Biomaterials in Tissue Engineering and Regenerative Medicine: From Basic Concepts to State of the Art Approaches 9811600015, 9789811600012

Table of contents :
Contents
About the Editors
List of Abbreviations
Part I: Fundamentals of Biomaterials
1: Biomaterials, Tissue Engineering, and Regenerative Medicine: A Brief Outline
1.1 Introduction
1.1.1 Biomaterials
1.1.2 Tissue Engineering
1.1.3 Regenerative Medicine
References
2: Metallic Biomaterials in Tissue Engineering: Retrospect and Prospects
2.1 Introduction
2.1.1 Traditional Metallic Biomaterials
2.1.2 Advanced and Revolutionizing Metallic Biomaterials
2.1.3 Metallic Biomaterials and Biocompatibility
2.2 Properties of Metallic Biomaterials
2.2.1 Phase Transformation and Elastic Moduli
2.2.2 Porosity
2.2.3 Corrosion Resistance
2.2.4 Anti-Bacterial Properties
2.2.5 Bioactivation of Metallic Biomaterials
2.2.6 Biodegradation
2.2.7 MRI Compatibility
2.2.8 Radiopacity
2.3 Permanent Metallic Biomaterials
2.3.1 Stainless Steel
2.3.2 Co-Based Biomaterials
2.3.3 Ti-Based Biomaterials
2.3.4 Tantalum and Its Alloys
2.3.5 Zirconium Alloys
2.4 Biodegradable Metallic Biomaterials
2.4.1 Mg-Based Biomaterials
2.4.2 Zinc-Based Biomaterials
2.4.3 Iron-Based Biomaterials
2.5 Advanced Metallic Biomaterials
2.5.1 Bulk Metallic Glasses
2.5.2 Shape Memory Alloys
2.6 Tissue Engineering Applications of Metallic Biomaterials
2.6.1 Bone Tissue Engineering
2.6.2 Cartilage Tissue Engineering
2.6.3 Cardiovascular Tissue Engineering
2.6.4 Dental Tissue Engineering
2.7 Future Prospects of Metallic Biomaterials in Tissue Engineering
References
3: Bioceramics in Tissue Engineering: Retrospect and Prospects
3.1 Introduction
3.2 Background Perspective
3.3 Bioactivity of Calcium Phosphate
3.3.1 Calcium Phosphates: Variants and Effects
3.3.2 CaPO4 Bioceramics in Tissue Engineering
3.3.3 Clinical Vignettes
3.4 Summary and Outlook
References
4: Polymeric Biomaterials in Tissue Engineering: Retrospect and Prospects
4.1 Introduction
4.2 Extracellular Matrix-the Framework Enabling Tissue Growth
4.3 Polymeric Materials as Ideal Scaffold
4.4 Natural and Synthetic Polymers as Scaffolds
4.5 Natural Biodegradable Polymers
4.5.1 Collagen
4.5.2 Gelatin
4.5.3 Chitosan
4.5.4 Alginate
4.5.5 Fibrin
4.5.6 Hyaluronic Acid
4.5.7 Silk
4.6 Synthetic Biodegradable Polymers
4.6.1 Poly Lactic Acid (PLA)
4.6.2 Poly (glycolic acid) (PGA)
4.6.3 Poly (lactic-co-glycolic acid) (PLGA)
4.6.4 Poly(caprolactone) (PCL)
4.6.5 Poly Vinyl Alcohol (PVA)
4.6.6 Poly-β-hydroxybutyrate
4.6.7 Polyethylene Glycol-Based Polymers
4.7 Polymer Scaffold Fabrication Techniques
4.7.1 Conventional (Traditional) Manufacturing Techniques
4.7.2 Nano Fabrication-Based Techniques
4.7.3 Additive Manufacturing-Based Techniques
4.8 Conclusion and Perspectives
References
5: Composite Biomaterials in Tissue Engineering: Retrospective and Prospects
5.1 Introduction
5.2 Bio-Composite Components: Classes and Desirable Properties
5.3 Strategies of Bio-Composite Development
5.3.1 Conventional Blending and Mixing Technique
5.3.2 Advanced Bio-Fabrication Methods
5.3.2.1 Co-electrospinning
5.3.2.2 Bioprinting
5.3.2.3 Reinforcement Methods
5.3.3 Nano-Particle Reinforced Bio-Composites
5.3.4 Surface Modifications
5.3.5 Surface Effects and Characterization
5.4 Retrospectives of Composite Biomaterials in Tissue Engineering
5.4.1 Composite Biomaterials for Hard Tissue Regeneration
5.4.1.1 Bone Tissue Regeneration
5.4.1.2 Dentistry
5.4.2 Composite Biomaterials in Soft Tissue Engineering
5.4.2.1 Vascular Grafting
5.4.2.2 Cardiac Tissue Engineering
5.4.2.3 Contact Lens and Cornea
5.4.2.4 Neural Tissue Engineering
5.5 Bottlenecks of Composite Biomaterial Applications
5.6 Prospects of Composite Biomaterials
5.7 Conclusion
References
Part II: Trends in Biomaterials
6: Trends in Bio-Derived Biomaterials in Tissue Engineering
6.1 Introduction
6.2 Concept of Bio-Derived Biomaterials and their Applications in Tissue Engineering
6.3 Decellularized Extracellular Matrix (DECM) as Biomaterials
6.3.1 ECM and Decellularization
6.3.2 Methods of Decellularization
6.3.3 Regenerative Properties of DECM
6.3.4 Decellularized Material Systems: Applications in Tissue Engineering
6.4 Naturally Derived Biomaterials
6.4.1 Proteins Based Bio-Derived Biomaterials
6.4.1.1 Collagen
6.4.1.2 Gelatin
6.4.1.3 Fibrin
6.4.1.4 Silk
6.4.1.5 Keratin
6.4.2 Polysaccharides Based Bio-Derived Biomaterials
6.4.2.1 Glycosaminoglycans
6.4.2.2 Alginates
6.4.2.3 Agarose
6.4.2.4 Carrageenan
6.4.2.5 Chitosan
6.4.3 Other Bio-Derived Biomaterials
6.5 Microbial Derived Biopolymers
6.5.1 Types of Bacterial Polymers
6.5.2 Biosynthesis and Purification of Bacterial-Derived Polymers
6.5.2.1 Polyamides
6.5.2.2 Polyesters
6.5.2.3 Polysaccharides
6.5.3 Microbial Derived Biopolymers for Tissue Engineering
6.5.3.1 Poly-γ-Glutamic Acid (γ-PGA)
6.5.3.2 Polyhydroxyalkanoates (PHAs)
6.5.3.3 Polysaccharides
6.6 Conclusion and Future Directions
References
7: Trends in Functional Biomaterials in Tissue Engineering and Regenerative Medicine
7.1 Functionalized Biomaterials
7.2 Surface Functionalization Methods
7.2.1 Surface Roughening and Patterning
7.2.2 Surface Films and Coatings
7.2.2.1 Physical Methods
7.2.2.1.1 Physical Adsorption of Active Biomolecules
7.2.2.1.2 Langmuir-Blodgett Method
7.2.2.1.3 Physical Vapor Deposition
Evaporation
Deposition by Sputtering
Plasma immersion ion implantation and deposition (PIIIandD)
7.2.2.1.4 Electrophoretic Deposition
7.2.2.1.5 Spraying Techniques
7.2.2.2 Chemical Methods
7.2.2.2.1 Adsorption Via Covalent Bonding
7.2.2.2.2 Alkali Acid Hydrolysis
7.2.2.2.3 Chemical Vapor Deposition
Plasma-Enhanced Chemical Vapor Deposition
Plasma Polymerization
Atomic Layer Deposition
7.2.2.2.4 Sol-Gel Technique
7.2.2.2.5 Layer-by-Layer (LbL) Deposition
7.2.2.3 Radiation Methods
7.2.3 Surface Modification by Addition of Signaling Biomolecules
7.3 Functionalized Scaffolds Towards Organ Development
7.3.1 Cardiac Tissue
7.3.2 Liver
7.3.3 Lung
7.3.4 Bone
7.3.5 Dental Implants
7.4 Conclusion and Future Perspectives
References
8: Trends in Bioactive Biomaterials in Tissue Engineering and Regenerative Medicine
8.1 Tissue Engineering
8.2 Bioactive Scaffolds
8.3 Incorporation of Bioactive Components
8.3.1 Bioactivity by Incorporation of Adhesion Sites
8.3.2 Nanopatterning
8.3.3 Bioactivity by Incorporation of Growth Factors
8.3.4 Bioactivity by Physiochemical Interactions
8.3.5 Bioactivity by Material Transformation
8.4 Bioactive Inorganic Biomaterials for Tissue Engineering
8.5 Injectable Biomaterials
8.6 Bioactive Scaffolds: Tissue Engineering Applications
8.6.1 Neural Tissue Engineering
8.6.2 Vascular Tissue Engineering
8.6.3 Cardiac Tissue Engineering
8.7 Biomaterial Based Stem Cell Therapy in Regenerative Medicine
8.8 Scaffolds for Biomolecule Delivery
8.8.1 Properties
8.9 Biomolecule Delivery Systems
8.9.1 Hydrogel-Based Systems
8.9.2 Nanoparticle Based Systems
8.9.3 Liposomes
8.9.4 Micelles
8.9.5 Microparticles
8.9.6 Dendrimers and Elastomers
8.9.7 Microchips
8.10 Scaffold Based Biomolecule Delivery
8.10.1 Delivery of Therapeutic Drugs
8.10.2 Delivery of Therapeutic Cells
8.10.3 Scaffold Based Peptide Delivery
8.10.4 Scaffolds for Gene Delivery
8.11 Biomolecule Loaded Scaffolds in Tissue Engineering: Applications
8.11.1 Bone Tissue Engineering
8.11.2 Skin Tissue Engineering
8.11.3 Cartilage Tissue Engineering
8.12 Future Perspectives
References
9: Trends in Stimuli Responsive Biomaterials in Tissue Engineering
9.1 Introduction
9.2 Stimuli Responsive Biomaterials in Tissue Engineering
9.2.1 Electroactive Biomaterials
9.2.1.1 Conducting Polymers
9.2.1.1.1 Conducting Polymers in Tissue Engineering
9.2.1.2 Piezoelectric Material
9.2.1.2.1 Piezoelectric Materials in Tissue Engineering
9.2.1.3 Electrets
9.2.1.3.1 Electrets in Tissue Engineering
9.2.1.4 Photovoltaics
9.2.1.4.1 Photovoltaic Materials in Tissue Engineering
9.2.1.5 Carbon Based Nanomaterials
9.2.1.5.1 Carbon Based Nanomaterials in Tissue Engineering
9.2.2 Magnetoresponsive Biomaterials
9.2.3 Thermoresponsive Biomaterials
9.2.4 Photoresponsive Biomaterials
9.2.5 Chemical Stimuli Responsive Biomaterials
9.2.6 Biological Stimuli Responsive Biomaterials
9.3 Conclusions and Future Outlook
References
Part III: Applications of Biomaterials
10: Biomaterials for Hard Tissue Engineering: Concepts, Methods, and Applications
10.1 Introduction
10.2 Biomaterials for Bone Tissue Engineering
10.2.1 Polymers and Hydrogels
10.2.2 Hybrid Scaffolds in Bone Tissue Engineering
10.3 Applications of Tissue Engineering in Dentistry
10.3.1 Tooth Regeneration
10.3.2 Bone Regeneration in Dental Application
10.3.3 Enamel Regeneration
10.3.4 Dentin and Dental Pulp Regeneration
10.4 Biomaterials Used in Dentistry
10.5 Dental Stem Cells in Hard and Soft Tissue Engineering in Dentistry
10.6 Advanced Tissue Engineering Strategies
10.6.1 3D Printing in Hard Tissue Engineering
10.6.2 3D Bioprinting in Hard Tissue Engineering
10.7 Shape Memory Polymers in Hard Tissue Engineering
10.8 Tissue Engineering Challenges in Dentistry
10.9 Current Clinical Trials in Dentistry
10.10 Concluding Remarks and Outlook
References
11: Biomaterials for Soft Tissue Engineering: Concepts, Methods, and Applications
11.1 Introduction
11.2 The Properties of Scaffolds for Soft Tissue Engineering
11.2.1 Biological Properties
11.2.2 Physicochemical Properties
11.2.2.1 Cytotoxicity
11.2.2.2 Fabrication Techniques
11.2.2.3 Surface Properties of TE Scaffolds
11.2.3 Mechanical Properties
11.3 Application of TE Scaffolds in Soft Tissue Engineering
11.3.1 Vascular Tissue Engineering
11.3.1.1 Structure of Blood Vessels
11.3.1.2 Need for Vascular Tissue Engineering
11.3.1.3 Tissue Engineered Vascular Graft
11.3.1.3.1 Electrospun Scaffold-Guided Vascular Grafts
11.3.2 Skin Regeneration
11.3.2.1 Structure of Skin
11.3.2.2 Need for Skin Tissue Engineering
11.3.2.3 Tissue Engineered Skin Grafts
11.3.2.3.1 Injectable Hydrogels for Skin Tissue Engineering
11.3.2.3.2 Nanofibrous Scaffolds for Skin Tissue Engineering
11.3.3 Cartilage Tissue Engineering
11.3.3.1 Structure of Cartilage
11.3.3.2 Need for Cartilage Regeneration
11.3.3.3 Tissue Engineered Cartilage
11.3.3.3.1 Injectable Hydrogels
11.3.3.3.2 Nanofibrous Scaffolds
11.3.4 Intervertebral Disc (IVD)
11.3.4.1 The Structure of the IVD
11.3.4.2 Need for the Disc Repair
11.3.4.3 Tissue Engineered Disc
11.3.4.3.1 Nanofibrous/Hydrogel Scaffolds for Disc Repair
11.3.5 Tendon Repair and Regeneration
11.3.5.1 Structure of Tendon
11.3.5.2 Need for Tendon Repair
11.3.5.3 Tissue Engineered Tendon
11.3.5.3.1 Injectable Hydrogels Systems
11.3.5.3.2 Implantable Fibers System
11.3.6 Skeletal Muscle Tissue Engineering
11.3.6.1 Structure of Skeletal Muscle
11.3.6.2 Need for Skeletal Repair/Regeneration
11.3.6.3 Tissue Engineered Skeletal Muscle
11.3.6.3.1 Injectable Hydrogels for Skeletal Muscle Regeneration
11.3.6.3.2 Nanofibrous Scaffolds for Skeletal Muscle Regeneration
11.4 Future Perspective
11.5 Conclusion
References
12: Biomaterials for Specialized Tissue Engineering: Concepts, Methods, and Applications
12.1 Introduction
12.2 Biomaterials for Nerve Tissue Engineering
12.2.1 Nerve Guidance Conduits
12.2.1.1 Biological Conduits
12.2.1.2 Synthetic NGCs
12.2.1.3 Surface Micropatterning of NGCs
12.2.1.4 NGC Luminal Fillers
12.2.1.5 Stem Cell-Based NGCs
12.2.1.6 NGCs with Sustained Release of Growth Factors
12.2.1.7 Conductive NGCs
12.2.1.8 Carbon-Based Nanomaterial-Interfaced NGCs
12.2.1.9 Ultrasound Treatment Following NGC Implantation
12.2.1.10 Porcine Small Intestine Submucosa Made NGCs
12.2.2 Scaffolds for Nerve Tissue Engineering
12.2.2.1 Synthetic Scaffolds
12.2.2.2 Piezoelectric Scaffolds
12.2.2.3 Electroconductive Scaffolds
12.2.2.4 Conductive Hydrogels
12.2.2.5 Magnetic Scaffolds and Nanoparticles
12.2.2.6 ECM-Derived Scaffolds
12.3 Biomaterials for Pancreatic Tissue Engineering
12.3.1 Biomaterials in Restoring Pancreatic Function
12.3.1.1 Biological Polymer Scaffolds
12.3.1.2 Synthetic Polymer Scaffolds
12.3.1.3 Silk Fibroin
12.3.2 Decellularized Pancreas as Native ECM Scaffold
12.3.3 Surface Engineering of the Pancreatic Islets
12.4 Future Perspectives
References
13: Biomaterials and Stem Cells in Tissue Engineering and Regenerative Medicine: Concepts, Methods, and Applications
13.1 Introduction
13.1.1 Biomaterials
13.1.2 Stem Cells
13.1.3 Concept of Stem Cell
13.1.4 Different Types of Stem Cells
13.1.5 Tissue Engineering and Regenerative Medicine
13.2 Biomaterials and Stem Cells in TE and RM
13.3 Applications of Biomaterials and Stem Cells in TE and RM
13.3.1 Stem Cells and Biomaterials in Bone Tissue Engineering
13.3.2 Stem Cells and Biomaterials in Cardiovascular TE and RM
13.3.3 Stem Cells and Biomaterials in Pancreatic Tissue Engineering
13.3.4 Stem Cells and Biomaterials in Nerve TE
13.3.5 3D Bioprinting and Stem Cells in TE
13.4 Conclusion
13.5 Future Prospects
References
Part IV: Advances in Biomaterials
14: Biomaterials in Tissue Engineering and Regenerative Medicine: In Vitro Disease Models and Advances in Gene-Based Therapies
14.1 Introduction
14.2 In Vitro Disease Models
14.2.1 Different In Vitro Disease Models Used in TE andRM
14.2.1.1 Primary Skin Fibroblasts as a Model of Parkinson´s Disease
14.2.1.1.1 Advantage of Skin Fibroblasts as an In Vitro Model of PD
14.2.2 In Vitro Model Study of Fibroblast Activation Using Hydrogel Scaffolds
14.2.3 Induced Pluripotent Stem Cells as In Vitro Disease Models
14.2.4 Human Mesenchymal Stem Cells as In Vitro Disease Models
14.2.5 Progress in In Vitro Disease Models
14.3 Gene Therapy and Its Applications
14.3.1 Applications
14.3.2 GT in Tissue Engineering and Regenerative Medicine
14.3.2.1 Heart Diseases
14.3.2.2 Lungs Diseases
14.3.2.3 Liver Diseases
14.3.2.4 Kidney Diseases
14.3.2.5 Brain Diseases
14.4 Advances in Gene-Based Therapies and Its Applications in TE and RM
14.4.1 Adenovirus as Gene Therapy Vectors
14.4.1.1 Adenovirus Based Therapy Using Antisense/Small Interfering RNA
14.4.1.2 Cancer Vaccines Based on Adenoviruses
14.4.1.3 Gene Therapy: Applications with Haematopoietic Stem Cells
14.4.1.4 Gene Therapy and CAR-T
14.4.1.5 Gene Therapy in the Treatment of Adult-Onset Glaucoma
14.5 Biomaterials in TE Based on GT
14.6 Challenges and Future Prospects
References
15: Nanobiomaterials in Tissue Engineering and Regenerative Medicine: Current Landscape and Future Prospects
15.1 Introduction
15.1.1 Bio and Immuno Compatibility of Nanobiomaterials
15.2 Nanobiomaterials in Tissue Engineering and Regenerative Medicine
15.2.1 Neural Tissue Engineering
15.2.1.1 Types of Nano-Based Scaffolds Used in NTE
15.3 Nanobiomaterials and Bone Tissue Engineering/Regeneration
15.3.1 Nanobiomaterials Used in BTE
15.3.2 Nanohydroxyapatite (nHA)
15.3.2.1 nHA in Stem Cell Differentiation During Bone TE
15.3.2.2 nHA in Skeletal Defects Restoration
15.3.2.3 nHA in Internal Fixation
15.3.2.4 nHA in Spinal Fusion
15.3.3 Nanostructured Calcium Phosphate (CaP)
15.3.4 Graphene Nanobiomaterials
15.3.5 Titanium Nanobiomaterials
15.3.6 Silica Nanobiomaterials
15.3.7 Bioactive Glass Nanobiomaterials
15.4 Nanobiomaterials in Tissue Engineering of Bone Associated Tissues
15.4.1 Craniofacial and Dental Tissue Engineering
15.4.1.1 nHA in Dental Restoration
15.4.1.2 Nano-Titanium in Dental Regeneration
15.4.1.3 Synthetic Silicate Nanoparticles in Dentistry
15.4.1.4 Graphene in Craniofacial Bone Tissue Engineering
15.4.2 Cartilage Tissue Regeneration (Temporomandibular Joint)
15.5 Nanobiomaterials in Corneal Tissue Engineering
15.5.1 Natural Polymers
15.5.2 Synthetic Polymers
15.5.3 Nanobiomaterials in Corneal Epithelial Tissue Engineering
15.5.4 Nanobiomaterials in Corneal Endothelial Tissue Engineering
15.5.5 Nanobiomaterials in Corneal Stroma Tissue Engineering
15.5.6 Cell Sheet Engineering in Corneal Tissue Engineering
15.6 Limitations and Future Prospects
References
16: Intelligent Biomaterials for Tissue Engineering and Biomedical Applications: Current Landscape and Future Prospects
16.1 Introduction
16.2 Historical Account of Intelligent Biomaterials
16.3 Shape Changing Materials
16.4 Thermoresponsive Biomaterials
16.5 Photoresponsive Biomaterials
16.6 pH-Responsive Biomaterials
16.7 Magneto-Responsive Biomaterials
16.8 Electro-Responsive Biomaterials
16.9 Bio-Responsive Biomaterials
16.9.1 Enzyme-Responsive Biomaterials
16.9.2 Stress-Responsive Biomaterials
16.9.3 Immuno-Responsive Biomaterials
16.10 Other Stimuli-Responsive Biomaterials
16.11 Summary and Future Prospects
References
17: 3D Bioprinting in Tissue Engineering and Regenerative Medicine: Current Landscape and Future Prospects
17.1 Introduction
17.2 Background to 3D Bioprinting
17.2.1 Historical Account of 3D Printing/Bioprinting
17.2.2 Set Up and Work Flow of 3D Printing/Bioprinting
17.2.3 Types and Principles of 3D Bioprinting
17.2.3.1 Extrusion Bioprinting
17.2.3.2 Inkjet Bioprinting
17.2.3.3 Laser Assisted Bioprinting
17.3 Bioinks in 3D Bioprinting
17.4 Approaches in Bioprinting
17.4.1 Single Component Bioink Based Approaches
17.4.2 Multi-component Bioink Based Approaches
17.4.3 Approaches Involving Bioinks with Sacrificial Elements
17.4.4 Combinatorial Approaches in 3D Printing/Bioprinting
17.5 Summary and Future Prospects
References
Index

