid also is highly osmotic and may hold up 10,000 times its weight in water. Its high osmotic property enables it to control the hydration of oral mucosa during inflammation and tissue injury [29]. A caseecontrol study conducted on Gengigel also established its efficacy in the treatment of gingivitis [30]. Chronic periodontitis, a bacterial infection of the gum, is a manifestation of substantial increase in rod-shaped anaerobic bacteria [31]. Poor oral hygiene causes bacterial build up and plaque formation which causes an immune response resulting

TABLE 18.1 Hydrogel products containing natural polymers meant for diseases of oral cavity [12]. Brand

Active ingredients

Polymers used

Indication

Manufacturer

Dosage form

Buccastem M

Prochlorperazine maleate

Povidone K30, xanthan gum, locust bean gum

Controls nausea and vomiting in migraine

Alliance Pharmaceuticals, Chippenham, UK

Tablet

Gengigel

Hyaluronan

Hyaluronan

Mouth and gum caregingivitis

Oraldent Ltd

Gel

Nicotinell

Nicotine

Xanthan gum, gelatin

Smoking cessation

Glaxo SmithKline, London, UK

Chewing gum

Nicorette

Nicotine

Hydroxypropyl methylcellulose

Smoking cessation

Glaxo SmithKline, London, UK

Chewing gum

Extracted from Ghosh B, Kirtania MD, Clinical applications of biopolymer-based hydrogels. In: Plant and algal hydrogels for drug delivery and regenerative medicine. 2020. p. 535e568.

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in inflammation and tissue destruction and tooth loss. Green tea catechin is identified with anticariogenic, antiinflammatory, anticollagenolytic properties, and found to be effective in improving periodontal status [32]. A chitosan-based hydrogel system is being evaluated against chronic periodontitis (CTRI/2021/09/036764). Medicated chewing gum is a unique type of dosage form which contains two types of ingredientsda water-insoluble gummy part and a water-soluble active ingredient [33]. The gummy part is usually an elastomeric polymer which makes it flexible against breaking and cracking [33,34]. Gelatin contained in the Nicotinell gum has elastomeric properties [35]. It is used in nicotine withdrawal symptoms associated with tobacco dependence. A number of buccal hydrogels have been made with synthetic and semisynthetic polymers too. Some of prominent products of this type are Lubragel (mouth moisturization), Hydrogel 15% (contains ozonized sunflower oil for cleaning oral cavity), Biotene (treats dry mouth), SCHALI (prevents infection of oral cavity), Zilactin B Gel (a buccal film to treat sores), Imdur (delivers isosorbide dinitrate for preventing angina), etc.

18.3.2 Hydrogels in oral delivery system Recently the repertoire of drugs has been enriched by peptides which can act as hormones, growth factors, neurotransmitters, ion channel ligands, or antiinfective agents. However these molecules need special strategies to make them permeable through the GI barrier [36]. Therapeutic peptides are big molecules having molecular weights in the range of 500e5000 Da. Basically, they are amino acid chains linked by amide bonds, lacking the stability of secondary or tertiary structures. Without protection, the amide bonds can be easily hydrolyzed at the acidic pH of stomach. Second, their oral delivery is also limited by their inability to cross the GI barrier. Hence protein and peptides are usually delivered as parenteral formulations [37]. Oral absorption of protein and peptides is restricted by three factors: hydrolysis, degradation by luminal enzymes, and inability to pass through the GI wall. Formulation of peptides in hydrogel dosage form along with permeation enhancers and enzyme inhibitors has opened the prospect of oral delivery for these macromolecules [20,38,39]. To achieve success in oral delivery, the proteolytic enzymes present in the GIT must be countered. Hydrogels made from polyethylene glycolgrafted acrylic polymers have exhibited this desirable property. Hydrogen bonding between carboxylic group of the methacrylic acid and oxygen of the polyethylene glycol chain can form pH-sensitive temporary complexes. The internal environment of GIT changes from site to site. Stimuli-responsive hydrogels make use of this change to deliver protein and peptides both locally and systemically [40]. Hydrogels rich in carboxylic groups shrink in the acidic media, which prevents the entry of acids inside the dosage forms, thereby preventing the hydrolysis of proteins [41]. In small intestine, the pH is around 7, which is favorable for deprotonation of COOH. The resultant ionic repulsion causes swelling of the polymer leading to an increase of the mesh size. The drugs entrapped in the polymer chain can get released through the pores under this condition. Theoretically, three transepithelial pathways are available for the passage of molecules from the intestinal lumen to bloodstream: transcellular active/carrier-mediated, transcellular passive, and paracellular. Passive absorption of large hydrophilic molecules is limited through the paracellular pathway which restricts the molecules with molecular radii >11Å. Permeation enhancers open up the intracellular tight junctions and thereby enhance the permeation of megamolecules. However, a concern exists regarding the use of enhancers as the same pathway can be used by toxins to enter the microcirculation [42]. Microfold cells, residing at the border colon and intestine, are unique in having high transcytotic capacity and low lysosomal activity. They can pick up materials from the intestinal lumen and deliver them to the immune cells. So they can be used for the delivery of the macromolecular drugs, provided the drugs can be carried intact up to site of their residence. Hydrogels can be used to carry the drugs intact up to this location.

18.3.3 Hydrogels in controlled release dosage form The average duration of GI residence of the oral dosage forms is around 40 h. Passive diffusion is the dominant mechanism of oral absorption, but only drugs of small and moderate sizes get absorbed passively. So maintenance of the plasma concentration to a desired level can only be achieved through preprogrammed drug release. Release of drugs from hydrogel dosage forms is diffusion controlled and depends upon the pore size and matrix tortuosity [43]. Hydrogels made from synthetic polymers allow greater control on these parameters and are mainly used in oral delivery. On the flip side, in the making of synthetic hydrogels, cross-linkers and chain progression initiators are widely used. Traces of them are often present in the final product too. Fortunately, human GI system is well protected against the harmful influences of toxic chemicals and microorganisms. Cost is an important consideration in the success of a commercial product. Synthetic

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polymers allow greater tuning in terms of molecular weight, mechanical strength, and cost less than the natural polymers. A number of drugs have been developed for oral delivery using hydrogels of synthetic origin. Below is a list of marketed products developed with synthetic hydrogels. Polymers are at the heart of hydrogel. The properties of a hydrogel depend upon the nature of its constituent polymers. The nature of functional groups of a polymer and its tacticity decide the affinity/interaction for other molecules and degree of cross-linking it can achieve while forming a network [44]. A wide range of synthetic polymers have already been developed depending on their intended usage in hydrogels. Hydrogels made of polyethylene glycol and polyvinyl alcohol find wide use in tissue engineering. Hydrogels made of cellulose derivatives are commonly used for drug delivery [25] (Table 18.2). Hydrogels are basically polymer network cross-linked by various forces; for example, ionic or hydrophobic interactions, temperature, electric field, pH change, and magnetic field. Fig. 18.1 depicts some important biomedical fields in which hydrogels have made important contributions. The first hydrogel on the market that served a medical purpose was a synthetic sponge, named Ivalon. Made of polyvinyl alcohol, the negatively charged Ivalon sponge assumed small surface area when compressed and dry. This small piece was delivered through a catheter as an embolic material to the site of action. In contact with blood/fluid, it absorbed water and expanded in size and stuck to the affected area. Ivalon was well accepted, but it is the polyethyl methacrylatebased soft contact lenses that revolutionized the biomedical market of hydrogel [45].

18.3.4 Injectable hydrogels Dosage forms meant for injection have to comply with strict regulatory standards. In addition, they should be affordable enough to ensure marketability. Injectable hydrogels are well suited for local drug delivery, releasing drug at controlled

TABLE 18.2 Commercial hydrogels from controlled release polymers [25]. Product (molecule)

Company

Excipients

Advil (ibuprofen)

Pfizer Inc

Hydroxypropyl methylcellulose

Aplenzin (bupropion hydrobromide)

Valeant Pharmaceuticals International, Inc.

Ethylcellulose and polyvinyl alcohol

Concerta (methylphenidate)

Alza Corporation

Hydroxypropyl methylcellulose (hypromellose) and polyethylene oxide

Gaviscon (sodium bicarbonate and aluminum carbonate)

Reckitt Benckiser Healthcare Ltd.

Sodium alginate and carbomer 974P

Kaletra (lopinavir-ritonavir)

AbbVie Ltd

Polyvinyl alcohol

Levora (levonorgestrel and ethinyl estradiol)

Mayne Pharma Inc.

Croscarmellose sodium

Lopid (gemfibrozil)

Pfizer Inc

Hydroxypropyl methylcellulose

Portia (levonorgestrel and ethinyl estradiol)

Teva Pharmaceutical Industries Ltd.