Citation preview

Birru Bhaskar · Parcha Sreenivasa Rao  Naresh Kasoju · Vasagiri Nagarjuna  Rama Raju Baadhe  Editors

Biomaterials in Tissue Engineering and Regenerative Medicine From Basic Concepts to State of the Art Approaches

Biomaterials in Tissue Engineering and Regenerative Medicine

Birru Bhaskar • Parcha Sreenivasa Rao • Naresh Kasoju • Vasagiri Nagarjuna • Rama Raju Baadhe Editors

Biomaterials in Tissue Engineering and Regenerative Medicine From Basic Concepts to State of the Art Approaches

Editors Birru Bhaskar Prof Brien Holden Eye Research Centre, LV Prasad Eye Institute Hyderabad, Telangana, India Naresh Kasoju Department of Bio-Medical Technology Sree Chitra Thirunal Institute for Medical Sciences Trivandrum, Kerala, India

Parcha Sreenivasa Rao Department of Biotechnology National Institute of Technology Warangal Warangal, Telangana, India Vasagiri Nagarjuna Society for Biological Chemists India Bangalore, India

Rama Raju Baadhe Department of Biotechnology National Institute of Technology Warangal Warangal, Telangana, India

ISBN 978-981-16-0001-2 ISBN 978-981-16-0002-9 https://doi.org/10.1007/978-981-16-0002-9

(eBook)

# The Editor(s) (if applicable) and The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 This work is subject to copyright. All rights are solely and exclusively licensed by the Publisher, whether the whole or part of the material is concerned, specifically the rights of translation, reprinting, reuse of illustrations, recitation, broadcasting, reproduction on microfilms or in any other physical way, and transmission or information storage and retrieval, electronic adaptation, computer software, or by similar or dissimilar methodology now known or hereafter developed. The use of general descriptive names, registered names, trademarks, service marks, etc. in this publication does not imply, even in the absence of a specific statement, that such names are exempt from the relevant protective laws and regulations and therefore free for general use. The publisher, the authors, and the editors are safe to assume that the advice and information in this book are believed to be true and accurate at the date of publication. Neither the publisher nor the authors or the editors give a warranty, expressed or implied, with respect to the material contained herein or for any errors or omissions that may have been made. The publisher remains neutral with regard to jurisdictional claims in published maps and institutional affiliations. This Springer imprint is published by the registered company Springer Nature Singapore Pte Ltd. The registered company address is: 152 Beach Road, #21-01/04 Gateway East, Singapore 189721, Singapore

Contents

Part I 1

2

Fundamentals of Biomaterials

Biomaterials, Tissue Engineering, and Regenerative Medicine: A Brief Outline . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Birru Bhaskar and Vasagiri Nagarjuna

3

Metallic Biomaterials in Tissue Engineering: Retrospect and Prospects . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Suvro Kanti Chowdhury, Vasagiri Nagarjuna, and Birru Bhaskar

19

3

Bioceramics in Tissue Engineering: Retrospect and Prospects . . . . . P. R. Harikrishna Varma and Francis Boniface Fernandez

4

Polymeric Biomaterials in Tissue Engineering: Retrospect and Prospects . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Lynda Velutheril Thomas

5

61

89

Composite Biomaterials in Tissue Engineering: Retrospective and Prospects . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 119 Charu Khanna, Mahesh Kumar Sah, and Bableen Flora

Part II

Trends in Biomaterials

6

Trends in Bio-Derived Biomaterials in Tissue Engineering . . . . . . . 163 Dimple Chouhan, Sharbani Kaushik, and Deepika Arora

7

Trends in Functional Biomaterials in Tissue Engineering and Regenerative Medicine . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 215 Deepika Arora, Prerna Pant, and Pradeep Kumar Sharma

8

Trends in Bioactive Biomaterials in Tissue Engineering and Regenerative Medicine . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 271 G. P. Rajalekshmy and M. R. Rekha

9

Trends in Stimuli Responsive Biomaterials in Tissue Engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 305 Rajiv Borah, Jnanendra Upadhyay, and Birru Bhaskar v

vi

Contents

Part III

Applications of Biomaterials

10

Biomaterials for Hard Tissue Engineering: Concepts, Methods, and Applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 347 Manju Saraswathy, Venkateshwaran Krishnaswami, and Deepu Damodharan Ragini

11

Biomaterials for Soft Tissue Engineering: Concepts, Methods, and Applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 381 Chelladurai Karthikeyan Balavigneswaran and Vignesh Muthuvijayan

12

Biomaterials for Specialized Tissue Engineering: Concepts, Methods, and Applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 423 Divya Sree Kolla and Bhavani S. Kowtharapu

13

Biomaterials and Stem Cells in Tissue Engineering and Regenerative Medicine: Concepts, Methods, and Applications . . . . 469 Vasagiri Nagarjuna

Part IV

Advances in Biomaterials

14

Biomaterials in Tissue Engineering and Regenerative Medicine: In Vitro Disease Models and Advances in Gene-Based Therapies . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 485 Swathi Dahariya and Vasagiri Nagarjuna

15

Nanobiomaterials in Tissue Engineering and Regenerative Medicine: Current Landscape and Future Prospects . . . . . . . . . . . . 505 Nagaraju Shiga, Dumpala Nandini Reddy, Birru Bhaskar, and Vasagiri Nagarjuna

16

Intelligent Biomaterials for Tissue Engineering and Biomedical Applications: Current Landscape and Future Prospects . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 535 M. S. Anju, Deepa K. Raj, Bernadette K. Madathil, Naresh Kasoju, and P. R. Anil Kumar

17

3D Bioprinting in Tissue Engineering and Regenerative Medicine: Current Landscape and Future Prospects . . . . . . . . . . . . 561 J. Anupama Sekar, R. K. Athira, T. S. Lakshmi, Shiny Velayudhan, Anugya Bhatt, P. R. Anil Kumar, and Naresh Kasoju

Index . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 581

About the Editors

Birru Bhaskar works at Prof Brien Holden Eye Research Centre, LV Prasad Eye Institute, Hyderabad, India. Previously, he had worked at the Indian Institute of Technology Guwahati towards the development of novel biomaterials. He has completed his graduation in Biotechnology from the National Institute of Technology, Warangal, India, where he has developed novel biomaterials & bioreactors for tissue engineering. He was a Newton Bhabha Fellow and he had also undertaken a sabbatical in the ‘Department of Materials Science and Engineering’ at ‘The University of Sheffield’. His thrust area of research is Biomaterials and Tissue Engineering. He has several publications to his credit. Parcha Sreenivasa Rao is an Associate Professor at the Department of Biotechnology, National Institute of Technology, Warangal. He was bestowed as Associate Fellow by Telangana Academy of Sciences and honoured with ‘Established Scientist Award’ by Scientific Planet Society. Earlier, he has served as the Head of the Department of Biotechnology at the National Institute of Technology, Warangal. His research group works toward the development of biomaterial scaffolds that can be utilized in tissue engineering. He is entrusted as a scientific advisor to various biotech industries including Syngene, Lonza International, ATGC International Gland Pharma, Vcare Biolabs, and BVTL Pvt. Ltd. He is a reviewer and editorial board member of renowned international journals. He has a patent to his credit and has published more than 40 research articles in journals of international repute and has also authored 5 book chapters. Naresh Kasoju is a Scientist at the Department of Bio-Medical Technology (BMT) Sree Chitra Tirunal Institute for Medical Sciences and Technology (SCTIMST), Thiruvananthapuram, India. He has earlier obtained his Ph.D. from the Indian Institute of Technology, Guwahati, India (2012) and post-doctoral training from the Institute of Macromolecular Chemistry, Prague, Czech Republic and University of Oxford, Oxford, United Kingdom (2012–2017). His research areas are focused on developing novel biomaterial structures, unraveling molecular organization in polymeric biomaterials, and understanding cell-material interactions at the molecular

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level. Till now, he has published more than 40 research articles and is associated with many international journals as a reviewer, guest editor, and editorial board member. He is a member of the prestigious National Academy of Sciences, India and Royal Society of Biology, United Kingdom. Vasagiri Nagarjuna is an alumnus of UOH and is a member of various scientific organizations (SBCI, ISCA, etc.). He has worked as a Scientist at the NIAB, Hyderabad and as a research fellow at CCMB (CRF-NIMS) Medical Biotechnology Centre & NITW. His research interests are in cellular and molecular pathobiology, chemical biology, cancer biology, stem cells, and regenerative medicine. His research contributions include pioneering work on mitochondrial-targeted curcumin, which is highly acclaimed and cited. Rama Raju Baadhe is an Assistant Professor at the Department of Biotechnology, National Institute of Technology, Warangal. He has been honored with various awards, including the Young Scientist award from Telangana Academy of Sciences (2016), Scientific Planet Society (2015), K V Rao Scientific Society Hyderabad (2013). He is bestowed as an Associate Fellow of Telangana Academy of Sciences (TAS) in 2015. His research is focused on the development of natural biomaterial which can be utilized in tissue engineering through the biorefinery concept. He is serving as an editorial member of Journal of Enzymology and Metabolism. He has published more than 20 research articles in peer-reviewed international journals and has authored 10 book chapters.

List of Abbreviations

bFGF BMP2 ECM FDA GAG GMP HA hASC HASMC HEPM IPN P4HB PANI PCL PEEK PGA PHA PHBV PLA PLGA PMMA PPy RASMC SF TE

Basic Fibroblast Growth Factor Bone Morphogenetic Protein-2 Extracellular matrix Food and Drug Administration Glycosaminoglycan Good manufacturing practices Hyaluronic acid human adipose-derived stem cells Human aortic smooth muscle cells Humanembryonic palatal mesenchymal Interpenetrating network poly 4-hydroxybutyrate Polyaniline Polycaprolactone Poly(ether ether ketone) Polyglycolic acid Polyhydroxylalkanoate poly(3-hydroxybutyrate-co-3-hydroxyvalerate) Polylactic acid Polylactide co-glycolide Poly(methyl methacrylate) Polypyrrole Rat aortic smooth muscle cells Silk Fibroin Tissue engineering

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Part I Fundamentals of Biomaterials

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Biomaterials, Tissue Engineering, and Regenerative Medicine: A Brief Outline Birru Bhaskar and Vasagiri Nagarjuna

Abstract

The frontiers of future human health care system lies in advanced interdisciplinary areas of tissue engineering (TE) and regenerative medicine (RM). The success of TE and RM is indispensible with biomaterials. Advances in the development and fabrication of biomaterials paved the way for TE and RM to revolutionize modern medical treatment. In the present chapter we have detailed the various aspects of biomaterials, TE and RM. Glimpses of brief historical aspects of these topics are presented in this chapter along with the scope and concepts of TE, RM, and biomaterials. Keywords

Biomaterials · Tissue engineering · Regenerative medicine

1.1

Introduction

Biomaterials have emerged as an integral part of medicine today and are being used successfully for the reconstruction or repair of tissue. Initially, these materials are used as prosthetic devices in the replacement of body parts lost due to trauma, congenital disorders, or diseases. These prosthetic devices improved the quality of life in physically disabled individuals. Design and suitability are only considered in

B. Bhaskar Prof. Brien Holden Eye Research Center, LV Prasad Eye Institute, Kallam Anji Reddy Campus, Hyderabad, Telangana, India V. Nagarjuna (*) Society for Biological Chemists India, Bangalore, India # The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 B. Bhaskar et al. (eds.), Biomaterials in Tissue Engineering and Regenerative Medicine, https://doi.org/10.1007/978-981-16-0002-9_1

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the development of prosthetic devices since these devices are used for replacement or repair of tissues. The advancements in material sciences and the human expectancy for quality life standards have led the way in the direction to develop biocompatible materials for restoring the function of tissue. For example, prosthetic implants enhanced the quality of life for millions of people, bioresorbable sutures made life easy for surgeons and changed the surgical procedures and artificial grafts saved millions of lives. Further exploration of material science helped in developing biocompatible and biodegradable materials with modulated properties like mechanical strength, degradation kinetics, and structure to mimic the native tissue. Various tissues in the human body possess different physical, biological, and biochemical properties. The choice of materials and their composition, design aspects, degradation kinetics, and material–cell interaction are essential for consideration in the development of implants wherein the principles of material science, thorough characterization, and biological interactions in the human body are considered. The regulatory approvals are essential to commercialize any biomedical device or implant. The U.S. Food and Drugs Administration (FDA) has approved many implants, scaffolds, fabrication techniques, and biomaterial-based delivery systems for health care applications to this end. The advancements in the material sciences and biological research have carved the niche for various therapeutic solutions, which ignited the research towards the development of engineered grafts and biomedical devices for improved health care standards and quality of human life. This is a multidisciplinary area of research comprised of material scientist, chemist, biologist, physicians, and biomedical engineers. Rapid progress in the biomaterialbased health care solutions is the need of the hour and herein through this book we summarize the concept of biomaterials, types of biomaterials, their application in tissue engineering and regenerative medicine, advances in fabrication techniques, and recent trends in biomaterials and applications. Thus, this book is aimed at providing the insights, from basics to advances, significant outcomes, of biomaterials in tissue engineering and regenerative medicine.

1.1.1

Biomaterials

Biomaterials are substances/materials which possess the properties or engineered with properties to interact with biological systems to serve various purposes. These are the materials derived either from synthetic or natural route and are used in health care to restore or repair the body part or tissue and in medical applications, wherein the principle involved in these materials development is to allow these materials to interact with biological interface systems. At the same time, these materials can be used in drug or bio-factor delivery applications. The design and development of these biomaterials rely on the need of application. In ancient times, gold was used for dental applications. Considering the tooth structure for uncompromised chewing ability and enhanced bone integration, the use of dental implants made up of sea shells was introduced by the Mayans.

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Biomaterials need to possess certain properties for them to make up for ideal scaffold materials. These properties possessed by biomaterials help in readily supporting tissue formation through cell adhesion, colonization, proliferation, and transmission of physical and chemical cues that biomimetic natural environment. Some of the ideal properties of biomaterials are discussed below. Biocompatibility Biocompatibility is the property of the material to perform specific function without eliciting harmful immune or inflammatory reactions at the desired site. In TE, as all the cellular functions such as adhesion, proliferation, migration, and differentiation occur within the scaffold, biocompatibility of the scaffold is very important to obtain the desired outcome. Various factors like material of the polymer used, its chemical and structural aspects upon functionalization determine the biocompatibility of scaffolds including choice of polymers, structure, and chemistry used for functionalization (Asghari et al. 2017). Biodegradability Biodegradability refers to the degradation of the biomaterial over time. For an ideal scaffold the rate of degradation of the scaffold should match the rate of tissue regeneration. Also, the biodegradable materials must assisting the healing and regeneration of the concerned tissue while undergoing degradation. The primary parameters that should be ensured in the selection of biodegradable materials are biosafety and physical and chemical nature of biodegraded byproducts. Mechanical Properties Mechanical properties like tensile strength, elasticity, etc. should be considered before selecting a biomaterial for scaffold preparations. Various TE applications need scaffolds of varied mechanical properties to ideally fit their purpose. Structural Properties Structural properties such as size, shape, etc. play a vital role for the usage of biomaterials in various applications. An ideal scaffold biomaterial is one which has high surface to volume ratio so as to enable cell attachment and drug delivery. Porosity Porosity of the scaffold is of utmost importance as it determines the migration of cells, supply of biochemical factors, etc., and also determines the flexibility and shape of scaffold. Scaffolds with high degree of porosity and interconnected pore networks are ideal for integration with host tissue. Processability Processability refers the property of biomaterial to be able to get fabricated into desired shape, ease of handling, etc. Stability This is another important aspect in biomaterial selection. Biomaterials should be stable at physiological conditions with respect to the application of their usage. The physical and chemical structure and properties and also the biological activity of the biomaterials should be stable.

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Fig. 1.1 Historical evaluation of biomaterials for medical applications

The evolution of biomaterials and the intrigued properties considered in development of potential biomaterials for clinical applications since 1940 to till date is depicted in Fig. 1.1. These first generation of biomaterials considered only the basic concepts of biocompatibility and functionality. The gained experience in using these materials in medical applications and pooled knowledge from multidisciplinary fields realized the necessity of introducing additional features to produce potential functional materials for health care application. The understanding of fundamental principles and integration of technical insights of various subjects which includes materials science, biology chemistry, mechanical, electrical, chemical, and biomedical engineering and medicine can bring the materials into real time medical applications. The second-generation biomaterials were developed through the incorporation of bioresorbable property in addition to the bio-inertness and biocompatibility. The concept of biomaterial evolved only to augment, replace, or repair any body part, organ, or tissue that has been lost due to injury, trauma, or disease. The clinical practice for replacement of tissue using autograft, replacing tissue from the patient own body has limited application in medicine. The limited availability, donor site morbidity, and the need of second surgery insisted further to use allograft. Allograft is the transplantation of organ or tissue from a donor to the recipient. Other potential alternative is xenograft, wherein the organ or tissue graft was taken from the donor of other species. The immunological reactions associated with the use of allograft and xenograft in patient’s body and the adverse reactions question the biocompatibility of the graft. The limitations of these grafts have drawn the attention towards biomaterials-natural or synthetic for the engineered grafts. The application of the biomaterials in health care has geared up in the twentieth century. In

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third-generation biomaterials, the biomaterial interaction with the cells, importance of physical and biochemical cues in stimulation of cell response, and immunogenic reactions were considered for use as a potential biomaterial in medical application. The present goal of scientist is to explore the development of biomimetic materials, which comes under the class of fourth generation biomaterials. At the end of World War I, physicians used the implants made up of metals, ceramics, and polymeric materials, which are inert, and durable for use to replace the diseased or damaged parts of the body. The materials used for the manufacture of automobiles, radios, and clocks were taken by the surgeons for implantation. These materials include methacrylates, Teflon, titanium, polyurethane, nylon, silicone, and stainless steel. The historical evolution of biomaterials for clinical applications till today finds its appreciation in the contributions of physicians and dental practitioners. Other than dental implants, arthroplasty, heart valves, and intraocular lenses were successfully developed using biomaterials. Later, the complexity of biological interactions of the materials and their integration with host tissue contributed for the next generation of biomaterials in the year 1960s. In addition to the biocompatibility, bioresorbable nature was introduced for the use of material in medical applications. Synthetic materials having bioresorbable nature and biocompatibility were used for medical applications. Bioresorbable nylon or silk-based sutures were developed and successfully used in surgeries. Synthetic and natural materials have been used for development of biocompatible and bioresorbable grafts that are being successfully used in clinical applications than auto or allografts. Biomaterials are classified into four categories at large. They are (1) metals, (2) ceramics, (3) polymers, and (4) composites. These materials are used for hard and soft tissue engineering applications. The first part of this book deals with the metallic, polymeric, ceramic, and composite biomaterials in tissue engineering. They have been detailed from Chaps. 2, 3, 4, and 5. In particular, recent trends and historical evolution of these materials for medical applications are briefly provided in separate chapters. The need of implantation, biomaterial intervention, and the huge demand for organ transplantation has directed the research towards organ development. Thus, Tissue Engineering has emerged for demonstrating the technological potential in organ development using biomaterials.

1.1.2

Tissue Engineering

Generation of artificial tissues, organs, has been a matter of myth and dream throughout the history of mankind until last few decades. With the introduction of clinical medicine this vision became feasible. Tissue engineering and regenerative medicine are the two terms in the field of biomedicine which deals with the transformation of these fundamental ideas to practical approaches. In the sense of modern dental implantology, the first attempt to replace teeth dates back to as early as Galileo-Roman period. The anthroposophic evidence that metallic implants found in the jaw of human skull further supports the early attempts of humans in regaining the lost function by tissue substitution (Crubézy et al. 1998).

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There were evidences of using reconstructive medicine clinically even before this from the work of Ambroise Pare` (1510–1590) described in Dix livres de la chirurgie (Parey 1649) for to reconstruct teeth, noses, and other parts of the body. Homologous transplantation of teeth was a common method in the eighteenth century to replace teeth in humans. The basis for scientific approach on transplantation medicine was pioneered by John Hunter (1728–1793) with his investigations on homologous transplantation of teeth clinically and by working on the fate of transplants experimenting in animal (Hunter 1771). Skin grafting techniques are the tissue-based therapies developed first and the use of skin graft was a milestone in the modern view of tissue engineering. Johann Friedrich Dieffenbach (1792–1847), a famous surgeon, performed animal experimental and clinical work on skin transplantation which was described in Nonnulla de Regeneratione et Transplantatione (Dieffenbach 1822). Dieffenbach is one of the early practitioner of transplantation medicine and the modern founder of plastic and reconstructive surgery. The first successful autologous skin transplantation by Heinrich Christian Bünger made breakthroughs in the clinical use of skin grafts (Bunger 1823); other contributions include use of small graft islets by Jaques Reverdin (1842–1929) and the use of split thickness grafts by Karl Thiersch (1827–1895) (Mangoldt 1895). Allograft skin banking came into existence with the advent of techniques which enabled to preserve cells and tissues making these skin grafts an off-the-shelf product. The first tissue-engineered skin products that were successful grafted are made in the late 1970s and early 1980s. Apparently this started the era of modern tissue engineering although the term “tissue engineering” was coined later in 1987. It was defined as “Tissue engineering is the application of the principles and methods of engineering and life sciences toward the fundamental understanding of structure-function relationships in normal and pathologic mammalian tissue and the development of biological substitutes to restore, maintain, or improve function.” Rudolf Virchow (1821–1902) work on biological mechanisms describing that cell proliferation is vital for tissue regeneration which decides the fate of transplants in his “Cellular pathologie” was fundamental in the investigation of tissue healing through cellular effects and cultivation of cells outside the body (Virchow 1858). The first researchers who attempted to cultivate cells outside the body are C.A. Ljunggren and J. Jolly (Ljunggren 1898). R.G. Harrison (1870–1959) made a breakthrough in in vitro cell cultivation by demonstrating active growth of cells in culture (Harrison 1910) which has realized the idea of TE to a great extent and expanded its scope in regeneration of tissues and organs in lab. In the year 1991 the first recorded use of the term tissue engineering was found to be mentioned in an article entitled, “Functional Organ Replacement: The New Tissue Engineering” Volume 12, Number 5, 2006 # Mary Ann Liebert, Inc. History of Tissue Engineering and A Glimpse Into Its Future by CHARLES A. VACANTI, M.D. published in Surgical Technology International. Dr. Joseph Vacanti from Boston Children’s Hospital and Dr. Robert Langer from M.I.T. in their article in Science (Langer and Vacanti 1993), described this new technology, as the beginning of a new biomedical discipline.

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TE has its own limitations in human applications in spite of its significant progress in research. The important problems associated with implantation are “scaling up” and cell death. For TE to be effective it must generate relatively large volumes of tissue from very few cells. But to generate a small volume of tissue large numbers of cells are needed. Also hypoxic environment leading to cell death can occur due to cell implantation and its associated vascular disruption. Since mature cells require relatively high oxygen and have a low potential for expansion (scale up) they lose efficacy when expanded in vitro. To overcome these immature cells, commonly referred to as stem cells or progenitor cells are being explored along with their potential to expanded in vitro and survive a relatively hostile environment at the time of implantation. For the purpose of TE cells are mostly derived from either the donor tissue, which are often very limited in supply or from stem cells/progenitor cells. The properties of stem cells that make them highly acceptable for use are their high proliferative capacity and pluripotency, the ability to differentiate into cells of multiple lineages. Although there are ethical concerns in the usage of human embryonic stem (ES) cells that impedes their use significantly, use of induced pluripotent stem cells (iPS cells), adult stem cells, and stem cells derived from placental and umbilical sources have replaced the ES cells as feasible sources. One of the critical factors that has to be taken into account for the effective outcome in TE is the cellular microenvironment which allows cells to function as they do in the native tissue. This can be achieved by using appropriate materials with requisite mechanical and chemical properties that can biomimetic the in vivo settings. Cell scaffolds have to serve at least one of the following purposes: 1. cell adhesion and migration; 2. retention of biochemical factors and their presentation; 3. porous microenvironment for adequate diffusion of cells, nutrients, expressed products, and waste; 4. mechanical strength; rigidity or flexibility. Tissue Engineering Society (TES) was founded by Drs. Charles A. and Joseph P. Vacanti in Boston in late 1994 and was officially incorporated in the state of Massachusetts on January 8, 1996 which later made its way for the international Tissue Engineering Society, TESi. TESi was renamed the Tissue Engineering Regenerative Medicine International Society (TERMIS), to reflect the evolution of the TE, which had expanded to include regenerative medicine. The aim of TE is to produce a functional engineered tissue or organ to restore, repair, or maintain the diseased or damaged tissue or organ. The scaffold matrices integrate and interact with the cells/tissue under the provision of appropriate physical and biological cues. It is essential to provide bioactive molecules and maintain the physical cues such as mechanical stimuli and electrical stimuli for promoting the cell proliferation towards tissue formation. The scaffolds using synthetic or natural materials were fabricated using various techniques. The structure, porosity, mechanical strength, and degradation kinetics are considered while fabricating the scaffold

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Fig. 1.2 Tissue Engineering principles contributes for the development of various engineered grafts. MSCs mesenchymal stem cells, iPSCs induced pluripotent stem cells, IVD intervertebral disc

and these parameters are tissue specific. The fabrication technique and material composition for preparation of three-dimensional (3D) scaffold are depending on the host tissue properties. Cell-biomaterial interaction is a crucial factor in the development of engineered tissue, which is depend on the physical and biological cues. The artificial tissue or organ developed using TE approach can be used in regenerative medicine, pharmaceuticals, diagnostics, and for understanding the molecular mechanism involved in disease onset and progression. In vitro disease models can be developed using TE for drug screening and evaluation as these models address the ethical and economical concerns. TE had made great advances in providing health care solutions for repair of organ or tissues including bone, cartilage, cardiac, vascular, and pancreatic tissues. The ability of stem cells to differentiate into any cell type has boosted the research to address the disease or organ failure of all the tissues or organ in human body, wherein the biomaterial scaffold, differentiation medium, growth factors, and physical cues were chosen specific to the desired tissue or organ to be developed. The combination of cells, bioactive molecules, and the mechanical cues for the development of biomaterial scaffolds can be used for various TE applications (Fig. 1.2). Overall, TE is broadly classified into two types: hard and soft TE, wherein hard TE represents the tissues in the human body that possess firm intercellular matrix and it has mineralized tissue. In contrary, soft tissues connect and support the surrounding tissues or organs in the

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body. These include tendon, ligament, blood vessel, muscle, fat, fascia, synovial membrane, and nerves. The challenges associated in the development of TE graft are optimization of isolation, proliferation and differentiation of cells, scaffold designing along with the delivery of bioactive molecules for the provision of conducive environment and organizing the growth of 3D tissue. Stem cells harvested from the patient and used in 3D scaffold have attracted the attention of researchers as they avoid the immunogenic reaction issues. A number of biomaterials, both natural and synthetic, are being developed for TE applications. Since more focus of research has shifted towards the use of natural biomaterials for TE applications, we have dedicated one chapter on detailing the bioderived biomaterials in TE (Chap. 6) in this book. The main prerequisite for biomaterial scaffold for use in TE is to possess the biodegradable nature. The improved cell–biomaterial interactions and enhanced cellular attachment to the scaffolding surface are quite essential for the development of engineered tissue using TE. Therefore, research was focused on the surface modifications of biomaterials to provide conducive environment for cell adhesion, cell proliferation and colonization, and differentiation. Chap. 7 of this book, Trends in Functional Biomaterials in Tissue Engineering and Regenerative medicine, has detailed the concepts and methods involved in surface/bulk modifications to make the biomaterials into functional biomaterials for TE. Although surface modifications favor the cell adhesion and proliferation, it is essential to provide the bioactive molecules such as growth factors or drug loaded scaffolds to induce the regeneration ability and promote the tissue growth. Bioactive molecule delivery using nanoparticles, and composite biomaterials loaded with bioactive molecule have been considered as potential strategy in TE. There is rapid increase in use of the bioactive biomaterials in the development of various engineered grafts and regeneration of tissue or organ, which has been discussed in detail in the Chap. 8. The mechanical, electrical, temperature, and pH switchable materials possessed the ability to provide the stimuli response, that is, quite essential for cell–biomaterial interactions, accelerate the cell proliferation, and effectively induce the differentiation of stem cells into specific cell lineage. Stimuli responsive biomaterials are being developed by several researchers intend to mimic the stimuli presented in the in vivo tissue environment, which could exacerbate the development of engineered tissues at in vitro level to meet clinical demand of artificial grafts. The current trend on the development of variety stimuli responsive biomaterials which are being used for nontherapeutic approach includes bio and non-biomoieties must be well explored further, warranting the need of providing trends in stimuli responsive biomaterials in tissue engineering, and detailed information on which is discussed in Chap. 9. The biomaterials developed so far have been investigated for their clinical applications, challenges associated in the implantation, subsequent modifications in making them suitable for implantation have recorded vast knowledge and the studies in the same direction are being conducted with aim of providing the engineered tissues or organs. The applications of the biomaterials have been already explored vastly for tissue engineering applications and the concepts and protocols were wellestablished. We wish to provide all the information pertained to concepts, methods

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involved in the development of biomaterials specific to the tissue requirements and their application for hard, soft, and specialized tissue engineering in this book. The tailored mechanical strength, degradation kinetics, and porous structure could be achieved using composite biomaterials. These properties were adequately tailored to meet the requirement of hard and soft tissue engineering applications and even for various fabrication techniques that were employed for the development of the scaffolds. The advances in the field of biomaterial scaffold fabrication for the hard and soft TE have been explained in detail including the concepts, methods, and applications, which provides the overall picture of biomaterials and their real time applications in clinical use at one glance in this book, in Chap. 10 which is dedicated to hard tissue engineering includes bone and dental tissue engineering, whereas Chap. 11 dealt with soft TE which includes skin, tendon, muscle, articular cartilage, fascia, intervertebral disc, synovium, joint capsule, and blood vessels tissue engineering. Nerve and pancreas TE were explored to address the clinical problems exerted by these sensory and endocrine tissues. The use of biomaterials in nerve and pancreas TE, especially the development of smart materials for providing conducive cues for regeneration makes an important aspect of research in TE. The established concepts, methods, and application of these biomaterials in specialized TE intend to provide the progress in pancreas and nerve TE are provided in Chap. 12. Overall, combination of biomaterials and stem cells are the foundation for TE, wherein variety of intrinsic properties tissues have considered to promote cell–material interaction, proliferation, and differentiation. The added information on the concepts, methods, and application of biomaterials in TE may not provide exactly how stem cells isolated form different sources can be used in a combination with variety of biomaterials. Cell–biomaterial interactions, fundamental concepts on stem cells and biomaterials, physical, chemical, and biological properties play a key role in the use of biomaterials and stem cells for artificial tissue or organ development. A notable advancement of the use of nano-based biomaterials and stem cells is in regenerative therapeutics, wherein mostly novel delivery systems which have been explored to improve the on-site, targeted regeneration are of importance. In Chap. 13, biomaterials and stem cells in tissue engineering and regenerative medicine—Concepts, Methods, and Applications of these are presented. The regenerative medicine aspects using functional biomaterials, bioactive biomaterials along with the overall view of biomaterials and stem cells have been covered in first three parts of the book. The fundamentals, trends in biomaterials, and application of biomaterials have been equally given priority in this book to provide outstanding, emerging concepts in the field of biomaterials, and TE with aim of giving substantial information and resources to the material scientist, biomedical engineers, and clinicians helps in establishing bridge connection among them, which foster the new innovations to meet the current global demand of artificial organ or tissue.

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Regenerative Medicine

Regenerative medicine (RM) is an interdisciplinary field of research and clinical applications which lay emphasis on the repair/replacement and regeneration of cells, tissues or organs to restore their impaired function resulting from any cause, including congenital defects, disease, trauma, and aging. In short, Regenerative medicine (RM) implies the replacement or regeneration of human cells, tissue, or organs, to restore or establish normal function [C. Mason]. The term “regenerative medicine” is widely considered to be coined by William Haseltine during a 1999 conference in Lake Como, in the attempt to describe an emerging field, which blends knowledge deriving from different subjects: tissue engineering (TE), cell transplantation, stem cell biology, biomechanics prosthetics, nanotechnology, and biochemistry (Mason and Dunnill 2008). Historically, this term was found for the first time in a 1992 paper by Leland Kaiser, who listed the technologies which would impact the future of hospitals (Kaiser 1992). The potential outcomes of medical interventions of RM were recognized by mankind way before the term “Regenerative Medicine” was coined. Discoveries and techniques in medicine pioneered by ancient civilizations like Indian, Chinese, South American, Egyptian, and Sumerian have had their impact on the field still today. One of the common phenomenon of leaving beings is regeneration of body parts which was noticed and studied since ages in organisms like salamanders which can regenerate an amputated limb in a few days (Kragl et al. 2009). Though humans also have the potential to regenerate a served fingertip, their regenerative capacity decreases with age and is limited up to the age of 11 years (Illingworth 1974). Ancient mythologies have mentioned the regeneration of human liver of Prometheus and it was noted in literature that ancient Greeks had the knowledge regarding the regenerating capacity of liver. They termed it “hepar” (Greek: meaning to “repair oneself”). Aristotle in three of his works on natural history, History of Animals, Generation of Animals and Parts of Animals has described that undifferentiated matter was able to give rise to life. The Aristotelean thesis stated that animals possess higher regenerative potential in their early stages of development. He had detailed the process of regeneration on the limbs of salamanders and deer antlers (Barnes 1984). His observations on the origin of biological forms from undifferentiated matter clearly emphasized the concept of “epigenesis.” Later, William Harvey (1578–1657) coined the word “epigenesis” in his work “on the generation of animals” which was a repetition of Aristotle’s works. Abraham Trembley (1710–1784) produced several publications on the regenerative phenomena on freshwater polyps (Vartanian 1950). It was also established from the data of more than 10,000 mummies that cranial trepanation was a routine procedure in prehistoric Peru as early as 3000 B.C. Only after the advent of cell therapy regenerative medicine became a tangible area of science. First therapeutic surgeries in medicine were a result of work done in the field of transplantation in the mid-1950s. Some of the early transplants that made key milestones in the history of transplantation were the first kidney transplant in 1954