Hypromellose

Ranexa (ranolazine)

Gilead Sciences

Polyvinyl alcohol and hydroxypropyl methylcellulose

Suprax (cefixime)

Sanofi Aventis

Hydroxypropyl methylcellulose

Toviaz (fesoterodine)

Pfizer Inc

Polyvinyl alcohol and hydroxypropyl methylcellulose

Vicoprofen (hydrocodone bitartrate and ibuprofen)

AbbVie Ltd

Hydroxypropyl methylcellulose

Voltaren (diclofenac free acid)

GlaxoSmithKline

Hydroxypropyl methylcellulose and polyethylene glycol

Xartemis (oxycodone hydrochloride and acetaminophen)

XR Mallinckrodt Pharmaceuticals

Hydroxypropyl cellulose and polyvinyl alcohol

Extracted from Cascone S, Lamberti G. Hydrogel-based commercial products for biomedical applications: a review. Int J Pharm 2020;573:118803.

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FIGURE 18.1 Biomedical applications of hydrogels.

rate at the desired site of action, which helps to minimize the systemic side effects [46]. At present, they are getting routinely used in the reconstruction and repair of the tissues damaged by trauma, injury, and degenerative diseases [47]. The first hydrogel-based biological implant, Ivalon, was introduced as an embolic material to manage the GI bleeding and arteriovenous defects [45]. Now it is being used to effect microvascular decompression in the treatment of trigeminal neuralgia [48]. As biodegradability is a requirement for dosage forms that come in direct contact with the body tissues, injectable hydrogels are made from biodegradable polymers. Of the natural polymers, mainly collagen, gelatin, hyaluronic acid, and silk fibroin serve as excellent carriers in tissue engineering [49]. In the drug delivery side, they have made important contribution as efficient vehicles for therapeutic agents.

18.3.4.1 Injectable hydrogels in the treatment of prostate cancer One of the diseases which hydrogel-based products is effective in controlling is prostate cancer. Prostate cancer cells depend on testosterone to sustain its rapid growth and spread. A strategy to stop/minimize the growth of prostate cancer cells is to reduce the supply of testosterone. Testosterone production in the body is a complex process. In mammals, the production involves gonadotrophin-releasing hormone (GnRH). GnRH is produced by hypothalamic neurons. Flowing through the hypophyseal circulation, it reaches the anterior pituitary, where it binds with specific receptors to trigger the production and release of luteinizing hormone (LH) and follicle-stimulating hormone (FSH), which in their turn stimulate the production of sex hormones. In case of men, the product of this stimulation is testosterone [50e52]. Endo’s Vantas is a hydrogel-based implant containing histrelin acetate, a GnRH analog [53]. The cylindrical implant releases histrelin, which is a testosterone agonist. Under normal conditions, GnRH release is pulsatile. But the implant releases the GnRH in a continuous manner. The chronic exposure of the drug desensitizes LHRH receptors leading to a decrease LH production which lowers the testosterone level [54]. Starved of testosterone, cancer cells stop growing. When LHRH agonists like histrelin are first introduced to the patients, testosterone levels went above the normal value before falling to subnormal levels. This was called tumor flare. Tumor flare, although a temporary phenomenon, can cause significant damage. So, strict control should be exercised on drug release to counteract the effects of tumor flare. Hydrogels made of synthetic polymers allow greater manipulation in drug release than their natural counterparts. So far, all the hydrogels used in prostate cancer treatment are made of synthetic polymers (Table 18.3). During radiation therapy of cancer, nearby organs are also accidentally exposed. SpaceOAR is a hydrogel that spaces the organs at risk (OARs) during radiation therapy. SpaceOAR, an absorbable form of polyethylene glycol, is used in the treatment of prostate cancer. Injected into the perirectal space, it pushes the rectal wall away from the prostate which partially protects the rectum from the intense radiation delivered in prostate during the therapy [55]. The protective effect of SpaceOAR is supported by the findings of meta-analysis of seven studies where patients used the spacer during prostate

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TABLE 18.3 Some commercially successful injectable hydrogels [53]. Year of regulatory approval

Product

Company

Drug/polymer

Indication

Vantas

Endo Pharmaceuticals

Histrelin acetate, poly(2hydroxyethyl methacrylate), poly(2-hydroxypropyl methacrylate)

Palliative treatment of prostate cancer

FDA 2004

SpaceOAR hydrogel

Augmenix, Inc.

Polyethylene glycol

For protecting healthy tissues during prostate cancer radiotherapy

FDA 2015

Gut guarding gel

National ChengKung University Hospital

Sodium alginate/calcium lactate

Gastroenterological tumor and polyps

NCT03321396 (NA)

Absorbable radiopaque tissue marker (between pancreas and duodenum)

Sidney Kimmel Comprehensive Cancer Center at Johns Hopkins

Polyethylene glycol/TraceIT

Imaging of pancreatic adenocarcinoma

NCT03307564

Absorbable radiopaque hydrogel spacer

Thomas Zilli, University Hospital, Geneva

Polyethylene glycol/TraceIT

Spacing in radiation therapy for rectal cancer

NCT03258541(NA)

Absorbable radiopaque tissue marker (resection bed)

Washington University School of Medicine

Polyethylene glycol/TraceIT

Imaging of oropharyngeal cancer

NCT03713021 (Ph I)

Absorbable radiopaque tissue marker (between pancreas and duodenum)

Sidney Kimmel Comprehensive Cancer Center at Johns Hopkins

Polyethylene glycol/TraceIT

Imaging of pancreatic adenocarcinoma

NCT03307564

cancer radiotherapy. Absorbable hydrogel spacer was used in two patients suffering from rectal cancer (one male and one female). For the female, the implant was introduced between rectum and vagina; for the male, it was placed between prostate and rectum before nCRT (neoadjuvant chemoradiation therapy) and curative surgery. Results indicated that vagina was well protected from radiation, but benefit was minimal in case of the male patient [56]. Administering drugs as subcutaneous implants is a favored strategy for delivering protein and peptide drugs that need prolonged administration at a controlled rate [57]. As injectable hydrogels can be designed and inserted into the exact site of action with minimally invasive techniques, they have become popular for creating in situ implants. Inserts/implants made of hydrogels are better tolerated than those made from other materials. SUPPRELIN LA, marketed by Endos, contains histrelin acetate, a synthetic hormone used to stave off early puberty in children. Histrelin, a nonapeptide, cannot be administered orally. Supprelin LA is a cylinder shaped (3.5 cm
3 mm) hydrogel implant, carrying 50 mg drug in its core. A study conducted on 33 children suffering from central precocious puberty resulted in excellent clinical response in favor of this implant [58]. In a phase III multicenter clinical trial too, Supprelin LA implant showed sustained gonadotropin suppression with a reasonable safety level [59].

18.3.4.2 Injectable hydrogels as dermal fillers Formation of wrinkles is a natural phenomenon. With aging, skin undergoes degenerationdboth structurally and functionally. Aging causes a decline in the production of elastin and collagen, the fibers responsible for firmness, and elasticity of skin [60]. The production of proteoglycans (PGs) and glycosaminoglycan by dermal fibroblasts is also reduced [61]. Glycosaminoglycans are hydrophilic molecules that keep the skin moisturized, thereby retain its volume, elasticity, and firmness. With decreasing level of glycosaminoglycan, skin starts losing its power of moisture retention, becomes dry, and develops depression and fine wrinkles [62]. Hence restoring the level of these essential molecules through direct injection into dermis has become a popular strategy to reverse the effects of skin aging.

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As research has shown that collagen supplementation can reverse the wrinkles, collagen was introduced into the cosmetic formulations for antiaging effects [63]. Zyplast and Zyderm were the first dermal fillers introduced into the market by Allergan, (formerly Inamed). They contained animal-derived collagen [64]. But, in some people, exposure to collagen causes allergy and skin test is required as a precaution before application of the products [65]. Popularity of bovine collagen as a dermal filler was highest in early 1980s. Then it became known that other molecules present in the extracellular matrix (ECM) also have a major influence on skin function. Hyaluronic acid, a key molecule that binds water to the skin [66], has become the new craze in the field of soft tissue augmentation. Oral absorption of hyaluronic acid remains a controversial issue, but they can be easily formulated into hydrogels. The effect of collagenbased dermal fillers lasted for twoefour months only. The effect of the hyaluronic acid-based fillers lasted longer. Juvederm, a hyaluronic acid-based hydrogel product, was introduced to compensate the age-related volume loss. Once injected into the skin, this product gets lodged into the loose folds of the skins and stay there until degraded. Due to its high affinity for water, hyaluronic acid attracts and binds interstitial water giving the skin an appearance of tightness. Later the same company (Allergan plc) introduced Juvederm XC which contains 0.3% lidocaine to ease the pain of injection [67]. Hyaluronic acid is effective against severe facial wrinkles too. Merz pharmaceuticals introduced Belotero balance (R)þ which contains hyaluronic acid. The product is intended to fill up the gaps created by loss of collagen in moderate to severe facial wrinkles. Injected into the dermis, it solidifies into the folds of the skin to create an appearance of youthfulness. The product got FDA approval in 2019. All these products, intended to fill up the gaps created by loss of collagen in moderate to severe facial wrinkles, were injected into the dermis as liquid. It solidifies into the folds of the skin to create an appearance of youthfulness (Tables 18.4 and 18.5).