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followed by the first liver and lung transplants in 1963, pancreas transplant in 1966, and the first heart transplant in 1967. Later the real excitement came with the breakthrough in bone marrow transplantation in leukemia patients that changed the course of history. This lead to a wave of enthusiasm with cell biologists questioning the capabilities of the integrity of the tissues being transplanted and working on the possibilities to create, grow, and harvest these tissues in the laboratory. Thus the modern era of Tissue Engineering began leading the way for developments in the field of Regenerative Medicine (George et al. 2017). Inception of cyclosporine into medicine and its first successful use in kidney recipients in late 1970s lead the way to cyclosporine era in 1980s for use in solid organ transplantation reducing the risk of rejection drastically which was otherwise before due to adverse immunological effects. At present the major challenges in transplantations are an ever increasing demand for organ transplantation and lifelong immune suppression which leads to a number of side effects. With increase in life expectancy and an aging population with the need for transplantation of diseased/ injured organs due to age related issues coupled severe shortage of donors the problem is ever increasing. Traditional treatment of diseased organs by surgery faces three major problems, i.e., repair of autologous tissues, resection of lesions, and replacement with allografts known as three “R” paradigm. Revolutionize modern medicine and overcoming these issues the fourth “R- Regenerative medicine” came into limelight offering ways to cure and regenerate damaged/diseased tissues. RM is a combination of several technological approaches including TE, stem cell therapy, gene therapy, etc. which moves beyond traditional transplantation therapies addressing various problems that are clueless before. The possible health care associated infection during the organ transplantation and surgical intervention for engineered tissue replacement pushed the search for alternative treatment approach. The use of immune suppression drugs and antibiotics after transplantation attributes for several side effects. The reprogrammed cells using genetic engineering tools, cell therapies, and gene therapies have been developed to regenerate the tissue and restore the normal function of tissue or organ. This classical approach, providing non-invasive procedure, and effective treatment strategy was simply termed as RM. The central and fundamental focus of RM to be productive and promising is in the selection of human cells. These cells could be embryonic derived, adult or somatic cells, and reprogrammed cells (iPSCs). The moral and ethical issues bounded with use of embryonic derived cells have attracted the attention towards iPSCs and somatic cells. The advances made in genetic engineering and molecular biology help in introducing genes into human cells for the treatment of many disorders, and improved function of tissues or organs. The use of RM in the health care would cut down the medical prosthesis and pharmaceuticals and largely mitigate the patient economic burden with the provision of effective treatment with better compliance. For example, β-islet cells derived from stem cells can be used to replace the need of insulin injection in diabetic patients. The autologous chondrocytes—Carticel were first FDA approved product for the management of cartilage defects. Herein, the

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autologous chondrocytes have been expanded in ex vivo for implantation at the site of injury, and these chondrocytes are harvested from articular cartilage. The regenerative medicine often uses the materials in current practices owing to their ability to mimic extracellular matrix (ECM) of native tissue and favors the structure and function of neo tissue with the local administration of growth factors. Polymer based 3D scaffolds have been investigated for their potential use in cartilage regeneration, for example MACI (matrix-induced autologous chondrocyte implantation) are being successfully used for cartilage regeneration. The human fibroblast derived artificial graft—“Derma graft” has been widely used for the treatment of chronic wounds and venous ulcers. The concepts evolved to use the decellularized cells for regenerative medicine have geared up the clinical investigations to promote healing or substitute the tissue. In some cases, material alone can induce the regeneration, acts as graft substitute, favors host integration. Several bone grafts help in host integration and fusion with the bone. The local and sustained growth factor delivery has augmented to use the biomaterials for promoting regeneration and healing of the tissue. Bone morphogenic proteins (BMP-2 and 7) have been investigated for their potential ability of bone regeneration using biomaterials. Infuse, Stryker’s OP-1 are FDA approved products aimed to deliver the BMP-2 and 7 growth factors at the site of injury and help in the bone regeneration. Notable development is made in health care research using regenerative medicine to bring the most efficient treatment modalities. Several studies are in their clinical investigative stages as huge regulatory procedures should be followed for getting the FDA approval. Most likely, tissue engineering and regenerative medicine are interchanging the principles, concepts, and methods, at the same time, aimed to deal the repair or replacement of tissue or organ. The advances in biomaterials for the tissue engineering and regenerative medicine aimed to develop the engineered grafts or regenerative therapies, wherein the improved host tissue integration, addressing the immunological reactions, and neo-vascularization and innervation are the key considerations. Biomaterial-based 3D disease models are being widely preferred for drug screening approaches and development of in vitro models helps in understanding the molecular mechanism involved in disease progression and diagnosis. The rapid screening of drug with reduced cost and time, and enhanced disease diagnosis can be achieved using 3D disease models. Gene impregnated biomaterial scaffolds are the novel strategy in promoting the healing and regeneration of tissue, which comes under the class of gene therapy. Biomaterials play a significant role in the gene delivery, herein conducive cues provided by the biomaterials favor the cell adhesion and gene fusion in the cells augment the healing and regeneration of tissue. The targeted, localized, and sustained gene delivery approach is essential to improve the efficacy of the treatment, which boosted the research towards the biomaterial advancement. The 14th chapter “Biomaterials in tissue engineering and regenerative medicine- In vitro disease models and advances in gene-based therapies”, details the potential use of genetic engineering in tissue regeneration that could advance the development of cost-effective drug screening models, improved regenerative therapy in addressing the graft failures at large. Nano scale biomaterials have emerged to mimic native structure of tissue, improve the cell adhesion, promote tissue

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regeneration, and help in host tissue integration with the goal of developing artificial grafts for successful transplantation. Nanoscale biomaterials have been developed using various fabrication techniques to aptly mimic the native tissue structure and are successfully being assessed for their potential application in tissue engineering and regenerative medicine. This has been discussed in Chap. 15. The conducive cues for promoting healing or regeneration of tissues are essential and the advancement in the field of tissue engineering has pooled all the failures, logistics in the biomaterials use without incorporating variety of stimuli responses. Simple scaffolds only can provide adequate physical and mechanical properties, which alone could not help in the functional tissue formation. The progress in the biology and understanding the native tissue environment, stimuli response role in promoting the tissue regeneration with the enhanced cell-biomaterial interactions have augmented the research towards the development of intelligent biomaterials for TE and RM. The intelligent biomaterials possessing biomimetic nature and tailored properties according to the requirement of stimuli response promote the desired functional tissue regeneration. Numerous studies are being conducted on the development of intelligent biomaterials and their application for TE and RM, which are discussed in detail in the Chap. 16. The anatomical structure of tissue is very important in developing engineered organ or tissue, and traditional methods for scaffold fabrication cannot mimic 3D structure. 3D printing has been used successfully for organ development, especially multilayered scaffolds with cell laden constructs are being developed for the organ development. Bioinks are being developed for bioprinting of biomaterials, this bioprinting approach would make it most reliable, precise 3D organs and rapid production is possible with this technology. The advances in biomaterials led to the preparation of printable bioinks and their application enormously has been geared up in TE and RM. The 3D bioprinting of biomaterials for TE and RM: Current landscape and future prospects have elaborately discussed in Chap. 17. Thus, this book as a whole provides all the basic to advanced, essential topics in the field of biomaterial development with respect to their application and role in TE and RM. Through this chapter we tried to provide insights into the basic concepts of biomaterials, TE and RM and the different aspects presented in this book so as to get a better conceptual understanding of the subject.

References Asghari F, Samiei M, Adibkia K et al (2017) Biodegradable and biocompatible polymers for tissue engineering application: a review. Artif Cells, Nanomed Biotechnol 45:185–192. https://doi. org/10.3109/21691401.2016.1146731 Barnes J (1984) Complete works of Aristotle, volume 1: The revised Oxford translation. Princeton University Press, Princeton Bunger C (1823) Gelungener versuch einer nasenbildung aus einem vollig getrennten hautstuck aus dem beine. J Chir Augenheilk 4:569 Crubézy E, Murail P, Girard L, Bernadou JP (1998) False teeth of the Roman world. Nature 391 (6662):29. https://doi.org/10.1038/34067 Dieffenbach J (1822) Nonnulla de regeneratione et transplantatione. Richter, Herbipoli

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George J, Manjusha W, Jegan S, Mahija S (2017) A review of stem cells in regenerative medicine. Int J Sci Res Sci Technol 3(8):806–815 Harrison RG (1910) The outgrowth of the nerve fiber as a mode of protoplasmic extension. J Exp Zool 9:787–846 Hunter J (1771) Natural history of the human teeth etc, London Illingworth C (1974) Trapped fingers and amputated finger tips in children. J Pediatr Surg 9:853–858 Kaiser L (1992) The future of multihospital systems. Top Health Care Financ 18:32–45 Kragl M, Knapp D, Nacu E et al (2009) Cells keep a memory of their tissue origin during axolotl limb regeneration. Nature 460:60–65. https://doi.org/10.1038/nature08152 Langer R, Vacanti J (1993) Tissue engineering. Science 260:920–926 Ljunggren CA (1898) On the safe survival of skin epithelial cells outside of the human organ ism with special reference to skin transplantation. Nordiskt Medicinskt Arkiv 31:1–10 (in Norwegian). https://doi.org/10.1111/j.0954-6820.1898.tb00376.x Mangoldt F (1895) Die Ueberhäutung von Wundflächen und Wundhöhlen durch Epithelaussaat, eine neue Methode der Transplantation). DMW - Deutsche Medizinische Wochenschrift 21 (48):798–799 Mason C, Dunnill P (2008) A brief definition of regenerative medicine. Regen Med 3:1–5 Parey A (1649) The works of that famous Chirurgion. Clarke, London Vartanian A (1950) Trembley’s Polyp, La Mettrie, and 18th century French materialism. J Hist Ideas 11:259–286. Virchow R (1858) Die Cellularpathologie in ihrer Begründung auf physiologische und pathologische Gewebelehre. Hirschwald, Berlin

2

Metallic Biomaterials in Tissue Engineering: Retrospect and Prospects Suvro Kanti Chowdhury, Vasagiri Nagarjuna, and Birru Bhaskar

Abstract

Metallic biomaterials and their use in tissue engineering (TE) have always been the focus of study ever since research in tissue regeneration initiated. With the day-to-day emergence of tissue engineering and its applications in health care and research, development of novel biomaterials became imperative. The importance of metallic biomaterials in tissue engineering is ascribed to their exceptional amalgamation of properties like high mechanical strength, shape memory, controlled degradation, anti-microbial activity, radiopacity and their blending for preparation of different alloys and amalgams. With the advent of modern processing technologies like additive manufacturing, metallic biomaterials with controlled porosity and degradation rates have been fabricated. These fabricated biomaterials exhibited enhanced cytocompatibility and anti-corrosion activity which have been the major drawbacks for their usage in the past. New innovations in fabricating metallic biomaterials paved the way for regeneration of tissues with desired shape and size for their potential use in bone, cartilage, dental and cardiovascular tissue engineering. Our present chapter deals in detail the advent of metallic biomaterials in TE, their evolution, importance and drawbacks, properties, types and applications of metallic biomaterials in TE, along with future prospects.

S. K. Chowdhury Department of Biosciences and Bioengineering, Indian Institute of Technology Guwahati, Guwahati, India V. Nagarjuna (*) Society for Biological Chemists India, Banglore, India B. Bhaskar (*) Prof. Brien Holden Eye Research Centre, LV Prasad Eye Institute, Hyderabad, Telangana, India e-mail: [email protected] # The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 B. Bhaskar et al. (eds.), Biomaterials in Tissue Engineering and Regenerative Medicine, https://doi.org/10.1007/978-981-16-0002-9_2

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Keywords

Shape memory · Radiopacity · Additive manufacturing · Cytocompatibility · Biofunctionality

Abbreviations ACL Ag Al ALP BMG BMP2 BrdU Ca Co Cr CT Cu CVD CVI DMLS DREAMS ESR Fe FGF GAGs HA HNS HUVECs Mg Mo MRI MSC MTT Ni Pd PEG PLA Pt SLM SMA SPD Sr

Anterior cruciate ligament Silver Aluminium Alkaline phosphatase Bulk metallic glass Bone morphogenetic protein-2 Bromodeoxyuridine Calcium Cobalt Chromium Computed tomography Copper Chemical vapour deposition Chemical vapour infiltration Direct metal laser sintering Drug eluting absorbable metal scaffold Electroslag remelting Iron Fibroblast growth factor Glycosaminoglycans Hydroxyapatite High nitrogen steel Human umbilical vein endothelial cells Magnesium Molybdenum Magnetic resonance imaging Mesenchymal stem cell (3-(4,5-Dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide) Nickel Palladium Polyethylene glycol Polylactic acid Platinum Selective laser melting Shape memory alloy Severe plastic deformation Strontium

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Metallic Biomaterials in Tissue Engineering: Retrospect and Prospects

Ta TE TGF β THR Ti TiO2 TKR VEGF Zn Zr

2.1

21

Tantalum Tissue engineering Transforming growth factor-β Total hip replacement Titanium Titanium dioxide Total knee replacement Vascular endothelial growth factor Zinc Zirconium

Introduction

The exploitation of metals for biomedical applications dates back as early as the eighteenth century when they were first utilized for fixing bone fractures. Back then, only biomechanical properties of these metals are taken into consideration while implanting them, leading to several drawbacks like insufficient strength, lack of corrosion resistivity, microbial infections, etc. (Lambotte 1909; Sherman 1912). Later advances in science and technology lead to preparations of stainless steel (Hudetz et al. 2008) and other metallic biomaterials like cobalt (Co)-based alloys and titanium alloys for use as implants in replacement of ruptured hard tissues. With the evolution of engineering and material sciences, sophistication in the fabrication of metallic implants for different hard tissue applications became feasible, wherein the development of fabrication techniques so far has made it easy to tailor the shape and size as per the tissue requirements. Initially, easy fabrication of implants had geared up their usage for dentistry and orthopaedic applications. Fabrication techniques such as forging, casting and machining have been extensively used for the preparation of metallic implants (Niinomi 2008a). Major advantages of metallic biomaterials for use in TE are their inherent strength, ductility, shape memory, resistance to degradation and easy sterilization process. However, some metals have the disadvantage of possessing properties like high elastic modulus, non-bioactivity, ease of corrosion and release harmful metal ions in the body. During the design and fabrication of metallic biomaterials the ideal properties that are taken into consideration are their high mechanical strength, biocompatibility and resistance to corrosion. Also, for implantation these biomaterials must be bio-inert towards body fluids and should exhibit desirable bioactivity (Vallittu 2017). The biocompatibility of the metallic biomaterials is very important as corrosion of implant under the impact of internal body microenvironment results in loss of material, which will not only incapacitate the implant but also results in release of corrosion products which can lead to harmful effects inside the body (Wang et al. 2007; Ivanova et al. 2014). For this purpose, different alloys have been prepared for implantation with the aim of inhibiting the toxic metal ion release (e.g. Ni in Ti-Ni

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alloy). Electrical conductivity is another important property of metallic biomaterials which extend their use as neuromuscular devices and cardiac pacemakers (Prasad et al. 2017). The improvement in mechanical strength, biodegradability, biocompatibility and functionality of the metallic biomaterials brings them to the forefront of many clinical applications including dentistry, orthopaedic, cardiology and gynaecology (Bian et al. 2016). Mg- and Fe-based metallic biomaterials have exhibited biodegradability (Bian et al. 2016; Staiger et al. 2006) but have low mechanical strength, which has augmented the research towards making composites/alloy-based implants with improved mechanical strength, corrosion resistance along with the inclusion of biodegradability. High mechanical strength, fracture resistance and long durability of the implant with enhanced performance make metallic biomaterials competent for use in load-bearing tissues (Yang et al. 2017). Metallic biomaterials are widely preferred for various tissue engineering applications like bone, cartilage, dental and vascular tissues (Fig. 2.1). The inherent properties possessed by metallic biomaterials have led to their use in various soft tissue applications including artificial heart valves, vascular conduits, stents, balloon catheters and neurovascular implants (Niinomi 2008a; Ma et al. 2016). For compatibility in implant applications, mechanical properties of biomaterials must match with tissue of interest. Metal implants must integrate with host tissue along with possessing the desired mechanical properties to restore the normal function of tissue. For example, the tissue facilitating a movement in the body requires higher compressive strength; development of spine cage implants should have higher compressive strength (Matena et al. 2015) and osseointegration (Zhao et al. 2016). Another example is the use of titanium alloys for the repair of maxillofacial hard tissue for which the material must follow stress–strain behaviour and tensile strength for mandibular implantation (Wang et al. 2017b). For developing the stents, bioabsorbability and haemolytic properties of the biomaterials are to be taken into consideration. The design and fabrication of tailored mechanical properties (stiffness, stress, corrosion resistance and fatigue) play a vital role in the development of efficient metal implants (Prasad et al. 2017). Recent advances for improved functionalization of metallic biomaterials through surface treatment have emerged as a gateway for the development of next generation tissue engineering metallic scaffolds (Mani 2015; Su et al. 2017). Advances in the metallic biomaterials and their clinical applications have led to successful attempts towards the development of functionalized anti-bacterial and biodegradable implants. Release of certain ions also helps in functionality, e.g. Zn, Ca, Sr and Mg are the osteoinductive ions. Fabrication of metallic biomaterials with the amalgamation of different properties like biocompatibility, bioactivity, bio-functionalization, biodegradation and anti-bacterial properties is the need of the hour in bringing a new class of biomaterials for the development of engineered grafts. Further studies for advancement in properties can do wonders in the field of metallic biomaterials and their use in TE in future.

Metallic Biomaterials in Tissue Engineering: Retrospect and Prospects

Fig. 2.1 Representation of tissue engineering applications of metallic biomaterials

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2.1.1

S. K. Chowdhury et al.

Traditional Metallic Biomaterials

Metallic biomaterials are in use for a long time, owing to their biocompatibility and mechanical integrity. For example, in 1565, gold (Au) plate was used for the repair of cleft-palate defects (Greenspan and Hench 1976). Superior mechanical properties like fracture toughness, yield strength, fatigue strength and ductility of metallic biomaterials increased their application for load-bearing tissues, wherein these implants do not undergo deformation. Metallic biomaterials (metals and alloys) including stainless steel, carbon steels, silver (Ag), platinum (Pt), tantalum (Ta), palladium (Pd), zinc (Zn), copper (Cu), iron (Fe), magnesium (Mg), aluminium (Al), titanium and its alloys, Co-Cr alloys have been used for clinical applications in vascular therapy, trauma, orthopaedics, dental care and cardiology. Stainless steel, Ti and CoCr alloys are the majorly used metallic biomaterials. Improvements in stainless steel by addition of molybdenum (Mo) and reduction of carbon (C) increase its corrosion resistivity (Atanda et al. 2010). Titanium is known for least density compared to other metals (Smithells and Brandes 1992). Even its alloys like Ti6Al4V possess effective strength and corrosion resistance. Another alloy of titanium and nickel (nitinol) has been used for dental wiring owing to their property of shape memory. CoCr alloys have been used effectively till date for manufacturing artificial joints. Iron-based (FeMn) (Hermawan et al. 2008) and magnesium-based alloys (MgAl, MgRE, MgCa (Li et al. 2008)) are prioritized as bioimplants for their biodegradability. More focus is being laid on fabricating metallic biomaterials by blending properties like biodegradability, corrosion resistance, increased strength, good wear resistance and enhanced biocompatibility. Even biodegradable metals (Hermawan and Mantovani 2009) are being researched upon for advanced applications.

2.1.2

Advanced and Revolutionizing Metallic Biomaterials

Recent research on metallic biomaterials is directed towards the improvisation of existing properties like bio-functionality, biodegradability, inertness, biocompatibility, mechanical strength, corrosion resistance, etc. Most of the metal implants used earlier had some shortcomings on bio-functionality. The role of these revolutionized metallic biomaterials is to add up useful properties like anti-bacterial activity, enhanced osteogenesis, MRI compatibility, reduced in-stent restenosis and radiopacity to the already existing ones. Some of them have been engineered with smart coatings for mimicking bio-functional systems for the role of biosensors (Al-sehemi et al. 2017), orthopaedic (Farraro et al. 2014), cardiovascular and neural applications. Advanced metallic biomaterials are fabricated keeping in mind the anti-bacterial properties, which are absent in the earlier metal implants. To induce anti-bacterial activity in these metals, they have often been alloyed with metals like Cu, Ag, Fe, Zn, Al and Mg (Berry et al. 1992; Agarwal et al. 2010). However, one study represented that Cu, in spite of having anti-microbial effect, can be cytotoxic to

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human cells (Kishimoto et al. 1992) depending on its dosage. More research is being performed on titanium (Shirai et al. 2009; Liu et al. 2014; Kang et al. 2012) and stainless steel alloys (Hong and Koo 2005; Liao et al. 2010; Ren et al. 2011; Zhang et al. 2012) in combination with Cu or Ag for exploring their anti-bacterial activities. A great deal of research has been done in this area which is discussed herein. Layered metallic biomaterials have been used as bone implants with hydroxyapatite and other ceramics to enhance osseointegration (Habibovic et al. 2002). Studies are done on Ti-based porous biomaterials coated with calcium phosphate (CaP) to explore its effects on osteogenesis (Lopez-Heredia et al. 2008). Zinc has been effectively used as a component in scaffolds for fabricating biomimetic bone grafts (Moses et al. 2019). Porous nitinol scaffolds have also been constructed for bone tissue engineering (Gotman et al. 2013). Coronary stents which have been made using Cu bearing stainless steel to inhibit restenosis limit the proliferation of the vascular smooth muscle cells and support endothelial cells (Ren et al. 2012). Some revolutionized metallic biomaterials are also composed of Mg, Cu and Zr to enhance the MRI compatibility requirements (Li and Xu 2014).

2.1.3

Metallic Biomaterials and Biocompatibility

Biocompatibility is a vital issue that needs to be addressed as most of the metallic biomaterials used for tissue engineering are pro-inflammatory and are subjected to corrosion effects due to the physiological conditions prevailing inside the body. The interactions of the metal with the body fluids, enzymes, hormones, proteins may induce allergy, toxicity or even implant failure. Modern metal implants like 316 L stainless steel (Atanda et al. 2010) and titanium (Shirai et al. 2009) have high corrosion resistance and therefore release less ions inside the body. The formation of metal oxide films on their surfaces renders them physiologically inert (Zhen et al. 2013; Bauer et al. 2013), thereby reducing the chances of restenosis and inflammation. However, some reports (Takazawa et al. 2003; Staffolani et al. 1999) have stated that 316 L stainless steel implants cause pain as a result of nickel ion release. To address this, efforts are being made to reduce the nickel concentration from the existing 12–15% in 316 L stainless steel and adding nitrogen (Yang and Ren 2010; Mudali et al. 2002) for bolstering its corrosion resistance. It was also shown that addition of Zr, Nb and Ta to Ti-based biomaterials makes them more corrosion resistant with excellent biocompatibility (Li et al. 2010; Okazaki and Gotoh 2002). Some metals are also reported to induce allergies (Nosbaum et al. 2009; Brandão and Gontijo 2012) of type IV or delayed hypersensitivity reaction. Nickel and cobalt are mostly responsible for allergic contact dermatitis (Fernández Vozmediano and Armario Hita 2005). In vitro cytotoxicity tests have been performed over the years to check for any resulting adverse effect of metallic biomaterial implant. Often the results are found to be complicated and non-conclusive as cytotoxicity depends on the cell type chosen, culture conditions, duration of exposure to metal, concentration of metal ions, surface properties and physico-chemical parameters of the metallic biomaterial chosen (Wataha et al. 1994). Retamoso et al. (2012) found that the alloy,

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Ti–Cr–Mo–Fe, is more biocompatible with NIH/3 T3 cell lines than other cell lines. Cytotoxicity experiments by Garza-Cervantes et al. have shown that Ag when combined with Cu, Co, Ni, Zn, Cd exhibited enhanced anti-microbial potential on cell lines (Garza-Cervantes et al. 2017). The results yielded by various cytotoxic tests play a crucial role in determining the feasibility of a metal as a biomaterial for tissue engineering. As a consequence of accumulation of reactive oxygen species, metallic nanoparticles play a great regulatory role in cell signalling pathways resulting in altered gene expression. Except gold nanoparticles, most metal nanoparticles often degrade inside cells resulting in apoptosis and inhibition of cell differentiation (Lunov et al. 2010). However, some metal ions like Ca, P, Zn, Mg, Sr, Cu are found to have positive effects on osteoblast cell lines and are being looked upon as possible implant materials for orthopaedic applications, e.g. Zn induces osteogenesis, while Cu helps in angiogenesis. Some researchers found that Al promotes cell proliferation in lower concentrations (1000 μM) (Zhou et al. 2009). A comprehensive update on the various cytotoxicity analysis performed on different metallic biomaterial specimens has been stated in Table 2.1. More research needs to be conducted on each metal based on their physico-chemical properties and applications before drawing suitable conclusions.

2.2

Properties of Metallic Biomaterials

2.2.1

Phase Transformation and Elastic Moduli

Changes in the thermodynamic conditions of a metal, i.e. decrease in Gibbs free energy, result in its phase change. Rearrangement of atoms mostly occurs through a diffusive transformation process although some transformations are displacive or martensitic, where phase changes are dependent only on temperature. Such changes are visible in Ti-based and CoCrMo alloys used as joint replacement biomaterials. Ni-Ti alloys, also referred to as shape memory alloys (SMA), are found to exhibit this activity (Kaya and Kaya 2019). Biomechanical properties like strength and stiffness are an important aspect to be looked upon while designing biomaterials for tissue engineering. They should be strong enough when used in load-bearing sites and at the same time light-weighted to easily facilitate movements. Such properties are dependent on the metal microstructure and mode of processing the biomaterial. The elastic modulus of the native tissue should be mimicked by the metal used for the purpose. Elastic moduli of some tissues had been closely replicated by common metal implants as listed in Table 2.2. Large differences in it might lead to instability in the long term. Metallic biomaterials with minimal Young’s modulus that can lessen the impact of the implant on the tissue are the need of the hour.

Ag, Al, Cr + 3, Cr + 6, Ni, V

Au–Pt alloy, Co–Cr alloy, Ni–Cr alloy Ag, Cu, Zn

Cu, Al, Ti, Zr, V, Nb, Ta, Cr, Mo, Mn, Ni, Fe, Co, W Cd, Cr, Co, Fe, Mo, Ni, Ta, Ti, Zn

Metallic material Ti–Cr–Mo–Fe, Stainless steel, Ag–Sn–Co–Hg Au–Pt–Pd–Ag, Au–Pt–Pd–Ag–Cu– In–Ir, Au–Pt–Pd–Ag–In– Ru–Zn, Au–Pt–Pd–Ag–Sn– In–Ga Iron oxide

Metal ions of AgCl, AlCl3, CrCl3, CrCl2O4, NiCl2, VOCl3

Metal ions of Ag2SO4, CuCl2, ZnCl2

Metal alloy extract

Metal microparticle (0.1–150 μm) Human gingival fibroblast Primary human endometrial epithelial cells Primary human osteoblast, primary rat osteoblast

Human erythrocyte

Primary human fibroblast (h-TERT BJ1) Human osteoblastlike cell (MG-63)

Nanoparticles

Metal microparticle (1–147 μm)

BrdU assay; immunocytochemistry

Human fibroblasts

Metal block

MTT, ALP activity

MTT

MTT

Hemolysis measurement

Neutral red

MTT assay

Assay done MTT ssay

Cell treated Mouse fibroblast (NIH/3 T3)

Specimen type Metal alloy

Table 2.1 Studies reported on cytotoxicity of metallic biomaterials

Cr, Ni, V showed higher cytotoxicity when conc. was >100 μmol.L1

Uncoated nanoparticles had 40% cell viability and pullulan-coated ones had 92% cell viability Al, Ti, Zr, Nb, Ta, Cr, Mo and Fe demonstrated lower cell viability compared to cu, Si, V, W, co Co, Cr, Ni were the most toxic, whereas cd, Ti, ta, Zn showed very low levels of haemolysis Ni-Cr alloys exhibited more cytotoxicity (only 60% viability) Cytotoxicity % was in the order of ag+ > Cu2+ > Zn2+

Only Au–Pt–Pd–Ag–In–Ru–Zn and Au–Pt–Pd–Ag–Cu alloys showed higher % of BrdU positive cells (41.34
2.95, 37.61
2.42)% compared to control

Cell viability data Ti-Cr-Mo-Fe alloys had the best viability (84.92
0.11)%

Zhou et al. (2009)

Imirzalioglu et al. (2012) Wu et al. (2012)

Rae (1978)

Gupta and Curtis (2004) Sakai et al. (2002)

Grill et al. (1997)

References Retamoso et al. (2012)

2 Metallic Biomaterials in Tissue Engineering: Retrospect and Prospects 27

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Table 2.2 Mechanical properties of common metallic biomaterials Young’s modulus Application (GPa) Bone 114

Yield strength (MPa) 760–880

Elongation % 10–15

Tensile strength (MPa) 895–930

Bone

114

827–1103

–

860–965

Bone

193

170–310

–

540–1000

Bone

–

190–690

12–40

490–1350

Davis (2003)

Bone

28–41

70–140

~9

895

Ta

Bone

188

140–345

1–30

205–480

Co-Cr

Stents

200

550

3

720

Ti

Stents

100–120

310–490

10–20

380–640

Ta

Stents

61

200

>60

500

Mg-Zn

Stents

42

170

19

280

Niinomi (2008a) Niinomi (2008a) Niinomi et al. (2015a) Niinomi et al. (2015a) Niinomi et al. (2015a) Zhang et al. (2010)

Metal implants Ti6Al4V (cast) Ti6Al4V (wrought) 316 L stainless steel Austenitic stainless steel ASTM F138 Ni-Ti alloy

2.2.2

References Prasad et al. (2017) Paxton et al. (2016) Paxton et al. (2016)

Porosity

The extent of porosity on a metallic implant determines its cell adhesion, proliferation and differentiation ability. Dense metallic biomaterials lack adequate volumetric porosity which hinders their ability to adapt to the local tissue microenvironment. Surface modifications on metallic biomaterials for bone tissue engineering have found to be very beneficial in mimicking the native tissue (Habibovic et al. 2008; Shah et al. 2016b). Additive manufacturing or 3D printing had shown to help in this process of fabricating a controlled porous interconnected structure, thereby replicating the mechanical properties of the bone tissue (Shah et al. 2016b; Wang et al. 2017a). Other techniques used for creating porous metallic biomaterials are electrodeposition, alkali heat treatment and anodization (Zhang et al. 2005). Porosity aids in several ways, such as reducing the elastic modulus of the metallic biomaterial and enhancing cellular growth in the metal scaffold. Reports suggest that optimal

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pore size of 100–400 μm helps in vascularization and osteoblast growth in bone tissue (Bobyn et al. 1980). Porous titanium (Ti), porous magnesium (Mg) and porous tantalum (Ta) have been found to be more effective when compared to their dense counterparts (Mediaswanti et al. 2013). Additive manufacturing is the most widely used technique for constructing porous metallic biomaterials, especially for bone tissue engineering (Fig. 2.2). More research is being conducted to ensure that the surface modification technique used for generating a porous metal implant does not impair the mechanical and self-healing properties.

2.2.3

Corrosion Resistance

Metals that exhibit corrosive activity are not preferred for tissue engineering applications as the resulting ions may negatively impact the tissue microenvironment. These metal implants might affect the self-healing process of the body by inducing electrochemical variations resulting in alteration of equilibrium conditions and drastically reducing the pH of the body fluid. Sometimes, even H2O2 may be formed due to inflammatory reactions at the implanted site (Thomsen et al. 1991). Most of the in vitro experiments on metals to check corrosive properties are done using saline or isotonic solutions like Hank’s solution. Apart from body fluids, even proteins have been shown to affect the corrosive behaviour of metals (Virtanen 2008). The corrosion may be crevice, pitting, fetting, intergranular or galvanic. Mostly corrosive action initiates with the oxidation of the metal implant leading to insoluble products. Sometimes the debris resulting from the metal implants may induce an inflammatory cascade of reactions after mediation by the macrophages, ultimately resulting in local tissue damage. A surface oxide film protects some metals like Al, Ti, Cr from further corrosion, whereas metals like low-carbon steel keep on corroding until fully consumed. Rapid advancement in technology has led to the development of nickel-free austenitic stainless steel which has better corrosion resistance compared to the traditional 316 L stainless steel (Yang and Ren 2010). CoCr alloys and Ti alloys are also preferred for their anti-corrosive activities. Au and Pd alloys can resist corrosion to a great extent and are hence chosen as preferred dental implant metals. Several strategies like laser melting, plasma spraying, electropolishing and ion implantation have been implemented till date for improving the corrosion resistance of metal biomaterials.

2.2.4

Anti-Bacterial Properties

Metals like Cu, Ag and Zn are known for their widespread clinical applications as anti-bacterial agents (Berry et al. 1992). Ag, though toxic at higher concentrations, exhibits anti-microbial activity via electron transport binding to the DNA, thiol (-SH) group inactivation, and inhibition of DNA replication (Russell and Hugo 1994). Cu also exhibits anti-microbial activity and is much safer compared than Ag. It is necessary for the body in trace amounts and can be easily metabolized by

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Fig. 2.2 (a) Cylindrical samples of additively manufactured porous metallic biomaterial. (b) A femur structure with graded porosity. (c) Porous structure of hip stems. The titanium alloy, Ti-6Al-4 V has been used for most specimens at the Additive Manufacturing Lab, TU Delft. From Zadpoor (2019)

the human body. Reports suggest that a minimum of 5 wt % Cu content is required in alloys like Ti-Cu, for anti-bacterial activity. Conversely, Zn displays anti-microbial activity even in smaller concentrations and is widely used in many household

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products as mouthwash and shampoos (Roldán et al. 2003). Also, anti-bacterial stainless steel has been fabricated by adding 0.04–0.06 wt % Ag and 0.1 wt % Nb to 316 L stainless steel (Junping and Wei 2013). Cu and Nb alloyed stainless steel is also well-known for its bactericidal activities. Cu content of various stainless steel surfaces was altered by using plasma surface alloying technology and explored for possible anti-bacterial activity. Results showed that high Cu (90 wt %) containing surfaces showed better bactericidal effects compared to low Cu (2.5 wt %) containing surfaces (Zhang et al. 2012). Ti alloys, such as Ti-Ag and Ti-Cu, are also famous for their anti-bacterial activities. Zhang et al. also fabricated a layer of Cu-Ni alloyed to Ti to enhance the anti-bacterial activity (Zhang et al. 2013). Mg-based alloys (AZ31, Mg-4Y, Mg) have also been studied for their bactericidal effects (Lock et al. 2014) against E. coli. Mg, being biodegradable, releases Mg ions and forms Mg(OH)2, which leads to sufficient anti-bacterial activities. Even Zr-based bulk metallic glasses (BMGs) containing Cu exhibited sufficient antimicrobial effects over 4 hours.

2.2.5

Bioactivation of Metallic Biomaterials

Biological activation is yet another aspect to be looked upon while developing metallic biomaterials. Certain metals form an apatite layer on their surface after implant, thereby helping in osteoinduction. Metals that cannot form the apatite layer often require surface modification to ensure adherence to the bone and longevity of the implant (Zhou et al. 2014). Some metal oxide gels like TiO2, Nb2O5 and ZrO2 easily form an apatite layer in simulated body fluids. Titanium alloys are widely used for bone tissue engineering due to their excellent biocompatibility and ability to form apatite layers rapidly. Lately, more focus is being laid on fabrication of porous Ti. TiO2 is also used as a suitable coating material to render bioactivity to a metal via electrophoretic sol-gel coating and anodization. One of the most important surface modification techniques to make a metal bioactive is by using calcium phosphate} coating of hydroxyapatite (HA) (Shi and Somberg 2006). This can be done through electrodeposition, sol-gel processing and micro-arc oxidation (Mediaswanti et al. 2013). Mg is also gaining prominence as a scaffold for tissue engineering applications due to its strength, degradation and osteoinductive effects. Mg scaffolds are also being loaded with useful growth factors like bone morphogenetic protein2 (BMP-2), transforming growth factor-β (TGF-β), fibroblast growth factor (FGF) and vascular endothelial growth factor (VEG-F) for different TE applications (Staiger et al. 2006). Metals are also functionalized by immobilization of polyethylene glycol (PEG) on their surface, as PEG inhibits binding of proteins (Mahato 2004). TiO2- and gold-based biomaterials are functionalized using PEG-PLA and PEG-poly (DL-lactic acid). Since the presence of RGD (Arg-Gly-Asp) motif ensures the adhesion and viability of cells, researchers are trying to graft them onto metal surfaces via silanes, thiols or phosphonates (Reyes et al. 2007). Several other proteins, hydrogels, collagens and gelatin had also been explored for adding suitable bioactivity to metal surfaces.

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S. K. Chowdhury et al.

Biodegradation

Metals which exhibit suitable degradation behaviour are being increasingly chosen for soft tissue engineering applications as they help in the tissue regeneration process. Such metals undergo gradual and sustained corrosion, without initiating any adverse reaction due to their release of corrosive by-products. Only metals with appropriate degradation rates are chosen. Temperature, pH and electrolyte concentration of the body fluids at the site of implant determines the nature of corrosive activity. Biodegradable metals are more advantageous compared to other biodegradable ceramics, polymers and bioactive glasses for bone tissue engineering, owing to their tensile strength and similarity in Young’s modulus. Mg, being biodegradable and not skin-sensitizable, has been preferred for bone screws and stents (Staiger et al. 2006) since ages. It is required as a co-factor by many enzymes. Mg implants corrode fast due to the local pH change at the implant site resulting from possible trauma. The degradation rate can be controlled by using processing techniques like alloying, pore formation and coating. Another biodegradable metal, Fe, depends on the oxygen concentration at the local site of implant for its slow degradation activity (Eliaz 2012). However, it can only be used in patients who do not have any iron related disorders. Even Zn has gained prominence due to its biodegradability and bioabsorbability. Zn helps in inhibiting restenosis in arteries and triggering osteogenesis in bones (Katarivas Levy et al. 2017). Porous designs achieved by additive manufacturing can expand the potential applications of these biodegradable metals (Fig. 2.3).