18.3.5 Hydrogels used in the diseases of the skin Human skin, the multilayered organ that protects the human body from those environmental assaults, is a tough barrier, which allows few drugs to pass through it. Only low molecular weight drugs having high lipophilicity can cross this barrier well enough to elicit a pharmacological response. However, entry of hydrogels into the pharmaceutical market expanded this range. Hydrogels can moisturize the skin and enhance topical delivery compared to conventional dosage forms like emulsions and gels [68]. Research suggests that hydrogels have the potential to facilitate the delivery of liposomes and nanoparticles too [69]. Usually skin doesn’t favor the permeation of charged moieties [70]. But this shortcoming could be overcome by the application of electrical forces. Transdermal iontophoresis makes the use of electric potentials to drive the ionic drugs into the systemic circulation through the intact skin. Since the main driving force is electrical, the drug reservoir should be made of a conductive medium [71]. Hydrogels because of their water-holding capacity can conduct electricity while retaining its distinct three-dimensional structures. Patents demonstrating the use of hydrogels in drug reservoirs for iontophoretic drug delivery were also noted. The current trend of topical research demonstrates hydrogel formulations where the novel dosage forms like nanoparticles and microparticles are incorporated in hydrogel carriers [70,72] (Table 18.6).

TABLE 18.4 Collagen-based injectable hydrogels according to broad indication.a Chief ingredient

Year of FDA approval

Indication

Brand

Company

Wrinkle removal and contour deficiencies

Zyplast, Zyderm

Inamed corporation, Allergan Corporation

Collagen

1981

Correction of depressed cutaneous scar

Fibrel

Serono laboratories

Collagen

1988

Collagen Implant,a human-based collagen, CosmoDerm2 human-based collagen, CosmoPlast human-based collagen

Inamed Corporation/ Allergan, Inc.

Collagen

2003

a

Data extracted from Mandal A, Clegg JR, Anselmo AC, Mitragotri S. Hydrogels in the clinic. Bioeng Transl Med 2020; 5(2):e10158.

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TABLE 18.5 Hyaluronic acid-based injectable hydrogels according to broad indication [53]. Indication

Product

Company

Ingredients

Year

Correction of facial wrinkles and folds

Juve´derm XC

(Allergan, Inc.)

Hyaluronic acid with lidocaine

2010

Removal of deep wrinkles

Belotero balance

(Merz Pharmaceuticals)

Hyaluronic acid

2011

For correction of volume deficit, facial folds and wrinkles, midface contour deficiencies, and perioral rhytids

Restylane Lyft, Restylane Refyne, Restylane Defyne

(Galderma Laboratories, L.P.)

Hyaluronic acid with lidocaine

EMA (2010) FDA (2012)

For correction of volume deficit, facial folds and wrinkles, midface contour deficiencies, and perioral rhytids

Restylane Silk/Restylane Injectable Gel (Medicis Aesthetics Holdings, Inc.)

(Valeant Pharmaceuticals North America LLC/Medicis Aesthetics Holdings, Inc.)

Hyaluronic acid with lidocaine

EMA (2010) FDA (2012)

Facial wrinkles and folds

Teosyal RHA

(Teoxane SA)

Hyaluronic acid

2017

Moderate to severe facial wrinkles and creases

Revanesse Versa/Revanesse Ultra

Prollenium Medical Technologies, Inc.

Hyaluronic acid

2017

Moderate to severe facial wrinkles and creases

Revanesse Versaþ

(Merz Pharmaceuticals)

Hyaluronic acid with lidocaine

2018

Moderate to severe facial wrinkles and folds

Belotero balance (þ) Lidocaine

Merz Pharmaceuticals

Hyaluronic acid with lidocaine (chemical reaction)

2019

TABLE 18.6 Hydrogel-based commercial products used in transdermal drug delivery [25]. Main constituent

Polymers

Indication

Gensco Pharma

Lidocaine HCl (4%

Polyethylene glycol (PEG) 400

Fast pain relief 3e5 min

Collagen hydrogel mask

Skin republic

e

Collagen and sodium hyaluronate

Restore skin’s elasticity by hydrating it

Flector patch

IBSA Farmaceutici, Italy

Diclofenac epolamine

Gelatin and carboxymethylcellulose

Relief of arthritic pain and actinic keratosis

Lidoderm Lidotop patch

Teikoku Pharma-, Sanbonmatsu, Japan

Lidocaine

Gelatin and oleic acid

Relief of neuropathic pain

Lidothol patch

Terrain Pharmaceuticals

Lidocaine/ menthol

Carboxymethylcellulose sodium

Relieves nerve pain

Product

Company

Astero

18.3.6 Hydrogels in the diseases of the eye Human eye is a complicated sense organ that can be divided into anterior and posterior segments. In the anterior segment, we have cornea, conjunctiva, aqueous humor, iris, ciliary body, and lens, while the posterior segment consists of sclera, choroid, and retinal pigment epithelium [73]. Both the compartments of eye contain gelatinous semisolid material, which closely resemble hydrogels [74]. With increase in the life span of humans, the prevalence of the ocular diseases also increased. As per WHO’s VISION 2020 program, which aims to prevent blindness, about 1.4 million children under the age of 15 years suffered irreversible blindness, though 50% of them were preventable [75]. Apart from the diseases of refractive errors, common ocular diseases can be grouped into three classesdinflammation of the membranes (scleritis, keratitis, conjunctivitis, choroiditis, retinitis, and retinochoroiditis), hypertension-related

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FIGURE 18.2 Ocular delivery through injection.

(glaucoma), and age-related degeneration (macular degeneration and mucopolysaccharidosis) [74]. Many of these diseases are curable; provided drugs can be effectively delivered to their site of action. Yet in spite of the great advances of medical science, many of the eye diseases remain undiagnosed and lead to irreversible vision loss [76]. In the absence of other options, injecting drugs to the various chambers of the eye seemed to be the standard mode of ocular delivery (Fig. 18.2). The challenges in ocular drug delivery are mainly two. First is the prevention of drug loss by the lacrimal fluid; second, avoidance of the systemic absorption by the blood capillaries. Ocular bioavailability is affected by dilution in tears, lacrimal fluids, and corneal barriers. The combined effect of all these factors often reduces the ophthalmic bioavailability of drugs to less than 5%. When the site of action is close to the surface that is the anterior part of the eye, topical or subconjunctival applications are used. But when the target is deep inside the eye, intravitreal injection is the only option. Increase of the contact time can improve ophthalmic bioavailability. Ophthalmic hydrogels are suitable for intraocular as well as ocular surface therapy. To enter into the ocular compartments, drugs have to cross a series of physiological barriers like tear film, the corneal epithelial barrier, the endothelial barrier of iris capillary, the epithelial barrier of ciliary body, and the inner and outer retinal barrier. Keratoconjunctivitis sicca or dry eyes syndrome is a disorder of reduced tear production. Though not completely understood, it is believed that due to inadequate production/excessive evaporation; osmolarity of the tears goes up which activates the inflammatory cascade leading to the damage of ocular surface. The inflammatory cycle is broken by the application of cyclosporine which is available in the form of eye drops. However, eye drops need frequent administration due to short contact time (1e2 min) and continuous tear inflow. Hydrogels can stay on the ocular surface for prolonged period and help the drugs to penetrate into intraocular tissues, thereby provide sustained release. Another disease implicated for irreversible vision loss is glaucoma [77]. Some of the ophthalmic products made with natural hydrogels are presented in the Table 18.7.

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TABLE 18.7 Hydrogels used in ophthalmic products. Brand

Dosage form

Manufacturer

Active ingredients

Polymers used

Indication

Hylo gel

Eye instillation

Candorvision, Quebec, Canada

Sodium hyaluronate

Sodium hyaluronate, citrate, buffer, sorbitol

Treatment of chronic dry eye condition

Timoptic XE

Eye drop

Merck

Timolol maleate

Gellan gum

Glaucoma

XenGel

Subconjunctival implant

Allergan Inc. USA

None

Porcine gelatin

Glaucoma

Vismed Gel multi

Eye drop

TRB Chemidica, UK

None

Sodium hyaluronate

Dry eye syndrome

Neopt multi

Eye drop

Horus Pharma

Hyaluronic acid

Hyaluronic acid

Dry eye syndrome

Optive Fusion

Eye drop

Allergan India Pvt. Ltd

Sodium hyaluronate

Sodium carboxymethyl Cellulose, sodium hyaluronate

Dry eye syndrome

Triesence

Intravitreal injection

Alcon Laboratories, Inc.