2.2.7

MRI Compatibility

Magnetic resonance imaging (MRI) is a widely used technique that uses strong magnetic fields for medical scanning of tissues and organs in diagnostics. The biggest drawback of metallic biomaterials is their ability to get magnetized in the presence of the magnetic field waves of MRI resulting in generating inappropriate diagnostic results. This has been more visible in the case of intravascular stents, artificial joints, pacemakers and cochlear metal implants, since they are composed of ferromagnetic components (Li and Xu 2014). Of late, Mg, Nb and Zr alloys have been found to be MRI compatible (Suyalatu et al. 2009). A study comparing the movements of orthopaedic metal implants made from stainless steel, Co and Ti under the influence of the magnetic fields (0.3–3 T) generated by MRI machines showed significant visible movements of the artefacts altering the MRI scan results (Shellock 2002). It was concluded that Ti was much safer compared to Co and stainless steel implants, as it had less ferromagnetic content. Zirconium (Zr) was also found to be eminent as a MRI compatible metal owing to their low magnetic susceptibility (1.3  106 cm3 g1) compared to other metal alloys (Mantripragada et al. 2013). At present research is being focussed around Zr-Nb and Zr-Mo alloys. Nb has been a favourite for vascular stents as they are not easily influenced by magnetic fields. The alloy, Nb–xTa–Zr (30  x  70), exhibited optimal properties

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Fig. 2.3 Appearance of additively manufactured porous (a) Magnesium (Mg), (b) iron (Fe) and (c) zinc (Zn) implants during their 4-week in vitro degradation study. (a) Reproduced from Li et al. (2018b) (b) from Li et al. (2018a), (c) from Li et al. (2020a)

like low magnetic susceptibility (Li and Xu 2014) and favoured X-ray imaging, when chosen as a vascular stent. Pt stents are widely used now as balloon expandable stents for MR angiography since they cause only 30% or less artefact induced stenosis.

2.2.8

Radiopacity

Another important property to be looked upon while designing a metallic biomaterial is its radiopacity or the ability to be monitored using X-rays. This is particularly important for vascular stents, where the progression of catheter into the vascular branches needs to be monitored using X-rays during the operation. Techniques to enhance the radiopacity of metals include coating, alloying and introduction of contrast agents (Cheng et al. 2005). Till now, Ta stents have been reported to show highest radiopacity compared to stainless steel, Co-Cr and nitinol stents (Wiesinger et al. 2012). Stainless steel stents have low radiopacity and are hence

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coated with Au, Pt or Ir to increase their visibility (Habibzadeh et al. 2013). Even nitinol exhibits better radiopacity when coated with Pt. Another novel stent made from 33% Pt-Cr alloy was also found to demonstrate significant radiopacity and traceability (Menown et al. 2010).

2.3

Permanent Metallic Biomaterials

2.3.1

Stainless Steel

With the introduction of stainless steel in the metal industry as 18/8 stainless steel, problems of corrosion and poor mechanical stability for use in biomedical implants have been solved (Hatfield 1931). The enhanced corrosion resistance was due to the high chromium content (>12 wt %) together with molybdenum and nickel and lower carbon content. In biomedical or tissue engineering applications, Ni-free stainless steel is preferred. The presence of Ni in the traditional stainless steel adds toxicity to it, and hence nitrogen is added to minimize the risk of any metal allergy. Stainless steel is inferior compared to Ti-based implants in terms of biocompatibility and corrosion resistance, yet it is widely chosen for its low cost. 316 L stainless steel with carbon content less than 0.03% has been widely used for joint replacement (Sumita 1997). The corrosion resistance of stainless steel is dependent on the formation of a passive layer of Cr-Mo oxide. A lot of research work has been done on high nitrogen stainless steels (Katada et al. 2004; Katada and Taguchi 2015) since the 1990s. Pressurized electroslag remelting (ESR) technique is used for alloying nitrogen to the stainless steel (Stein et al. 1999), to get a final product having high corrosion resistance, high tensile strength and poor magnetism. Nitrogen concentrations up to 0.3 wt % are added to austenitic stainless steels to attach the property of crevice corrosion resistance. As traditional stainless steel with high Ni content might cause metal allergy with symptoms of inflammation, rashes, swelling or asthma, high nitrogen steel (HNS) is widely preferred as biomaterial for coronary stent applications in order to avoid the issue of restenosis, which has been visible after the use of Ni containing stainless steel (Costa et al. 2003). HNS steel has better biocompatibility compared to the traditional 316 L stainless steel, as evidenced by the rate of adherence and proliferation of human umbilical vein endothelial cells (HUVECs) which was higher on HNS steel (Costa et al. 2003). Rapid endothelialization and low in vivo inflammatory effects were visible in case of HNS steel (Fig. 2.4). HNS steel has been biofunctionalized by immobilizing vascular endothelial growth factor on its surface via ester bonds (Sasaki et al. 2011). This ensured better biological activity of the HNS coronary stents. As such is the case, Ni-free stainless steels will find wide applications in the tissue engineering sector in future.

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Fig. 2.4 Biocompatibility studies of HNS and 316 L stainless steel with HUVEC cells. (a) Proliferation of HUVECs on HNS (grey) and SUS316 L (black) over 7 days. Data of 5 samples presented as average
standard deviations (* ¼ p < 0.05, N.S. ¼ no significant difference). (b) Morphology of the stained HUVECs (green shows actin filaments and red shows nucleus). From Inoue et al. (2014)

2.3.2

Co-Based Biomaterials

Cobalt (Co)-based implants had been widely used as artificial joints due to their property of intense wear resistance, high strength and ductility (Niinomi 2002). Wrought Co-Cr alloys containing Ni are preferred more over cast Co-Cr alloys due to their inherent strength. They have high elastic modulus and strength compared to Ti-based alloys, but lack adequate biocompatibility. Studies are being conducted to improve the osseointegration capacity of these alloys. Apart from the traditional casting and forging techniques, selective laser melting (SLM) (Takaichi et al. 2013a) and metal injection moulding (MIM) (Tandon 1999) had been used for the fabrication of Co-Cr alloys. Electron beam melting (EBM) is yet another additive manufacturing (3D printing) technique utilized for generating porous Co-Cr alloys (Shah et al. 2016a), which exhibited better bone remodelling and osteogenesis when implanted in adult sheep. 3D printed porous Co-Cr-Mo alloys generated by using SLM technique provide enhanced corrosion resistance compared to the cast alloys (Hedberg et al. 2014). The microstructure of alloys like Co-28Cr-6Mo can be controlled by conjugation of other elements like Zr, and hence it is the most preferred Co-Cr alloy for bone tissue engineering applications. Co-Cr based biomaterials had been widely used as stents, catheters, dental implants. Some of the other applications of such alloys are listed in Table 2.3.

2.3.3

Ti-Based Biomaterials

Titanium is widely chosen for bone tissue engineering over stainless steel and Co-Cr alloys due to its tensile strength, biocompatibility, low density and ability to induce osteogenesis (Sasikumar et al. 2019). The Young’s modulus of Ti alloys closely mimics that of the native bone implanted on to. They are highly resistant to corrosion due to the formation of a stable layer of TiO2 oxide upon oxidation on their surface at

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Table 2.3 Clinical application of Co-Cr based alloys. From Niinomi et al. (2015b) Alloys Co-28Cr-6Mo

Trade name Vitallium (Howmedica, Inc) Haynes-Stellite 21(HS21) (Cabot Corp.) Protasul-2 (Sulzer AG) Zimaloy (Zimmer Inc.) BioDur CCM plus alloy (carpenter technology Corp.)

Co-20Cr-15 W-10Ni

Co-35Ni-20Cr10Mo

40Co-20Cr-16Fe15Ni-7Mo

Haynes-Stellite 25 (HS25) (Cabot Corp.) L-605 (carpenter technology Corp.) MP35N (SPS technologies, Inc.) Biophase (Richards medical co.) Protasul-10 (Sulzer AG)

Elgiloy (Elgiloy ltd.) Phynox (ArcelorMittal stainless and nickel alloys) Conichrome (carpenter technology Corp.)

Applications Stem, ball and cup of artificial joints Fixation screws Bone plates Joint replacements (hip, knee, shoulder) Fixation devices Fixation wires Vascular stents, heart valves Lead conductor wires Springs Stylets Catheters Orthopaedic cables Cardiovascular stents Archwires Springs Lead conductor wires Surgical clips Balloon expandable stents (annealed) Self-expanding stents (aged)

the implant site. Among the three types of Ti alloys, namely α, (α + β) and β types, β type alloys (Niinomi 2008b) are more commonly used in the biomedical industry due to their low elastic modulus. Processes such as cold rolling and severe plastic deformation (SPD) (Matsumoto et al. 2006) are particularly useful in reducing its Young’s modulus. Ti-Zr alloys have been explored for better biocompatibility since the addition of Zr renders high strength and inhibits calcium phosphate precipitation (Kobayashi et al. 2007). This eventually led to the development of Ti-Zr-Nb, Ti-ZrNb-Ta and Ti-Zr-Al-V alloys for clinical applications. Even Ti alloys with selftunable Young’s modulus have been fabricated, e.g. Ti-12Cr alloy (Nakai et al. 2011). Ti-Al-Nb alloys are more researched on now due to their high tensile strength and greater wear resistance (Boehlert et al. 2008) compared to other Ti-based alloys. They exhibited high biocompatibility with almost no signs of cytotoxicity as determined by using LIVE/DEAD assay (Boehlert et al. 2005). Ti-7Al-51Nb is particularly well-known in the biomedical industry for its enhanced cytocompatibility. Ti-1544 alloy is considered biologically safe and non-toxic (Okazaki et al. 1998). It has excellent biocompatibility and apatite formation abilities. Recent trend in the

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field of Ti-based alloy fabrication is the use of additive manufacturing, where the starter material, Ti powder, is fused using selective laser melting (SLM), electron beam melting (EBM) or direct metal laser sintering (DMLS). This technique gives an efficient control over the porosity of the final alloy, which helps in manipulating the strength and elasticity. Research has shown that the attachment of fibrin, which acts as a basal layer for cell adherence, is much better on porous or rough surfaces of Ti implants. Moreover, the 3D network of pores can help the mesenchymal stem cells to easily differentiate into osteoblasts (Schmidt et al. 2002). Rough surfaces have been shown to aid in osteogenesis on the surface of porous Ti-6Al-4 V fabricated by EBM, when human osteoblast-like cells (SAOS-2) were cultured on them for 4 weeks (Hrabe et al. 2013). This implant was found to be osteoconductive, leading to deposition of sufficient collagen. Thus the titanium implants developed through additive manufacturing have gained a lot of attention in recent years due to their fatigue resistance, resulting in their widespread clinical use.

2.3.4

Tantalum and Its Alloys

Tantalum (Ta) demonstrates fantabulous corrosion resistance properties even in acidic conditions owing to the formation of an oxide layer of Ta2O5 (Wang et al. 2012). It was often used to coat stainless steel and titanium biomaterials for enhancing the bioactivity. Studies by Wang et al. have shown that Ta2O5 nanotube films increase the anti-corrosion properties of pure Ta and enhance the adherence, proliferation, differentiation of rabbit bone marrow mesenchymal stromal cells (Wang et al. 2012). It is widely used in orthopaedics for joint replacement and spine fusion implants. It was shown that porous Ta helps in bone and fibrous tissue ingrowth without giving rise to any inflammatory response (Zardiackas et al. 2001). Also, interconnected pores resulted in tensile strength alike cancellous bone. Hence, porous Ta foam structures had been formulated for bone augmentation applications by chemical vapour deposition (CVD) and chemical vapour infiltration (CVI) of Ta on vitreous carbon lattices (Zardiackas et al. 2001). Porous Ta has been used for total hip replacement (THR) implants in cases where there have been substantial bone loss and primary total knee replacement (TKR) implants. They are also used for attachment of ligaments and tendons to the implants due to excellent fibrous tissue ingrowth and are hence chosen in foot and ankle surgery. Guo et al. used 3D printing for fabricating porous Ta scaffolds for bone tissue engineering via SLM technology (Guo et al. 2019). Further, in vitro studies revealed that Ta scaffolds were more biocompatible than Ti-6Al-4 V scaffolds after culturing hBMSCs. Osteogenesis and osseointegration were better visible in case of Ta scaffolds, and hence they can be a promising biomaterial for future bone tissue engineering applications (Guo et al. 2019).

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2.3.5

S. K. Chowdhury et al.

Zirconium Alloys

Zirconium (Zr) alloys are used in knee and hip replacement implants. They have excellent wear resistance due to the formation of protective layer of zirconium oxide. It was shown that Zr-2.5Nb alloy has better tensile strength and wear resistance compared to Ti and Co-Cr alloys (Good et al. 2003). They displayed similar osteogenesis potential as Co-Cr alloys when implanted in a rabbit tibia model. Most of the modern Zr alloys demonstrate low magnetic susceptibility (1.36  106 g cm3) and generate less artefacts compared to other alloys. Their poor strength has been improved by alloying with molybdenum (Mo) and niobium (Nb) (Kondo et al. 2011). Nano-hydroxyapatite has also been used to enrich Zr scaffolds to improve osseointegration.

2.4

Biodegradable Metallic Biomaterials

2.4.1

Mg-Based Biomaterials

Magnesium (Mg) is widely preferred for manufacturing scaffolds for tissue engineering as it is biocompatible, biodegradable and unlike permanent metallic biomaterials it will not initiate hypersensitivity reactions. These scaffolds are used in bone tissue engineering since the ions generated after corrosion are easily absorbed by the body and used for bone growth and strengthening. Also, while undergoing degradation, they form a layer of calcium phosphate around them which further helps in osteoblast proliferation and differentiation. Since Mg is important in calcium uptake by bone, Mg-Ca alloys make the first choice in orthopaedic applications (Serre et al. 1998). Mg-0.7wt%Ca was found to have better ductility and anti-corrosion properties compared to the alloy with 2 wt% Ca (Gao et al. 2009). Grain refinement is conducted to enhance the properties of these scaffolds since they are prone to stress corrosion cracking (SCC) in chloride solutions. Alloys like Mg-Ca, Mg-Sr, Mg-Zn, Mg-Si and Mg-Ag have been extensively studied for their high tensile strength (86–300 MPa) and biodegradability (Ramya et al. 2015; Staiger et al. 2006). In order to increase the degradation time of Mg, metals like Ca, Zn and Mn are alloyed to it. The rate of degradation of Mg alloys depends on the processing technique used. Additional advantage of Mg-based biomaterials is its low magnetic susceptibility which helps in proper diagnostic imaging without interference (Witte et al. 2005). Mg-based alloys have widespread applications in orthopaedics as they bolster the strength of bone–implant interface (Witte et al. 2005). Mg-Si alloys have been fabricated since Si is involved in improving the immune system and bone formation. Zn was also alloyed to the Mg-Si alloy for use in many biomedical applications. Zhang et al. (2010) studied the in vitro cytotoxicity of Mg-6Zn alloys using L929 cell cultures and found them to be potentially safe. Even the in vivo studies by them showed that Mg-6Zn alloys stimulate the bone cells and aid in osteointegration. At present, research is focussed around increasing the bioactivity of the Mg alloys by coating them with suitable substances like HA and

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TiO2 so as to enhance the biocompatibility and bone growth (Amaravathy et al. 2014). Further studies by Gao et al. have shown that ZEK100 Mg alloy coated with chitosan, sodium alginate and mechano-growth factor using a layer-by-layer strategy results in control of degradation rate (Gao et al. 2016). Such coated alloys having good biocompatibility and slow corrosion rate can be used for bone tissue engineering. Though Mg-based alloys have less tensile strength compared to Ti alloys and stainless steel, modifications using additive manufacturing techniques improve their structural properties for use in TE.

2.4.2

Zinc-Based Biomaterials

Zinc (Zn) is involved in many functions of the human body like bone formation and mineralization, enzyme activity, hormone function, nucleic acid metabolism and apoptosis regulation. Concentration of Zn decreases with age in humans. There are many advantages for use of Zn-based biomaterials. Zn has superior anti-corrosion properties when compared to Mg. Corrosion products of zinc-based biomaterials are found to be in no association with the evolution of hydrogen gas as is the case with Mg-based biomaterials. Moreover, the corrosion products of Zn implants help in inhibiting the proliferation of smooth muscle cells and stop restenosis (Bowen et al. 2015). Zn also has very low magnetic susceptibility (15.7  106) compared to Mg and other biodegradable metals. It also has pro-regeneration properties. Hence it is widely chosen as a biomaterial for biodegradable vascular stents. Along with these advantages its major disadvantage is its poor strength, which has been rectified by alloying it with other metals like Al or Mg to make binary Zn-based alloys. Thermal deformation processing is done to improve their corrosion resistance. These Zn alloys are suitable for bone implants as they retain majority of their original mechanical strength during the implant period unlike Mg. In another study by Li et al., Zn-1X (Ca, Mg, Sr) alloys were implanted in mice model for analysing biocompatibility till 8 weeks. It was found that there was no sign of inflammation at the implant site. Osteogenesis led to dense bone tissue formation at the site surrounding the pins. It was also shown that Zn-1Sr exhibited new bone tissue formation in large quantities compared to Zn-Mg and Zn-Ca (Li et al. 2015). Ternary Zn-based alloys have been fabricated by alloying Ca or Sr into the binary alloys resulting in achieving better tensile properties due to the homogeneous distribution of smaller grains and faster corrosion rates compared to the binary alloys (Liu et al. 2016). It was also noticed that the ternary alloys induce osteogenesis at a much faster rate than their binary counterparts. Zn-based ceramic nanomaterials have also gained prominence in the fields of drug delivery, imaging and cancer therapy (Zhou et al. 2015) as they can specifically target cancer cells and deliver therapeutic agents. Since Zn helps with cardiac functions, Zn-Cu and Zn-Li alloys have been preferred as a biomaterial for vascular stents. Zn-based ceramic biomaterials like ZnO, ZnAl2O4 and ZnS have also been explored for clinical applications such as orthopaedic regeneration, drug delivery and cancer therapy (Zhou et al. 2015; Zhu et al. 2016). The promising biological functions, properties

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and results of Zn-based biomaterials have made it a top choice among researchers to improvise it for future biodegradable biomaterial applications.

2.4.3

Iron-Based Biomaterials

Iron is one of the important elements found in humans. Fe can be easily metabolized inside the body by means of haemoglobin, hepatocytes, bone and muscle cells. Accumulation of excess Fe ions may lead to oxidative stress, liver failure and other organ malfunctioning. Hence, there is a need to control the rate of degradation of Fe-based biomaterials. Several in vitro studies have showed that Fe ions drastically reduce the proliferation of smooth muscle cells and thus inhibit vascular restenosis (Zhu et al. 2009). It was also shown that lower concentrations of Fe ions (50 μg mL1) had a passive effect on the endothelial cell growth. When pure iron stents were implanted for 28 days into the porcine coronary arteries, sufficient corrosion was visible compared to the Co-Cr alloy stents (Waksman et al. 2008). However, the endothelial area was more in case of the Fe stents (Fig. 2.5). Structurally iron (Fe) has better mechanical strength compared to other common metals. Its elastic modulus (211.4 GPa) is much higher than 316 L stainless steel (190 GPa) and magnesium (40GPa). Since pure Fe has high ferromagnetism which may obstruct magnetic resonance imaging (MRI) newer alloying materials and processing technologies are used for making them MRI compatible biomaterials. Alloys like Fe-Mn, Fe-Co, Fe-Al, Fe-Sn, Fe-Mn-Si, Fe-Mn-Pd and Fe-21Mn-1C have been fabricated to overcome the shortcomings of pre-iron based biomaterials (Liu and Zheng 2011). Modern processing techniques like electroforming, metal injection moulding, cold gas dynamic spraying and 3D printing are being utilized to fabricate Fe-based biomaterials which are more biocompatible. Fe-based biomaterials have been used in a variety of biomedical applications, the details of which are as listed in Table 2.4.