Triamcinolone acetonide

Sodium carboxymethyl cellulose

UVites, Temporal arteritis

Alcon

Timolol maleate

Xanthan gum

Glaucoma

Timoptic GFS

One FDA-approved ocular healing hydrogel, ReSure sealant, is made of a synthetic PEG-based hydrogel. A similar formulation called DuraSeal originally approved for cranial adhesion has also been explored for ocular wound sealing [78]. Postoperative endophthalmitis, an infection-driven inflammation of the intraocular fluids, is a severe disorder that occasionally happens after cataract surgery. During phacoemulsification, a tiny incision (approximately 3e7 mm) is made in the cornea through which a needle-thin probe is inserted into the site of the cataract [79]. As these small incisions heal on their own, the procedure may be suture-less, as astigmatism induced by the surgery is minimal. However, leaving incisions without closure may increase the risk for wound leaks [80,81]. ReSure, a PEG-based hydrogel sealant, has recently been approved by the FDA for closing these leaking corneal incisions. ReSure is shown to be highly effective for this purpose in a recent study [82]. A number of biological drugs are also used in the treatment of eye diseases. Delivery of biologics is difficult as these drugs have short half-lives in vitreous humor. Encapsulation of these biologics into hydrogels can improve drug stability as well as extend the duration of drug release (Table 18.8).

18.4 Tissue engineering The purpose of tissue engineering is to create tissue constructs that would replace or restore the functions of failing organs. The materials used in this process are mainly of three typesdcells, biomaterials, and growth factors. Tissue engineering may also be used for other purposes like diagnosis, research (testing of safety and efficacy of new drugs), and development of extracorporeal life support system like artificial liver and kidney. To be functional, these artificial structures must have properties similar to the natural tissue. Human body is a unique creation. Though more than 60% of it is water, it maintains an exceptional balance of rigidity and flexibility. Against the various types of internal and external stresses, cells retain their structural and functional identity and act in harmony with the different parts through a sturdy network of nerves and vascular systems. Hydrogels having an intermediate nature of solid and liquid have wide use in tissue engineering. They provide the basic materials starting from small orthopedic implants like screws, pins, plates, and wires to complete solutions like total knee replacement devices [83]. The first tissue-engineered product that contained live cells was a skin construct. In 1979, researchers at the Harvard Medical School had isolated keratinocytes from biopsy material and expanded it several thousand times by culturing it over a feeder layer of mouse mesenchymal cells. This technology gave birth to the first man-made epidermis which was later marketed under the trade name “Epicel” by Genzyme to treat patients who suffered severe burn injuries [84]. Starting from Epicel, a large number of tissue-engineered products have reached the market and they can

284

TABLE 18.8 Synthetic hydrogel-based ophthalmic formulations currently on the market or under clinical trials. Company

Drug

Gelling polymers

Application

Status

Lumecare

Medicom

e

Carbomer 980

Dry eye syndrome

In market

Viscotears

Novartis

e

Carbomer 980

Dry eye syndrome

In market

Xailin Gel

VISUfarma

e

Carbomer 980

Dry eye syndrome

In market

Clinitas Gel

Altacor

e

Carbomer 980

Dry eye syndrome

In market

GelTears

Bausch & Lomb

e

Carbomer 980

Dry eye syndrome

In market

Liquivisc

Laboratoires THEA

e

Carbomer 974P

Dry eye syndrome

In market

ReSure Sealant

Ocular Therapeutix

e

PEG

Seal clear corneal incisions following cataract surgery

In market

Tiopex

THEA Laboratoires

Timolol Maleate

PVA, carbomer 974P

Glaucoma

In market

Pilopine HS

Alcon Laboratories

Pilocarpine hydrochloride

Carbopol 940

Glaucoma

In market

AktenTM

Akten

Lidocaine hydrochloride

Hypromellose

Ocular anesthesia

In market

Zirgan

Sirion Therapeutics

Ganciclovir

Carbomer 974P

Acute herpetic keratitis

In market

Virgan

Laboratoires THEA

Ganciclovir

Carbomer 974P

Acute herpetic keratitis

In market

DuraSite/ Azasite

Inspire Pharmaceuticals

Azithromycin

Poloxamer Polycarbophil

Bacterial conjunctivitis

In market

AzaSite plus

Insite Vision Inc.

Azithromycin and dexamethasone

Polyacrylic acid

Eye infections

Phase III clinical trial

Dextenza

Ocular Therapeutix

Dexamethasone

4-arm PEG N-hydroxy-succinimidylglutarate (20K)

Postsurgical ocular inflammation and pain

In market

Dextenza

Ocular Therapeutix

Dexamethasone

4-arm PEG N-hydroxy-succinimidylglutarate (20 K)

Allergic conjunctivitis

Phase III clinical trial

OTX-TP (insert)

Ocular Therapeutix

Travoprost

PEG

Glaucoma, ocular hypertension

Phase III clinical trial

OTX-TIC (implant)

Ocular Therapeutix

Travoprost

PEG

Glaucoma and ocular hypertension

Phase I clinical trial

OTX-TKI

Ocular Therapeutix

Axitinib

PEG

(AMD, DME, RVO)a

Phase I clinical trial

a Age-related macular degeneration, diabetic macular edema, retinal vein occlusion. Extracted from Fang G, Yang W, Wang Q, Zhang A,Tang B. Hydrogels-based ophthalmic drug delivery systems for treatment of ocular diseases. Mater Sci Eng C 2021;127:112212.

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be broadly classified into three categoriesdproducts meant for structural support (bone and cartilage), products to augment the transport and barrier properties of organs (skin and blood vessels), artificial organs for restoring the production of biochemicals, and secretary molecules essential for the body (liver and kidney). Among the tissue engineered products, Infuse is perhaps the most successful in the field of regenerative medicine. It is mainly used in the treatment of lumbar disc degenerative disease. Recently it has been approved for use in tibial shaft fracture and dental procedures too. Infuse bone graft consists of two componentsda recombinant human bone morphogenetic protein solution and a scaffold which serves initially as a carrier for the protein solution and later the reconstructed bone. These components are used together as a system. ECM plays a major role in tissue engineering. The ECM is composed of two main classes of macromolecules: PGs and fibrous proteins. Hydrogels have strong similarity with ECM [84]. As matrix material, the natural and human-derived biomaterials fare better than the synthetic ones in almost all accountsdcell adhesion, proliferation, and differentiation [85]. The techniques of tissue engineering involve implantation of cell which depends on ECM materials for growth and maintenance [86]. The power of tissue regeneration is limited in mammalians, but subjected to the stress or injury, certain cells acquire the capability to reconstruct the defective tissue [87]. But it needs proper encouragement in terms of nutrients and structural support. This observation has given birth to the idea of scaffolds. Scaffolds play home to the cells and motivate them to grow. They should either get absorbed or allow easy withdrawal, once natural process of reconstruction is over. Regeneration of the damaged organs has been attempted almost in every field. But success is limited. Three major fields where success has been achieved are skin, cartilage, and liver. Cell survival is a major challenge in tissue engineering. The planted cells need oxygen and nutrients to survive and function effectively. So creation of a vascular network is essential to supply this need. Usually, growth factors and nutrients were loaded in the matrix prior to or concomitant with cell seeding. Vascular network may be created through bioprinting and soft lithography [88,89]. Once sufficient revascularization has taken place, the wound is covered with autologous graft. Next, the host cells invade the matrix and proliferate within it causing the formation of the desired tissue. The most preliminary types of tissue-engineered products are skin substitutes. They are used as templates for dermal regeneration. To support vascularization, these templates are made porous. Skins developed through tissue engineering are mainly used in wound healing (discussed in detail in wound healing section). Real skin is multilayered. Below epidermis lies dermis which is rich in collagen and nutrients. There is a basal layer at the deepest part of epidermis, which stores the stem cells. Beneath the epidermis lie the dermal appendages like hair follicles, sweat glands, and sebaceous glands. To construct all these layers along with their appendages is not an easy job. The outermost layer does not have blood supply. It receives its nutrients from the underlying microvascular circulation of the dermal layer. Depending on the thickness of the injury, three types of regeneration take place: regeneration of the dermal layer, regeneration of the outermost layer of the skin, and regeneration of the full thickness skin. However, the most important function of the skin is protection. So regeneration of the skin starts with the regeneration of the epidermal layer (Fig. 18.3). The first artificial skin was made by Ioannis Yannas, a mechanical engineer and John F Burke, a burn surgeon at Massachusetts Institute of Technology. They prepared a cross-linked porous matrix for skin growth by mixing bovine type I collagen with chondroitin-6-sulfate and drying the same through controlled lyophilization [90]. Skins developed through engineering was partially successful in serving four purposes: protection of the wound from external environment, allow time for slow pace healing, supply of nutrients and matrix components to the wound bed, and supply of cells and structural components of connective tissue to allow vascularization for cell survival. Acellular dermal matrix or ADM is basically decellularized dermal layers, with thin layers of ECM obtained from cadavers or animals. It is a form of surgical mesh from which cells have been removed leaving the support structure intact. During tissue growth, the matrix can provide both structural and nutritional support. It is used for cosmetic and reconstructive procedures too [91]. Removal of the cellular components minimizes the risk of immunological rejection in ADM recipients. ADM is basically collagen which constitutes 30% of body proteins. Mainly collagen, gelatin, hyaluronic acid, and silk fibroin are used to prepare hydrogels for tissue engineering [49]. The presence of natural growth factor along with collagen and elastin in ADM stimulates the integration process of the mesenchymal cells into the matrix. This accelerates healing of the damaged tissue (Table 18.9). AlloDerm is a basically freeze-dried ECM obtained from cadaver skin [92]. It is used for both tissue repair and reconstruction. However, all the cellular materials were removed from the skin to prevent transmission of diseases and immunogenicity. As the template does not have any keratinocyte, it needs to be covered with an autograft for reepithelization. Being extremely pliable, it can be used for plastic surgical procedures too. AlloDerm was initially introduced for

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FIGURE 18.3 Recent perspectives for tissue engineering applications on skin: (A) through regeneration of dermis; (B and C) through regeneration of epithelium; (D) through regeneration of full thickness composite skin.

the treatment of burn patients. Later its use was expanded in a wide variety of orthopedic, urogenital, and dental procedures that involved soft tissue augmentation and guided bone regeneration [93]. Diabetes is a worldwide problem, and a sizeable fraction of diabetic patients suffer from neuropathy and peripheral venous insufficiency. Foot ulcers are common in such patients. Various types of dressings are available to treat this condition. Hydrogel dressing for foot ulcers is an addition to the existing repertoire. A comparison of hydrogel dressings with the other conventional dressings indicates hydrogel dressings are more effective in healing diabetic foot ulcers [94]. Apligraf, an FDA-approved bioengineered skin substitute, is being used in the treatment of diabetic foot ulcers and venous leg ulcers, which are resistant to standard care for more than one month [95]. Apligraf is a bilayered construct that generates both epidermis and dermis. To create the first layer, human foreskin-derived neonatal fibroblasts are cultured in bovine type I collagen matrix. Next, epidermal keratinocytes obtained from the same source are cultured over it to build a stratified structure. This is especially supportive of the healing for difficult wounds as both cells and matrix are available from the same structure. Apligraf is immunologically inert as it does not have Langerhans cells. TissueMend, an acellular collagen membrane, is mainly used for tendon augmentation. It is derived from bovine fetal tissue which is rich in type III collagen. The scaffold creates a biological environment that favors tissue healing. It facilitates repair of difficult tissues like Achilles tendon, rotator cuff, quadriceps tendon, and patellar tendon.

TABLE 18.9 Some important tissue engineering products [25]. Product

Company

Polymer/Application

Reference

Acellular allograft dermal matrix

Graft-jacket Now

Wright Medical Group

Collagen/reinforcement of tendon and ligamentous tissue (Achille tendon, foot ankle, hip knee)

https://www.wright.com/healthcareprofessionals/graftjacket-now

Acellular dermal matrix

AlloDerm SELECTTM Regenerative Tissue Matrix

Biohorizons

Collagen and elastin/graft for gingival augmentation

https://www.biohorizons.com/alloderm.aspx

Acellular dermal matrix

AlloDerm SELECT RESTORETM RTM products

Biohorizons

Collagen and elastin Restoration of the soft tissue

https://www.nature.com/articles/sj.bdj.2009. 759

Viscous barrier gel

Hyalobarrier Gel/Hyaloglide

Anika Therapeutic Inc., USA

Hyaluronic acid/prevention of adhesion between body surfaces

https://anika.com/medical/products/surgicalsolutions

Woven gauze used graft wrap

Hyalonect

Anika Therapeutic Inc., USA

Hyaluronic acid Treatment of pseudarthrosis

https://anika.com/

Non-denatured collagen matrix

TissueMend

Stryker

Collagen Soft tissue repair

https://www.stryker.com/us/en/sports-medicine/ products/tissuemend.html

Acellular dermal matrix

Zimmer Collagen Repair Patch

Zimmer Biomet

Collagen and elastin from porcine dermis Repair of tendon injuries

ZimmerÒCollagen Repair Patch Surgical Technique (zimmerbiomet.com)

Collagen scaffold

OrthADAPT Bioimplant

Pegasus Biologics, Inc.

Collagen/repair of soft tissue in musculoskeletal procedures

Pegasus Biologics Inc.’ OrthADAPT(TM) Bioimplant Receives CE Mark | BioSpace

Injectable hyaluronate gel

Gel-One Hyaluronate

Zimmer Biomet

Hyaluronic acid Osteoarthritis of the knee

Gel-One Cross-linked Hyaluronate (zimmerbiomet.com)

Extracellular matrix hydrogel

Ventrigel

Ventrix

Phase I clinical trial for correction of cardiac damage

A Study of VentriGel in Post-MI Patients - Full Text View - ClinicalTrials.gov

Bilayered bioengineered skin substitute

Apligraf

Organogenesis

Collagen and fibroblast/living cells and structral proteins

Apligraf Living Cellular Skin Substitute

Soft tissue implant

Restore

DePuy, Inc.

Collagen and water/orthopedic tissue/orthobiologic soft tissue implant

http://www.accessdata.fda.gov/cdrh_docs/pdf3/ k031969.pdf

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Submucosal tissue of small intestine can serve as a wonderful scaffold material for tissue growth [96]. The Restore orthobiologic soft tissue implant is a round device, manufactured from submucosa of pig small intestine. It is composed predominately of water and collagen and is used for reinforcement of the soft tissues weakened during rotator cuff repair. The Restore implant provides a resorbable scaffold which is replaced by the growth of patient’s own soft tissue [97].

18.5 Wound repair The ability to heal the wounds is natural in humans, but in complicated cases like burns and diabetes, effective treatment is needed [98]. When the wound is deep, healing is not spontaneous and the area needs to be covered by skin substitutes [99]. Wound healing is a multistep proceduredhemostasis, inflammation, proliferation, and remodeling [100]. Strategies used include dressing wounds with polymeric scaffold, direct injection of cells into the wound site, or implantation of the cells encapsulated within the biomaterials [101]. Among the commercially available hydrogel products related to drug delivery, wound dressings are the most common [102]. The perfect wound dressing should mimic the human skin in every way. It should maintain adequate humidity, allow gaseous exchange, and protect the underlying tissue from bacterial invasion. Materials applied to the wound should be nontoxic and biocompatible. It should support painless application and removal and be able to control tissue necrosis. Mechanistically wound dressing can be classified into three types: passive, interactive, and bioactive products [103]. While the passive type acts only as barrier membrane, the other two serve as an active interface between the wound and the environment. Bioactive dressings deliver bioactive ingredients to the wound site. Complying all the specifications is a difficult proposition, but hydrogel dressings fulfill most of them. Hydrogel dressings provide an excellent moist environment for healing, which brings comfort to the patients. They are most suitable for dry or dehydrated wounds. They are also effective against radiation-related skin damage (after radiation therapy) but usually need a secondary dressing to keep them maintained in place and prevent their drying out. Because of the moist environment, hydrogels accelerate the wound healing [104]. Many of them contain natural alginate. Hydrogel-based wound-healing products are available in three forms: amorphous free-flowing hydrogels packed in tubes, spongy gauge pads impregnated with hydrogels, and fiber mesh-supported sheet hydrogels. Alginate is gifted with properties that fulfill the requirement for wound dressing. They are bacteriostatic, a property desired in products used in infected foot ulcers [105]. Some dressings contain both alginate and chitosan as the combination is thought to increase the absorbability. Alginate dressings are available in the form of freeze-dried porous sheets or foams. Alginate is a natural copolymer having glucuronic acid and mannuronic acid blocks. Higher proportion of mannuronic acid results in soft gels, whereas increased proportion of glucuronic acid produces firm gels. Usually to form hydrogels, alginates cross-link with calcium ions. In contact with the wound exudates, which contain calcium ions, a layer of calcium alginate is created over the wound. However, the film lacks mechanical strength. So a second support is usually required. Alginate’s inherent hemostatic property also favors wound healing [102]. In full-thickness wound, damage may extend into the subcutaneous tissue and underlying organs, after crossing all layers of the skin. If the size of the defect is large, usual method of reconstruction, split thickness skin graft, may not bring satisfactory result. Scar formation, inadequate elasticity, and hypo- or hyperpigmentation are drawbacks of this method. Combined use of splitethickness skin graft and ADM seems to overcome these limitations. Skin substitutes may be categorized into three groups: acellular scaffolds, temporary substitutes containing allogeneic skin cells, and permanent substitutes containing autologous skin cells. Some of the wound-healing products use the natural ECM as the base material for cell growth which is the idea behind ADM. The presence of natural growth factor along with collagen and elastin in ADM stimulates the integration process of the mesenchymal cells into the matrix, which accelerates wound healing. In the hyperglycemic environment, immune cells produce proinflammatory cytokines and wound healing is delayed. In this environment, matrix metalloproteinase causes breakdown of the ECM molecules, as soon as they are made by fibroblasts leading to chronic inflammation. Insulin serves as a chemoattractant and mitogen in the wound-healing cascade [106]. Injection of hydrogel (polyethylene glycol diacrylate)-encapsulated pancreatic islet cells caused healing of the wounds by creating normoglycemia. The presence of the hydrogel layer protects the islets from immune rejection. The immune-isolated cells have increased viability, and the released insulin accelerates wound healing [107] (Table 18.10).