2.5

Advanced Metallic Biomaterials

2.5.1

Bulk Metallic Glasses

Bulk metallic glasses (BMGs) are the newcomers in the field of biomaterial research. They exhibit excellent processing capabilities and properties as required by complex implants. Due to their amorphous structure along with constituent elements like Ti, Zr, Mg, Fe, Pd, etc. they possess excellent strength and elasticity. In short, they possess the properties of both bioglass and metal alloys. These amorphous alloys are prepared through modified casting leading to either biodegradable BMGs (Mg, Ca, Zn, Sr-based) or non-biodegradable BMGs (Ti, Zr-based) (Schroers et al. 2009). Porous BMGs are often used for research purpose due to their enhanced strength and lightweight. The similarity in the elastic modulus of the biodegradable BMGs and

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Fig. 2.5 Histology of the pure iron and Co-Cr stent embedded porcine coronary arteries after 28 days of implantation. From Waksman et al. (2008)

cortical bone makes them an appropriate candidate for bone tissue engineering (Schroers et al. 2009). In tissue engineering, even biodegradable polymers used till date demonstrated poor strength and uncontrolled degradation rates. Biodegradable BMGs are a wonderful alternative to the traditional metallic implants which require removal post-transplantation affecting tissue healing. Mg-Zn-Ca BMGs are the most preferred choice for biodegradable BMG implants due to their outstanding biocompatibility (Gu et al. 2010a). Mg66Zn30Ca4 and Mg70Zn25Ca5 have been studied for

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Table 2.4 Biodegradable metallic biomaterials used in tissue engineering Biomaterial Publication type year Research findings Magnesium-based biomaterials MgF2 2020 Sucrose is a good coated spacer agent for mg porous Mg scaffolds. The coated scaffold scaffolds had slow degradation rate Mg-Zn 2019 The scaffolds showed scaffold sustained drug release with for anti-bacterial tetracycline activity and osteogenic differentiation Nano-HA 2019 The bioactive foams embedded were porous uniformly. composite Mechanical properties Mg foam were similar to that of cancellous bone ZEK 2016 Layer-by-layer coating 100 alloy of sodium alginate, chitosan and mechanogrowth factor on the alloy reduced degradation rate Mg-2Zn2016 The alloy exhibited 2Gd alloy anti-bacterial activity PCL coated 2014 Increased strength and Mg reduced degradation scaffolds were observed. 2014 Higher osteoinduction, HA/TiO2 coated Mg better biocompatibility alloy and slow degradation were visible Porous pure 2010 The porous scaffold Mg showed minimal cytotoxicity and better corrosion resistance Zinc-based biomaterials Porous Zn 2020 The additively scaffolds manufactured scaffolds had good biocompatibility and anti-bacterial activity Ca-P coated 2020 They had good Zn alloy biocompatibility, no cytotoxicity and aided in osteogenesis

Applications

References

Bone tissue engineering

Toghyani et al. (2020)

Bone tissue engineering

Dayaghi et al. (2019)

Bone tissue regeneration

Parai and BandyopadhyayGhosh (2019)

Bone tissue engineering

Gao et al. (2016)

Not mentioned

Trivedi et al. (2016) Yazdimamaghani et al. (2014)

Bone tissue engineering Various biomedical applications

Amaravathy et al. (2014)

Tissue engineering

Gu et al. (2010b)

Bone tissue engineering

Cockerill et al. (2020)

Cranial bone regeneration

Zhuang et al. (2020)

(continued)

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Table 2.4 (continued) Biomaterial type ZnO composite scaffolds Ca/Sr added Zn-1.5 Mg alloy

Publication year 2020

2016

Zn wires

2015

Zn-1 Mg, Zn-1Ca, Zn-1Sr alloys Zn alloys

2015

2013

Iron-based biomaterials Porous iron 2020 scaffold

Porous iron

2018

Fe-based metallic glasses Fe-based metallic materials Nitriding iron stents

2016

Pure iron stent

2015

2013

2009

Research findings They helped in the osteochondral differentiation of MSCs The alloys demonstrated high yield strength, tensile strength and increased corrosion It inhibited arterial restenosis with no signs of inflammation They had enhanced biocompatibility and better mechanical properties They possess the ideal corrosion behaviour required by stents The 3D printed scaffolds demonstrated anti-platelet adhesion property, good cytocompatibility and hemocompatibility The scaffolds made by direct metal printing (DMP) mimicked mechanical properties of trabecular bone They were non-toxic and possessed apatite forming ability Poor biocompatibility and cytotoxicity were detected High tensile strength, fast corrosion rate and restenosis were observed after 1 year High Fe ion concentration may negatively affect the endothelial cells

Applications Bone and cartilage tissue engineering Not mentioned

References Khader and Arinzeh (2020) Liu et al. (2016)

Endovascular stents

Bowen et al. (2015)

Biodegradable implants

Li et al. (2015)

Cardiac stents

Bowen et al. (2013)

Bone tissue engineering

Sharma et al. (2020)

Bone tissue engineering

Li et al. (2018a)

Stents, orthopaedic implants Cardiovascular applications

Qin et al. (2016)

Fagali et al. (2015)

Stent

Feng et al. (2013)

Vascular stents

Zhu et al. (2009)

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their in vitro cytocompatibility using L929 and MG63 cells. Results showed that compared to pure Mg, they exhibited high strength, low elastic modulus, negligible hydrogen gas release, high cell viability and better cell proliferation and their corrosion products, Mg(OH)2 and Zn(OH)2 are found to be non-toxic and aided in making the BMGs corrosion resistant (Gu et al. 2010a). Ca-based BMGs had also been explored for biomedical applications, but studies are hampered owing to their faster degradation rates. Ytterbium (Yb) is added to Ca-based alloys to fabricate Ca-Yb-Zn-Mg-Sr BMG with enhanced corrosion resistance (Li et al. 2013). The Ca-Yb-Zn-Mg-Sr BMGs also showed better adhesion, proliferation and differentiation of osteoblasts. They successfully reduced degradation time and enhanced osteogenesis of the Ca-based BMGs. Zn-based BMGs like Zn38Ca32Mg12Yb18 exhibited better biocompatibility when evaluated with MG63 cells. Sr-based BMGs have been explored for bone tissue engineering since strontium (Sr) can induce osteogenesis and inhibit bone resorption. Sr40Mg20Zn15Yb20Cu5, a Sr-based BMG was found to possess enhanced strength, biocompatibility and corrosion resistance (Li et al. 2012). Among the non-biodegradable BMGs, Ti-based implants had gained immense attention due to their high yield strength, low elastic modulus and superior bioactivity. Studies have even reported the deposition of hydroxyapatite on the surface of Ti-based BMGs like Ti-Zr-Fe-Si alloys, when immersed in simulated body fluid, thereby assuring their biocompatibility. Zr-based BMGs have high tensile strength compared to the crystalline metallic biomaterials (Bai et al. 2008). This makes them potential for use in cardiovascular stents, where thin struts are preferred. Nb and Ag have been added to Zr-based BMGs in some studies and are reported for better pitting corrosion resistance (Lu et al. 2012). More attention is being focused on use of modern processing technologies like additive manufacturing for fabricating the future BMGs to be used in tissue engineering.

2.5.2

Shape Memory Alloys

Shape memory alloys (SMA) are the alloys which have the ability to recover their original shape even after undergoing extensive deformation under pressure. The deformed shape reverts to the original shape on application of heat to cross the transition temperature (Tadaki et al. 1988). The most prominent SMA studied nowadays is the nitinol (NiTi) alloy which has good fatigue, thermo-elasticity and corrosion resistance due to the TiO2 layer formed on it. It has been widely preferred over stainless steel in biomedical applications due to its better biocompatibility and less interference with MRI and CT imaging (Shabalovskaya et al. 2008). Ni-Ti wires have been used as self-expandable stents, intra-spinal implants and in dental applications (Schillinger et al. 2006). The surface properties of nitinol alloy can be controlled by various processing techniques like thin film coating, annealing, ion implantation, chemical etching and laser melting (Shabalovskaya et al. 2008). Also, TiNi SMA foams have been designed for use in bone tissue engineering using spaceholder sintering technique to increase their porosity, thereby resulting in decrease in their elastic modulus so as to resemble the mechanical properties of cancellous bone

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(Xiong et al. 2008). At present, Ni-free SMAs are in focus of research studies worldwide, since the release of Ni from TiNi alloys might cause cytotoxicity to an extent.

2.6

Tissue Engineering Applications of Metallic Biomaterials

2.6.1

Bone Tissue Engineering

Metallic biomaterials always remained at the forefront of biomaterial applicability for bone tissue engineering due to their longevity and tensile strength. Also, several metal ions help in the proliferation and differentiation of the osteoblasts, thereby influencing bone healing (Fig. 2.6). Metallic biomaterials have often been used to stabilize bone fractures. As Ca enhances bone tissue regeneration (Hu et al. 2014), Ca-based biomaterials have opted for bone tissue engineering. For use in permanent implants such as artificial joints, high nitrogen steel (HNS) is widely preferred over Co-Cr alloys for their excellent strength and biocompatibility (Katada et al. 2004). Further, studies have shown that Co-based hydrogels loaded with BMP-2 and

Fig. 2.6 Representation of the roles of metal ions in bone tissue engineering. From Glenske et al. (2018)

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Co-doped bioactive glass induce bone graft vascularization (Perez et al. 2015). Another metal, iron (Fe) is found to provide adequate support for osteoclast differentiation. Hence Fe-based biomaterials with controlled degradation rate are being used for osteogenic applications (Jia et al. 2012). Fe and Mg had been used as biodegradable bone scaffolds (Farack et al. 2011) for ages due to their excellent cytocompatibility and mechanical properties. Mg-based biomaterials are widely used for functional bone tissue engineering (Farraro et al. 2014). Mg-based biomaterials are found to form a layer of calcium phosphate as a corrosion product, which results in increase in bone area (Witte et al. 2005). Studies on porous Mg scaffolds coated with polycaprolactone (PCL) prepared by powder metallurgy technique to investigate their role in bone tissue healing showed that these coatings provided controlled degradation rates and enhanced mechanical properties confirming the positive role of such scaffolds in bone tissue engineering (Yazdimamaghani et al. 2014). Also, studies on Ti-based scaffolds loaded with growth factors such as TGF-β and BMP have shown sufficient osteoinduction when compared to the traditional Ti alloys making them suitable for use as bone scaffolds (Jansen et al. 2005). Even porous Ta had been demonstrated to be effective as a biomaterial for knee replacement implants (Bobyn et al. 1999). Shape memory alloys (SMAs) like TiNi alloy foams have also been used in bone tissue engineering (Xiong et al. 2008). Recent advances in orthopaedic research are focused on rectification of the stress shielding effects and surface modifications of metallic biomaterials to render them useful for future applications.

2.6.2

Cartilage Tissue Engineering

Biodegradable metallic biomaterials are the preferred choice of metals for tendon and ligament tissue engineering. Mg-based biomaterials are mostly used for this purpose. Studies demonstrated that Mg-based anterior cruciate ligament (ACL) interference screws could assist in the ACL graft healing (Farraro et al. 2014). Shape memory alloys (SMAs) have been shown to be successfully used as flexor tendon substitutes (Moneim et al. 2002) due to their high tensile strength and biocompatibility apart from their super-elastic property. Implantation studies in rabbits have shown that nitinol (Ni-Ti) alloy sutures exhibited more strength and biocompatibility compared to polyester sutures for application as artificial tendon (Kujala et al. 2004). Use of tantalum (Ta) in dynamic culture has shown that it is chondro-conductive and that 84.5% collagen type II and glycosaminoglycans (GAGs) deposited on the surface of the metal (Gordon et al. 2005). More research is yet to be done to justify the suitability of metallic biomaterials for chondrogenic implants.

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Cardiovascular Tissue Engineering

Cardiovascular health has been a major concern for both the developing and developed countries since ages. Coronary artery diseases are treated using stents coupled with balloon angioplasty. Traditional balloon expandable stents were made up of 316 L stainless steel and Co-Cr alloys that demonstrated sufficient strength, corrosion, resistance and radiopacity. Co-Cr alloy-based stents are found to be better when compared to the stainless steel as they reduce vascular restenosis by means of their thin struts (Pache et al. 2003). Ti-Ni shape memory alloys are preferred in fabrication of self-expandable stents owing to their elasticity and biocompatibility (Schillinger et al. 2006). Though with the advent of permanent stents, the need for bypass surgeries drastically reduced to below 0.5% and restenosis rate by 30%, major concerns remained in place since permanent stents required a second surgery for their removal and are prone for the corrosion by-products (Bowen et al. 2016). This necessitated the need for biodegradable stents like Mg-, Zn- and Fe-based stents which minimize the chances of chronic inflammation and release degradable bioactive ions on corroding (Bowen et al. 2016). A vascular stent made of Mg-based alloy, AE21 (Mg–2wt%Al–1 wt% rare earth element) displayed sufficient endothelialization upon implantation in a porcine artery (Heublein et al. 2003). Biocompatibility tests on pure Fe-based stents revealed that high iron ion concentration (>50 μg/ml) might lead to cytotoxicity of endothelial cells (Zhu et al. 2009). Zn-based stents are largely chosen as they have the appropriate degradation rate between that of Fe and Mg (Bowen et al. 2013). Zn wires have been shown to inhibit arterial restenosis with no signs of inflammation. The degradation products of Zn alloys inhibit the proliferation of smooth muscle cells. Many clinical trials have been conducted recently using biodegradable stents. Absorbable metal stent (AMS) (BIOTRONIK, Berlin, Germany), when implanted in 20 patients with critical limb ischemia, degraded after 6 weeks (Peeters et al. 2005). The first-in-man trial (BIOSOLVE-I) of drug eluting absorbable metal scaffold (DREAMS) on 46 patients reported no scaffold thrombosis (Haude et al. 2016). Subsequently, the BIOSOLVEII study with DREAMS 2.0 incorporating the sirolimus drug resulted in better endothelialization along with more efficacy (Kitabata et al. 2014). At present, rapid advancement in technology led by an increased demand for biodegradable metallic biomaterials has improvised the ongoing research work on metallic biomaterials for cardiovascular applications.

2.6.4

Dental Tissue Engineering

Metallic biomaterials are used in different dental applications like filling of cavities, constructing bridges and wires or in dental implants. The type of material used for the purpose is selectively chosen since some metals can easily corrode under the influence of enzymes or wear rapidly. Usually, orthodontic wires are made up of stainless steel, Co-Cr based alloys or Ti-based alloys like NiTi shape memory alloy. In the fabrication of dental amalgams for replicating the shape of the tooth cavity,

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Table 2.5 Recent advances and applications of metallic biomaterials in tissue engineering Applications Bone tissue engineering

Cartilage tissue engineering

Cardiovascular tissue engineering

Biomaterial Mg-based scaffolds

Properties Controlled biodegradability

Porous Zn scaffolds

Promotes osteogenesis

Porous iron scaffold ZEK 100 alloy

Enhanced cytocompatibility Controlled degradation

Ti6Al4V implants

Promotes bone maturation

Mg-based metal organic frameworks Sr-based metal organic frameworks Mg interference screws Self-expandable metallic stents Zn-Cu coronary stents

Promotes chondrocyte proliferation Treating osteoarthritis

Mg-, Zn-, Fe-based stents Zn alloys Fe-based metallic glasses Zn-based stents Dental applications

Ti-6Al-4 V scaffolds Porcelain-Ti dental alloys

Better ACL graft healing Effective for obstructive atelectasis Better recovery of vascular pulsatility Inhibits in-stent restenosis Better biodegradability and biocompatibility Apatite forming ability Appropriate biodegradability High yield strength and fatigue strength Increased strength

Reference(s) Toghyani et al. (2020) Cockerill et al. (2020) Sharma et al. (2020) Gao et al. (2016) Shah et al. (2016b) Li et al. (2020c) Li et al. (2020b) Farraro et al. (2014) Bi et al. (2020) Zhou et al. (2020) Sangeetha et al. (2018) Bowen et al. (2016) Qin et al. (2016) Bowen et al. (2013) Xiong et al. (2020) Antanasova et al. (2020)

mercury (Hg) is often added to metal alloys (Eichner 1983). Gold alloys supplemented with Ag or Cu and Ag alloys are preferred for dental castings. The addition of Cu enhances strength and decreases the melting point, whereas Ag reduces inflammation. Ti-based alloys like Ti-6Al-7Nb are also chosen for dental castings due to their enhanced mechanical properties and corrosion resistive nature (Eichner 1983). Selective laser melting (SLM) technique is used extensively for fabrication of Co-Cr-Mo alloys in manufacturing of dental implants and wires (Takaichi et al. 2013b) for dental applications. Ti-Ta alloys have also been explored for dental use and were found to be better than the Ti-6Al-7Nb alloy in terms of corrosion resistance (Mareci et al. 2009). Porcelain-Ti alloys have also been favoured for dental applications as stated in Table 2.5. With the advent and use of modern processing techniques rapid progress is being made in the fabrication of high

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performance metallic biomaterials for use in the field of dental applications nowadays.

2.7

Future Prospects of Metallic Biomaterials in Tissue Engineering

Metallic biomaterials play an indispensable role in tissue engineering. Porous metallic biomaterials are very significant in this field as they possess adequate mechanical properties suitable for supporting cell adhesion, proliferation and differentiation. Biodegradable metallic biomaterials like Zn-based materials are gaining prominence due to their controlled degradation rate and high in vivo cytocompatibility. Of late research is being focused to minimize the hydrogen gas evolution from the metallic scaffolds during degradation. Since surface topography of the scaffolds plays a crucial role in directing the cells towards differentiation, modern processing techniques like 3D printing are being explored for generating porous or nano-patterned surfaces on the metallic biomaterials. Additive manufacturing techniques need to be employed for fabricating scaffolds that mimic the native properties of the tissue to be replaced. Revolutionizing metallic biomaterials like bulk metallic glasses (BMGs) and shape memory alloys (SMAs) will be the most preferred choice of biomaterials for future applications in this field. Bio-functionalization of the metallic biomaterials is an area that needs to be focused on. Having said that, further research work is required to render the metallic biomaterials bio-functionality and amplify their potential applications in tissue engineering. Apart from scaffold fabrication, modern technology has also opened the doors for research in biosensors and bioresorbable electronic stents, where degradable metallic biomaterials are chosen. Rapid advancement in material sciences coupled with multi-disciplinary research upholds the tremendous potential of metallic biomaterials in the vast ever-growing field of tissue engineering.