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TABLE 18.10 Some natural hydrogels used in wound healing [25,108]. Brand

Manufacturer

Ingredient

Type

Indication

Hyalofill-F and R

Anika, Padua, Italy

Hyaluronic acid

In fleece and rope

Absorbs wound exudate, and promotes Granulation

Helix3-cm

Amerx Health Care Corp, Florida, USA

Type I bovine collagen

Dermal gauze pad

Management of burns, sores, blisters, ulcers, and other wounds

Purilon gel, Regenecare wound gel

Coloplast Corp., Minneapolis, USA

Lidocaine (2%), collagen, aloe, and sodium alginate

Wound gel

Pressure ulcers, cuts, burns, and abrasions

HemCon bandage PRO

TriCol Biomedical Inc., Oregon, USA

Chitosan

Bandage

Providing hemostasis, antibacterial barrier against wide range of microorganism

Condress

Smith and Nephew

Collagen

Spray

Chronic and acute wounds

Kaltostat

Convatec

Alginate

Dressing

Moderate to heavily exuding chronic and acute wounds. Should be applied dry.

NU-GEL

Systagenix

Alginate

Dressing

Management of dry to mildly exudating wounds throughout all stages of healing

18.6 Conclusion The addition of hydrogels in the biomedical field has expanded the possibilities for drug delivery and opened avenues for tissue engineering too. In the drug delivery field, injectable hydrogels have replaced many surgical procedures with minimally invasive techniques. In the tissue engineering field too, hydrogel-based devices and orthopedic implants have been successful to ease the treatment of cartilage and bone-related disabilities. Advanced level work is going on to develop hydrogels based oral implants [109], antiinfective orthopedic devices [110], as well as hydrogels for kidney [111], liver [112], and heart tissue regeneration.

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Further reading [1] Cai MH, Chen XY, Fu LQ, Du WL, Yang X, Hu PY. Design and development of hybrid hydrogels for biomedical applications: recent trends in anticancer drug delivery and tissue engineering. Front Bioeng Biotechnol 2021;9. [2] Berthiaume F, Maguire TJ, Yarmush ML. Tissue engineering and regenerative medicine: history, progress, and challenges. Annu Rev Chem Biomol Eng 2011;2:403e30.

Index ‘Note: Page numbers followed by “f ” indicate figures and “t” indicate tables’.

A Acetylsalicylic acid (AAS), 38 Acquired immunodeficiency syndrome (AIDS), 207 Addition reactions, cross-linking by, 133 Adipose-derived stem cell (ASC), 78, 165e166 Advanced in vivo imagining, 134 Algae-derived polysaccharides, 1 Amide coupling, 72e74, 75f Animal-derived polysaccharides, 1 Animal model, evaluation parameter in, 208e209 Animal sources, 2e7 chitosan, 3e4, 4f chondroitin sulfate, 5e7, 6f glycogen, 2, 3f heparin, 5, 5f hyaluronic acid, 2e3, 4f Antiadhesive properties, 135 Antifungal properties, 135 Antigen-presenting cells (APCs), 192, 207 Antiinflammatory, 177 Antineoplastic activity, 189e191 Antioxidant, 177 Antisense oligonucleotide (AS-ODN), 192 Antitumor, 177 Aqueous-based formulations, 262e263 Associated potential applications, 176e177 Azido-modified pullulan, 165

B Biocompatibility, 1 Biodegradability, 1 Biological function, 189e197, 190f Biotechnological approach, 14e16, 15f Bleeding control, 134 Blood substitute, 135 Bone tissue regeneration, 24e25, 97e99, 98f gellan gumebased hydrogels, 116e119, 118f Bottom-up approach, 164 Bovine serum albumin (BSA), 39 1,4-butanediol diglycidyl ether (BDE), 181, 183f

C Carboxymethylation, 206e207 Carboxymethyl cellulose, 166e167

Carboxymethyl-curdlan (CMCD), 206e207 Cartilage and bone regeneration, 41e42 Cartilage tissue regeneration, 99e101, 100f, 102f gellan gum-based hydrogels, 120e122, 121f Cell-free synthesis, 178 Cellulose nanofibrils (CNFs), 41 Charoenlarp, 169 Chemically crosslinked HEP HGLs, 69e70 Chemical methods, 132e133 addition reactions, cross-linking by, 133 click chemistry, 133 condensation reaction, cross-linking by, 133 cross-linking using enzymes, 133 Schiff base reaction, 132e133 Chitosan chitosan-based hydrogel application, 59e63 drug delivery, 59e62, 60t nasal delivery, 61 ocular delivery, 61 oral delivery, 59 parenteral delivery, 61e62 commercially available chitosan-based hydrogel, 63, 63te64t hydrogel classification, 47e49, 48f crosslinked junctions, 48e49 crosslinking, 48 electric charge, 48 hydrogel preparation, 49e59 chemical crosslinking, 49e51, 49f, 50t crosslinkers, 49e51 photopolymerization, 51 physical crosslinking, 51e54, 52t, 53f ionic crosslinking, 51e53 polyelectrolyte complex, 53e54 smart hydrogel, 54e59, 54f, 55te56t electro-responsive hydrogels, 57e58 magneto-responsive hydrogels, 58 nanogels, 58e59 pH-sensitive hydrogels, 57 temperature-sensitive hydrogels, 54e57 tissue engineering, 62e63, 62t Chitosan nanogels, 216 characterization methods, 228e229 chitosan (CS), 215e216, 215f chitosan-based nanogels, 218 chitosan derivative-based nanogels, 218e220 CMKGM/CS nanogel formation, 222f dentistry, 227e228 different delivery approaches, 225e226 drug delivery, 217e226, 228e229

functionalized chitosan-based nanogels, 220 other polymers, 221 preparation methods, 216e217, 217f regenerative medicine, 226e228 stimuli-responsive chitosan nanogels, 221e225 dual-responsive chitosan nanogels, 223e225 pH-responsive chitosan nanogels, 223 temperature-responsive chitosan nanogels, 221e223 structure of, 219f wound healing, 226e227 Chlorhexidine (CHX), 209e210 Chondroitin sulfate nanocomposites bioactivities of, 250e251 drug delivery applications, 252e253 future directions, 257 nanocomposite hydrogel systems, 251e252 codependency, 252 crosslinking method, 251e252 inwards diffusion method, 252 in situ method, 251e252 other applications, 256e257 structure, 250e251, 251f tissue engineering, 253e255 wound dressing, 255 Chronic obstructive pulmonary disease (COPD), 209 Click chemistry, 133 Codependency, 252 Collagen-hydroxyapatite hydrogels, 23 Colon, 113 Colon-targeted drug delivery, 145 Combination drug therapy, 161 Combined interaction (duel crosslinked), 80e81 Commercially available chitosan-based hydrogel, 63, 63te64t Condensation reaction, cross-linking by, 133 Controlled release dosage form, 276e277, 277t Coprecipitation and solvothermal method, 239 Cosynthesis method, 238 Covalent bonding, 69 Covalent functionalization, 238 Crosslinking method, 133, 251e252 Cryogels, 211 Crystallization, 132

293

294 Index

Curdlan, 13e14, 14fe15f animal model, evaluation parameter in, 208e209 application of, 209e211 biosynthesis of, 203e205 chemical structure of, 204f cryogels, 211 dental drug delivery, 209e210 drug delivery, 205e206, 209e211 metabolic pathway of, 204f oral drug delivery, 209 polymer properties with, 205 polysaccharides, 203 preparation method, 206e208 carboxymethylation, 206e207 phosphorylation, 207 sulfation, 207e208 protein delivery vectors, 210e211 tissue engineering application, 211 skin tissue regeneration, 211 transdermal drug delivery, 210 uridine diphosphate (UDP)-glucose, 203e204 uridine triphosphate (UTP), 203e204

D Dermatological applications, 36e37, 36f Dextran (DEX), 8e9, 9f, 129 biomedical application, 133e135 advanced in vivo imagining, 134 bleeding control, 134 drug carrier, 134 therapeutic vascularization, 134 tissue engineering, 133e134 tissue regeneration, 134 wound healing, 134 chemical methods, 132e133 addition reactions, cross-linking by, 133 click chemistry, 133 condensation reaction, cross-linking by, 133 cross-linking using enzymes, 133 Schiff base reaction, 132e133 drug delivery antiadhesive properties, 135 antifungal properties, 135 blood substitute, 135 wound-healing properties, 134e135 gelation, 131 hydrogel, 129 physical methods, 131e132 crystallization, 132 freeze-thawing, 132 heating-cooling polymer, 132 hydrogen bonding, cross-linked hydrogel by, 131 ionic interaction, 132 properties of, 131 source of, 130 structure of, 130, 130f Dibenzocyclooctyne (DBCO), 74e76 Differential scanning calorimetry (DSC), 39e40 Di-N-butyltin-dilaurate (DBTDL), 164

Divinyl sulfone (DVS), 38 Drug carrier, 134 Drug delivery, 16, 35f, 59e62, 60t, 177

E Electrically responsive transdermal delivery systems (ETDSs), 162e163 Electrospinning, 237e238 Electrostatic interaction, 76e77 Enzymatically responsive HGLs, 83 Enzyme-mediated crosslinked HGLs, 74 Epichlorohydrin (ECH), 210 Ethylene glycol diglycidyl ether (EGDGE), 210 Ethylene oxide (EO), 78 Exopolysaccharides, 211 Extracellular matrix (ECM), 36 Eye diseases, 281e283

F Fetal bovine serum (FBS), 166e167 Film formation, 177 Fixed concentration, 171 Fluorescein isothiocyanate (FITC), 171 Freeze-thawing, 132

G Gelation, 131 behavior, 188e189, 189f Gellan gumebased hydrogels, 8, 8f bone tissue regeneration, 116e119, 118f cartilage tissue regeneration, 120e122, 121f chemical structure, 110, 110f gelling properties, 111 hydrogel in drug delivery, 111e115 colon, 113 intestine, 113, 113fe114f oral cavity, 111e112 oral drug delivery, 111e113 stomach, 112 nasal drug delivery, 114e115, 115f neuron tissue regeneration, 122e124, 123f ocular drug delivery, 114 properties, 110 regenerative medicine, hydrogel in, 115e124, 116f skin tissue regeneration, 119e120 source, 109 types, 110 Gellan gum, in situ gels based on applications, 263e267 aqueous-based formulations, 262e263 compatibility of, 262e263 delivery routes, 267 injectable in situ gelling drug delivery, 267 intranasal in situ gelling drug delivery, 266 lipid-based formulations, 263 mechanisms of, 261e262, 264te265t ophthalmic in situ gelling drug delivery, 266 oral in situ gelling drug delivery, 266 Gene delivery, 191e195

Genipin, 165 Glutathione-responsive HGLs, 83e84 Glycogen-based hydrogels chemical structure, 22 composition, 22 drug delivery applications, 23e24 forming ability, 22e23 physicochemical properties, 22 source, 21e22 tissue engineering applications, 24e30 bone tissue regeneration, 24e25 self-healing hydrogel, 27e30 skin tissue regeneration, 25e26 wound healing, 25e26 tumor targeting, 24

H Heating-cooling polymer, 132 Heparin (HEP) amide coupling, 72e74, 75f chemically crosslinked HEP HGLs, 69e70 enzyme-mediated crosslinked HGLs, 74 HGLs, 69e76 Michael-type addition reactions, 71e72, 73fe74f other covalent bonding approaches, 74e76 photo-crosslinked HEP HGLs, 70 physically crosslinked, 76e78 combined interaction (duel crosslinked), 80e81 electrostatic interaction, 76e77 hosteguest interaction HGLs, 77e78 hydrogen bonding, 78 hydrophobic interaction HGLs, 78, 80f other physical interactions, 78, 81f smart HGLs, 81e84 stimuli-responsive HEP HGLs, 81e84 enzymatically responsive HGLs, 83 glutathione-responsive HGLs, 83e84 structure of, 70f Heparin-based nanocomposite hydrogels, 234e236, 234f application of, 239e245 biological, 235 biological systems, 236 coprecipitation and solvothermal method, 239 cosynthesis method, 238 covalent functionalization, 238 drug delivery, 239e240 electrical, 235 electrospinning, 237e238 layer-by-layer assembly, 238 magnetic, 235 mechanical, 235 nanocomposites, 240e245 physical self-assembly method, 238e239 preparation of, 237e239 spontaneous emulsion solvent diffusion method, 239 tissue engineering, 240e245 types of, 235 Hexamethylene diisocyanate (HMDIC), 169 Horseradish peroxidase (HRP), 83

Index

Hosteguest interaction HGLs, 77e78 Human mesenchymal stem cells (hMSCs), 70 Human umbilical vein endothelial cells (HUVECs), 70 Hyaluronic acid (HA), 36 cartilage and bone regeneration, 41e42 dermatological applications, 36e37, 36f drug delivery system, 35f extracellular matrix (ECM), 36 hydrogel structure, 35f inhalable hyaluronic acid hydrogels, 39e40, 39f injectable hydrogels, 38e39, 38f neuroregeneration, 42 ophthalmic applications, 37e38, 37f stem cells, 43, 43f tissue engineering, applications in, 40, 40f wounds treatment, 42 Hyaluronic acid aldehyde (HA-CHO), 39 Hydrogel, 129 classification, 47e49, 48f crosslinked junctions, 48e49 crosslinking, 48 electric charge, 48 preparation, 49e59 chemical crosslinking, 49e51, 49f, 50t crosslinkers, 49e51 photopolymerization, 51 structure, 35f Hydrogel-forming ability, 2 Hydrogen bonding, 78 cross-linked hydrogel by, 131 Hydrophobic interaction HGLs, 78, 80f

I Immunoadjuvants, 195 Indomethacin, 209 Inhalable hyaluronic acid hydrogels, 39e40, 39f Injectable hydrogels, 38e39, 38f, 277e280 dermal fillers, 279e280, 280te281t prostate cancer, 278e279, 279t Injectable in situ gelling drug delivery, 267 In situ method, 251e252 Intestine, 113, 113fe114f Intranasal in situ gelling drug delivery, 266 Intraocular pressure (IOP), 37e38 Inwards diffusion method, 252 Ionic interaction, 132

L Layer-by-layer assembly, 238 Levan, 11e12, 12f antiinflammatory, 177 antioxidant, 177 antitumor, 177 associated potential applications, 176e177 drug delivery systems, 177 film formation, 177 future perspectives, 181e184 levan-based hydrogels, 179e181, 182t

classification and preparation techniques, 180e181 rheology of, 179e180 mechanism, 175e176 prebiotic, 177 production mechanisms, 177e178 cell-free synthesis, 178 microbial production, 177e178 purification, 178, 179f properties, 176e177 solubility, 176 structure, 175e176, 176f support for cell proliferation, 177 tensile strength, 176 viscosity, 176 Levofloxacin (Lev), 38 Lipid-based formulations, 263 Long peptide antigen (LPA), 164 Low-density lipoprotein (LDL), 197 Low immunological activity, 1

M Matrix-assisted cell transplantation (MACT), 71e72 Matrix metalloprotease (MMP)-responsive, 71 Mesenchymal stem cells (MSCs), 167 Mesenchymal stromal cells (MSCs), 72e74 Methacrylate glycol chitosan (MeGC), 41 Methacrylic anhydride (MA), 80e81 Michael-type addition reactions, 71e72, 73fe74f Microbial and biotechnologically derived polysaccharides, 7e14 curdlan, 13e14, 14fe15f dextran, 8e9, 9f gellan gum, 8, 8f levan, 11e12, 12f pullulan, 10e11, 12f schizophyllan, 12e13, 13f scleroglucan, 9e10, 10f xanthan gum, 7, 7f Microbial organisms-derived polysaccharides, 1 Microbial production, 177e178 Migration inhibitory factor (MIF), 192e193

N Nasal delivery, 61 Nasal drug delivery, 114e115, 115f Neuron tissue regeneration, 104e105, 122e124, 123f Neuroregeneration, 42 Nonsteroidal antiinflammatory drug (NSAID), 36e37

O Octenyl succinic anhydride (OSA), 206e207 Ocular drug delivery, 61, 114 Ophthalmic applications, 37e38, 37f

295

drug delivery, 95 products, 283t in situ gelling drug delivery, 266 Oral cavity, 111e112 Oral drug delivery, 59, 111e113 controlled release, 92e95 in situ gelling, 266 sustained, 140e145, 141fe144f

P Paclitaxel (PTX), 195e196 Parenteral delivery, 61e62 Phosphate-buffered saline (PBS), 160 Phosphorylation, 207 Photo-crosslinked HEP HGLs, 70 Photopolymerization, 51 Physical crosslinking, 51e54, 52t, 53f, 76e78 combined interaction (duel crosslinked), 80e81 electrostatic interaction, 76e77 hosteguest interaction HGLs, 77e78 hydrogen bonding, 78 hydrophobic interaction HGLs, 78, 80f ionic crosslinking, 51e53 other physical interactions, 78, 81f polyelectrolyte complex, 53e54 Physical methods, 131e132 crystallization, 132 freeze-thawing, 132 heating-cooling polymer, 132 hydrogen bonding, cross-linked hydrogel by, 131 ionic interaction, 132 Physical self-assembly method, 238e239 Physicochemical properties, 22, 92 Plant-derived polysaccharides, 1 Platelet-derived growth factor (PDGF-BB), 168 Platelet extract (PE), 83 Poly (N-isopropylacrylamide) (PNIPAM), 23, 36 Polyethylene glycol (PEG) polymers, 38 Polyethylene oxide (PEO), 210 Poly(acrylamide)-graft-pullulan copolymer, 163 Polysaccharide-based hydrogels algae-derived polysaccharides, 1 animal-derived polysaccharides, 1 animal sources, 2e7 chitosan, 3e4, 4f chondroitin sulfate, 5e7, 6f glycogen, 2, 3f heparin, 5, 5f hyaluronic acid, 2e3, 4f biocompatibility, 1 biodegradability, 1 biomedical applications, 278f biotechnological approach, 14e16, 15f classification of, 1f controlled release dosage form, 276e277, 277t drug delivery, hydrogels in, 16 drug delivery system, 274e283

296 Index

Polysaccharide-based hydrogels (Continued ) eye diseases, 281e283 hydrogel-forming ability, 2 injectable hydrogels, 277e280 dermal fillers, 279e280, 280te281t prostate cancer, 278e279, 279t low immunological activity, 1 market trend, 274 microbial and biotechnologically derived polysaccharides, 7e14 curdlan, 13e14, 14fe15f dextran, 8e9, 9f gellan gum, 8, 8f levan, 11e12, 12f pullulan, 10e11, 12f schizophyllan, 12e13, 13f scleroglucan, 9e10, 10f xanthan gum, 7, 7f microbial organisms-derived polysaccharides, 1 natural, 2e14 ophthalmic products, 283t oral cavity, 275e276 oral delivery system, 276 plant-derived polysaccharides, 1 products, 275e276 regenerative medicine, 16 skin diseases, 280 synthetic hydrogel-based ophthalmic formulations, 284t tissue engineering, 283e288, 287t wound repair, 288, 289t Polyvinyl alcohol (PVA), 23, 196 Postsurgical tissue adhesion, 166 Propylene oxide (PO), 78 Pullulan-based hydrogels, 10e11, 12f drug delivery systems, 152e165 antiinflammatory drugs, 152e153 antimicrobial drugs, 153e158, 154fe155f, 157f cancer therapy drugs, 159e161, 160f diabetic disease drugs, 161e162 imagistic diagnostic drugs, 158 neurodegenerative diseases, 162e163 therapeutic proteins, 164e165 vaccine formulations, 163e164 wound healing drugs, 158e159 physicochemical properties, 151e152 regenerative medicine, 165e168 antiadhesion and tissue regeneration properties, 166e167 burn injuries hydrogels, 165e166 cartilage tissue engineering, 167, 168f dentin tissue engineering, 167e168 regenerative medicine, controlled and sustained release of drugs, 168e171, 170f, 172f rheological studies, 152 thermal properties of, 152 X-ray diffraction spectra, 152 Purification, 178, 179f

R Regenerative medicine, hydrogel in, 96e105, 115e124, 116f bone tissue regeneration, 97e99, 98f cartilage tissue regeneration, 99e101, 100f, 102f neuron tissue regeneration, 104e105 skin tissue regeneration, 102e104

S SARS-CoV-2 infection, 24 Schiff base reaction, 132e133 Schizophyllan (SPG), 12e13, 13f antineoplastic activity, 189e191 biological function, 189e197, 190f gelation behavior, 188e189, 189f gene delivery, 191e195 immunoadjuvants, 195 migration inhibitory factor (MIF), 192e193 other function of, 197 preparation, 187 structure, 187 wound dressing, 196 Scleroglucan, 9e10, 10f chemical structure, 140, 140f composition, 140 drug delivery applications, 140e146 colon-targeted drug delivery, 145 oral sustained drug delivery, 140e145, 141fe144f topical drug delivery, 145e146, 146f physicochemical properties, 140 source, 139 tissue engineering, 147 Self-assembled molecules, 164 Self-healing hydrogel, 27e30 Skin engineering, 165 Skin tissue regeneration, 25e26, 102e104, 119e120, 211 Smart HGLs, 81e84 Smart hydrogel, 54e59, 54f, 55te56t electro-responsive hydrogels, 57e58 magneto-responsive hydrogels, 58 nanogels, 58e59 pH-sensitive hydrogels, 57 temperature-sensitive hydrogels, 54e57 Sodium alginate (SA), 37 Sodium trimetaphosphate (STMP), 165e166 Solubility, 176 Spontaneous emulsion solvent diffusion method, 239 Staphylococcus epidermidis, 72e74 Stem cells, 43, 43f Stimuli-responsive HEP HGLs, 81e84 enzymatically responsive HGLs, 83 glutathione-responsive HGLs, 83e84 Stimuli-responsive hybrid materials, 160 Stomach, 112 Streptococcus pneumoniae, 163e164 Substrate specificity, 74

Sulfation, 207e208 Suprachoroidal space (SCS), 37e38 Synthetic hydrogel-based ophthalmic formulations, 284t

T Tensile strength, 176 Tetracycline hydrochloride (TCH), 207 Therapeutic vascularization, 134 Tissue engineering, 24e30, 62e63, 62t applications, 40, 40f bone tissue regeneration, 24e25 chitosan, 62e63, 62t chondroitin sulfate nanocomposites, 253e255 curdlan-based hydrogels, 211 DEX, 133e134 heparin-based nanocomposite hydrogels, 240e245 polysaccharide-based hydrogels, 283e288, 287t scleroglucan, 147 self-healing hydrogel, 27e30 skin tissue regeneration, 25e26 wound healing, 25e26 Tissue regeneration, DEX, 134 Topical drug delivery, 145e146, 146f Transdermal drug delivery, 210 Tumor cell lysate (TCL), 208 Tumor targeting, 24 Type-1 diabetes, 167 Tyramine-substituted carboxymethylated pullulan, 171

U Uridine diphosphate (UDP)-glucose, 203e204 Uridine triphosphate (UTP), 203e204

V Vaginal drug delivery, 95e96 Vascular endothelial growth factor (VEGF), 70 Viscosity, 176

W Water-soluble polysaccharides, 23 Wounds treatment, 25e26, 42, 134e135, 196 chitosan nanogels, 226e227 chondroitin sulfate nanocomposites, 255 Dextran (DEX), 134 polysaccharide-based hydrogels, 288, 289t pullulan-based hydrogels, 158e159 schizophyllan (SPG), 196

X Xanthan gum, 7, 7f chemical structure and composition, 90e91, 90f

Index

drug delivery applications, 92e96 ophthalmic drug delivery, 95 oral controlled release drug delivery, 92e95 physicochemical and rheological properties, 92 production, 91e92, 91f regenerative medicine, hydrogel in, 96e105

bone tissue regeneration, 97e99, 98f cartilage tissue regeneration, 99e101, 100f, 102f neuron tissue regeneration, 104e105 skin tissue regeneration, 102e104 source, 89 vaginal drug delivery, 95e96

Xanthomonas campestris, 90f X-ray powder diffraction (XRD), 39e40

Z Zinc chloride, 40

297