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Bioceramics in Tissue Engineering: Retrospect and Prospects P. R. Harikrishna Varma and Francis Boniface Fernandez

Abstract

Bioceramics are widely adopted in the field of skeletal tissue reconstruction, augmentation, and replacement. Apatite-based materials have shown great promise in the latter half of the nineteenth century in this area. Understanding demands on implants and providing a pathway to match it has been a learning experience. Clinicians and technologists have been on this path assiduously to optimize graft properties vis-a-vis biological demand. This is due to their widely acceptable traits of biocompatibility, osteoconduction, and osteointegration. This chapter attempts to track their development over time and track applications of the materials in their maturation from pure graft materials to tissue engineering scaffolds and smart materials. Keywords

Calcium phosphate · Hydroxyapatite · Hard tissue engineering · Osteointegration · Osteoconduction

P. R. H. Varma Biomedical Technology Wing, Sree Chitra Tirunal Institute for Medical Sciences and Technology, Thiruvananthapuram, Kerala, India F. B. Fernandez (*) Division of Bioceramics, Department of Biomaterial Science and Technology, Biomedical Technology Wing, Sree Chitra Tirunal Institute for Medical Science and Technology, Thiruvananthapuram, Kerala, India e-mail: [email protected] # The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 B. Bhaskar et al. (eds.), Biomaterials in Tissue Engineering and Regenerative Medicine, https://doi.org/10.1007/978-981-16-0002-9_3

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Abbreviations ALP BMP CaP Hap MSC TCP

Alkaline phosphatase Human bone morphogenic protein Calcium phosphate Hydroxyapatite Mesenchymal stem cells Tricalcium phosphate

3.1

Introduction

Hindsight is a privilege of history that allows us to be wise after an event and to be able to evaluate the effects of preceding actions and innovations. We are at the cusp of massive changes in the field of regenerative medicine that is driven by innovation from material scientists, biologists, and engineers. At this juncture, it is judicious to review stepping stones that have to lead us here and that which will guide us forward. The term bioceramics encompass ceramics with proven biocompatible nature that are applicable in biomedical or clinical use cases. They are generally classified based on their composition into mainly two groups: calcium phosphates and others. Calcium phosphate bioceramics, over the past two decades, have gained considerable space in orthopedics and dentistry. This is in sharp contrast to the massive number of materials and scaffolds that are proposed for translation with very few achieving clinical efficacy. Demands placed on hard tissue analogs regardless of their synthetic or natural origin are several. They can be permanent or biodegradable but should be biocompatible, ideally osteoinductive, osteoconductive, osteointegrative with sufficient porosity and mechanically compatible with native tissue milieu to fulfill their role in the desired manner. To develop necessary functionality diverse application forms have been developed ranging from cements, coatings, scaffolds, and paste forms. Table 3.1 is attempted to list the variety of properties demanded and their short definitions (Albrektsson and Johansson 2001; Lopes et al. 2018; Yang et al. 2005; Fernandez et al. 2020; Williams 2003). Calcium phosphate based bioceramics are the focus due to their wide application and acceptance in the commercial implant market. This demands their availability in various forms ranging from films, nanopowders, granules, porous, or dense bodies as required. Application needs dictate the use of these forms with nonload bearing areas opting for the application of highly porous apatite blocks. This chapter would aim to provide a comprehensive understanding of the background perspective of bioceramics with special attention to calcium phosphates, bioactivity of calcium phosphate, variants of calcium phosphate, their applications in tissue engineering including developments at clinical scale.

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Table 3.1 Desired properties and their definitions for biomaterials for hard tissue engineering applications Property Resorption Wettability Osteoinduction Osteoconduction Osteointegration Biocompatibility

Definition Gradual degradation of implanted material over time and replacement with natural host tissue The ability of the material to attract or repel water molecules translates into biological activity in situ Induction of osteogenesis. Denotes the cascade that promotes differentiation of undifferentiated capable cell types into bone-forming cell lineages The ability of a biomaterial to serve as support material in the bone-forming process or propagation of bony tissue Denotes the ability of the material to form firm anchorage in bone without ingress of fibrous tissue at the interface The ability of a material to perform with an appropriate host response in a specific application

Fig. 3.1 Synthesis of hydroxyapatite via the wet precipitation process: (a) The cloudy fine precipitate can be observed being collected under continuous stirring. (b) Collected apatite precipitate is washed and freeze-dried before spray drying. (c) Scanning electron micrograph of spraydried apatite powder, note size distribution of particles, and spherical nature of the powder

3.2

Background Perspective

Calcium phosphate ceramics are the common name for a large unit of materials that contain calcium ions (Ca2+) with orthophosphate (PO3
4), metaphosphate (PO
3), or pyrophosphate (P2O4
7) anions and sometimes hydrogen (H+) or hydroxide (OH
) ions. It forms the major inorganic component of enamel (90%) and that of bone (~60%). Of importance to us are the calcium phosphates that are derived to have an atomic ratio Ca/P between 1.5 and 1.67 called apatites such as hydroxyapatite. Synthesis by wet precipitation is a simple process to set up at room temperature as illustrated in Fig. 3.1. Many have carried out detailed investigations of the material with special emphasis on its attempts at translational application (Epinette

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and Manley 2004; Dorozhkin 2012). A well-documented study in 1900 details the replacement of a trepanation window using a bone autograft. This sparked interest in the use of autograft material harvested from the donor (Sparks et al. 2018) or from similarly aged patients (Macewen and Huxley 1881). These interventions were marked to remain as outlying curiosities until Fred Houdlette Albee attempted to implant lab-produced CaPs in a rabbit model (Albee 1920). This was followed by the observation that healing was rapid when tricalcium phosphate (TCP) was injected into the defect area than in the controls. More interest in apatite’s leads to the reporting of the crystal structure of various apatites (Hendricks et al. 1931) and also the analysis of biologically occurring apatites (Jensen and Möller 1948). The differentiation between the various phases of apatite was made in the 1930s. Concurrently more information on the influence of apatite materials on the healing process came to light (Schram and Fosdick 1948). The role of apatites and their ability to mediate the osteoinductive process were studied in detail. It was understood that a positive effect on osteoinductive processes will only contribute towards bone healing. Branemark coined the term osseointegration in 1952, a milestone in the field. The experiment conducted by him resulted in the integration of a titanium chamber into rabbit bone so completely that it could not be removed. The preliminary definition was that material is osseointegrated when there is a direct as well as a functional connection between the surface of the implant that is load carrying and the native bone (Brånemark 1983). It has evolved over the years to mean in 1986 the contact established without the interposition of nonbony tissue between normal remodeled bone and an implant entailing a sustained transfer and distribution of load from the implant to and within the bone tissue (Vaidya et al. 2017). The definition has evolved over the years to indicate the apparent direct attachment or connection of osseous tissue to an inert alloplastic material without intervening connective tissue (Jayesh and Dhinakarsamy 2015). Branemark went on to pioneer implantology in humans gaining the moniker “the father of modern dental implantology.” Interest in calcium phosphates and derivatives continued with Posner describing the structure of amorphous calcium phosphate and suggested the use of the term “Posner’s Cluster” to describe the smallest constituent unit (Posner and Betts 1975). From the year 1969 onwards there are reports of fabrication of hydroxyapatite (HAp) implants via hot pressing and their application (Levitt et al. 1969). From there we can see the rise in the application of materials as grafts in multiple use cases. The use of porous β-TCP scaffolds was reported in 1971 (Bhaskar et al. 1971a, b), implantation of resorbable and porous CaPs was reported from 1975 (Habraken et al. 2016). The immediate closure of tooth root areas was achieved by dense HAp cylinders by 1979 (Denissen and de Groot 1979). Films and layers of HAps were used from 1976 (León 2009) and the development of composites (Sudo et al. 1976) and hybrid materials followed shortly (Bonfield et al. 1981). Since the 1980s the use of the HAp blocks and coatings to assist in bone-anchoring was adopted by the dental community (Jarcho 1981). This leads to orthopedists adopting the same for building up of bone defects and interest in the coating of large metal implants with coatings to ensure anchorage (Epinette and Manley 2004). Furlong was instrumental in the

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Fig. 3.2 Hydroxyapatite forms showcase diverse graft material development based on application milieu. (a) Porous granules (2–3 mm size), (b) fine granules (250–1000 μm), and (c) powder 0 in the chemical composition of “precipitated hydroxyapatite,” one talks also about “calcium-deficient hydroxyapatite” (CDHA). Generally, x ¼ 1 so that CDHA has in most cases the composition Ca9(HPO4)(PO4)5OH. To narrow the subject further and focus the discussion we will concentrate on the undoped CaPO4 and move on. Readers interested in other variants may kindly consult the references provided.

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Table 3.2 The main variants of calcium orthophosphate compounds are listed below in a table adapted from (RMS Foundation, CH-2544 Bettlach, Switzerland and Bohner 2010). Current version adapted from Habraken et al. (2016) Name Monocalcium phosphate monohydrate Dicalcium phosphate Dicalcium phosphate dihydrate Octocalcium phosphate Precipitated hydroxyapatitea Precipitated amorphous calcium phosphate Monocalcium phosphate α-Tricalcium phosphate β-Tricalcium phosphate Sintered hydroxyapatite Oxyapatite Tetracalcium phosphate

Formula Ca(H2PO4)2H2O

Ca/P 0.50

Mineral –

Symbol MCPM

CaHPO4

1.00

Monetite

DCPA

CaHPO42H2O

1.00

Brushite

DCPD

Ca8H2(PO4)65H2O

1.33

–

OCP

Ca10
x(HPO4)x(PO4)6
x(OH)2
x

1.33–1.67

–

PHA

Mu(Ca3)(HPO4)3v(PO4)3yzH2O)b,c

0.67–1.50

–

ACP

Ca(H2PO4)2

0.50

–

MCP

α-Ca3(PO4)2

1.50

–

α-TCP

β-Ca3(PO4)2

1.50

–

β-TCP

Ca10(PO4)6(OH)2

1.67

Hydroxyapatite

SHA

Ca10(PO4)6O Ca4(PO4)2O

1.67 2.00

– Hilgenstockite

OXA TetCP

a - x may vary between 0 and 2. b - u may vary between 0 and 3, v may vary between 0 and 1.5, y may vary between 0 and 0.667, and z is unclear at this point. M is typically a monovalent cation (Na+, K+, NH4+) which is only present if there is an overall negative charge on the calcium phosphate. c - ACP produced in basic conditions has generally u ¼ 0, v ¼ 0, y ¼ 0.667, leading to the following composition: Ca3(PO4)2zH2O where z ¼ 3–4.5. In acidic conditions, u ¼ 3, v ¼ 1.5, y ¼ 0, leading to the following composition: M3(Ca3(HPO4)4.5zH2O) where z is unknown

3.3

Bioactivity of Calcium Phosphate

Calcium phosphates composed of calcium cations and phosphate anions are the major inorganic material in approximately 60% of human bones (Bose and Tarafder 2012) with their existence identified earlier in the 1700s and application in practical uses mentioned earlier. Several dental and orthopedic applications rely on their bioactivity to serve as bone cements, scaffolds, implants, and coatings. The key mechanism is their ability for partial dissolution and to generate ionic products in vivo that cause a hike in local levels of calcium and phosphate ions and thus

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Fig. 3.3 Hydroxyapatite foam-based templated lattice, this method generates openpore architecture

induces precipitation of a biological apatite on the surface of the implanted ceramics (Ben-Nissan 2014). This in turn affects the local expression of osteoblastic differentiation markers such as COL1, ON, RunX2, ALP, BMP’s, OPN, OCN (Frank et al. 2002; Whited et al. 2006). They can influence cell adhesion and tissue formation by modulating the adsorption of extracellular matrix proteins on the surface (Fujii et al. 2006; Tsapikouni and Missirlis 2008). The porosity of the implant is vital factor which facilitates cell ingrowth and colonization. The porous architecture of hydroxyapatite form lattice accounts for the generation open-pore architecture (Fig. 3.3). Calcium ions play a major role in many signaling pathways and serve as an indicator, modulator, and regulator in several body functions (Song et al. 2019a, b). The ion forms a major part of the bone matrix and is present locked in the form of calcium phosphates in bony tissue (Peacock 2010). They induce bone formation and maturation through calcification and affect bone regeneration through cellular signaling. The formation of nitric oxide and stimulation of osteoblastic synthesis pathway in bone precursor cells is also carried out by calcium stimulation. The bone synthesis pathway can be activated via the P13K/Akt pathways to increase osteoblast lifespan and the ERK1/2 pathways for bone synthesis (Danciu et al. 2003; Liu et al. 2008). Osteoclastic activity is also modulated by calcium signaling in diverse ways that are still being explored (Henriksen et al. 2011). Phosphorus ions in the body are a majority with involvement in proteins, nucleic acids, and the energy currency as adenosine triphosphate with the ability to affect physiological processes (Goretti Penido and Alon 2012; Khoshniat et al. 2011). Calcium phosphates lock up 80% of the phosphate ions with the majority of phosphate existing in the PO43
form. This has a major effect on cell growth, maturation, and functionality as more work presses on (Khoshniat et al. 2011). Phosphate ions play a major role in osteoblastic lineage modulation via the BMP’s increased expression and also via the IGF-1 and ERK ½ pathways (Tada et al. 2011). Reduction of RANK ligand:OPG ratio to negatively influence osteoclast

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Fig. 3.4 Hydroxyapatite burr hole closure devices. Note varying porosity in peg vs. cap structures in monolithic development for easy handling and integration. Burr hole closure devices are used in neurosurgery interventions to facilitate the complete reconstruction of intervention sites

differentiation and bone resorption is carried out. This is a negative feedback role (Tada et al. 2011). Based on activity and CaP ratios and availability of free ions there is a discernible difference between apatite families on biological activity (Oberbek et al. 2018). This has led to the development of specialized materials that depend on pure/mixed phase materials to achieve demanding performance characteristics. Burr hole closure devices made up of hydroxyapatite are used in the neurosurgery interventions (Fig. 3.4). Calcium phosphate ceramics (CPC) owe their adaptability as graft materials to their excellent properties captured in Table 3.1 and explained in detail concerning its biological activity. A key factor is their ability to induce progenitor cell differentiation into the osteoblastic lineage that is critical for defect repair as well as bone formation in a nonbony milieu (Samavedi et al. 2013). Their ability to conduct bone growth on surfaces via osteoconductivity is also well adapted for repurposing as graft units (Albrektsson and Johansson 2001). The ability to carry out the above is due to support for cell proliferation and cell adhesion as well as the ability to mediate the differentiation process that is undertaken by cell groups during tissue formation (Samavedi et al. 2013). Adhesion of cells to the surface is strongly governed by CPCs ability to bind extracellular matrix proteins, surface roughness, crystallinity, solubility, phase, porosity, and surface energy. Samavedi et al. (2013) provide a bird’s eye view of the variables involved in the development of biologically adaptable CPC systems (Fig. 3.5). The calcium phosphate crystal structure and its grain and particle sizes help determine the surface roughness that in turn controls protein adhesion to the surface (Deligianni et al. 2001). Protein adsorption has been demonstrated to improve at a roughness of less than 100 nm and is also linked to cell adhesion (dos Santos et al. 2008).

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Bioceramics in Tissue Engineering: Retrospect and Prospects

Crystallinity & Ca/P ratio

Surface roughness

Surface roughness

Solubility

Surface charge/ energy

69

Crystallinity & Ca/P ratio

Solubility

Protein adsorption

Cell adhesion

Grain & particle size Surface charge/ energy

Solubility

Cell differentiation

Surface roughness

Crystallinity & Ca/P ratio

Fig. 3.5 Schematic of key properties of CPC that affects the cascade of biological processes not limited to protein adsorption, cell adhesion, and cell differentiation. Reproduced from Samavedi et al. (2013) with permission #2013 Elsevier B.V

The porosity of materials is the result and function of their synthesis and process pathways and is of great interest to material scientists and biologists alike. This has a direct effect on bioactivity and draws great interest in its control and prediction (Sipaut et al. 2016). Various methods including surfactants have been adopted for the same (Nga et al. 2014). The rise in porosity increases overall contact with biological fluids that enhances dissolution (Sun et al. 2002) with an increase in the preferential dissolution of amorphous phases in all regions (Maté Sánchez de Val et al. 2016). Pore size also plays a role in bone ingrowth (Mygind et al. 2007) and also angiogenesis (Sakamoto 2010). 50 μm or greater pores were beneficial for the ingrowth of blood vessels and it drives angiogenesis as well as providing for bone conduction (Dorozhkin and Epple 2002; Saiz et al. 2007). About 100 μm or bigger pores would actively interfere with the mechanical strength and shape of calcium phosphate (Dorozhkin 2010). The cause of strength decrease in MP scaffolds (CaP with microporosity and macroporosity) in a biological milieu on comparison with NMP (CaP with the only macroporosity) scaffolds may be due to the increased degradation induced by osteoclasts at the reactive gran boundaries than that of NMP (Woodard et al. 2007). A larger specific surface area can be achieved by increasing the number of micropores; will be essential for osteoconductivity that will favor bone regeneration. This is attributed to the growth factor retention abilities of the observed microporosities especially applicable to bone formation in ectopic sites (Chen

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et al. 2017). This postulates that nanoporosity could boost osteoinductivity in hard tissue engineering by enhancing osteogenic differentiation. In an ectopic ovine model higher bone formation was seen in scaffolds with increased strut porosity (Coathup et al. 2012). Osteoinduction was observed in 50% of β-TCP materials that had a 60% porosity with nil osteoinduction observed in a 75% porosity group (Tsukanaka et al. 2015). Microporosity is a necessity for osteoconductivity that leads from inner pores to a large surface area which in turn is inductive for bone tissue formation as well as protein absorption which contributes to ion exchange and biological apatite formation (Schnieders et al. 2011). The dissolution processes of CaPs is also affected by surface area, temperature, and other conditional elements (Ben-Nissan 2014; Ambard and Mueninghoff 2006). The stability and solubility of calcium phosphates are determined by this, and as a thumb rule, it is inversely proportional to the ratio of Ca/P ions, purity, and crystal size of the material. Stable moieties indicate the same low ion exchange with local surroundings and slow crystallization on the surface which in turn determine the protein concentration and conformation by electrostatic interaction at the charged site. Phases with high solubility tend to dissolve faster causing changes in the local surroundings as well as precipitating a quick formation of biological apatite. This has a direct effect on protein adhesion, which may impact cell adhesion and determines the effectiveness of bone regeneration (Hu et al. 2007; Bodhak et al. 2009; Gustavsson et al. 2012).

3.3.1

Calcium Phosphates: Variants and Effects

In the perspective of the continuous development in the field of biomaterials, there is constant debate as to the status of CaPs. Are they just from an older generation that are functional and acceptable but lack in elegance? Or does their traditional acceptance and familiarity with use and adaptation lend to a bright future? A major tie-breaker here is that unlike polymers, composites, and novel combinations that are being proposed day-to-day CaPs are a mile ahead due to their natural presence in the body and ease of regulatory approval. In an aging world, this provides endless opportunities for the large-scale cheap production and engineering of these materials to meet new demands. They can be fabricated in the desired shape and sizes as depicted in Fig. 3.6. A quick overview in Table 3.3 provides us a top-down view of the origin, maturation, and burgeoning of this field. Extrusion of apatite into clinically significant shapes played a major role here (Velayudhan et al. 2000).

3.3.2

CaPO4 Bioceramics in Tissue Engineering

Tissue and organ repair have been the penultimate goal of surgery. This rings true from ancient times (Bose and Tarafder 2012) but with a new horizon based on augmentation or replacement insight (Mandrycky et al. 2017). This has been approached primarily to develop substitutes for organs that will serve for

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Fig. 3.6 Evolution of apatite forms fabrication—from perfectly machined forms in defined geometries to challenging sintered curved surfaces

transplantation or restoring lost function. Secondarily a need for fully competent systems built from the ground up is required for drug testing and disease modeling (Jensen et al. 2018). This is very promising in cases where scaffolds can be modified to mimic disease conditions. Tissue engineering has been brought to focus as it deals with providing abilities that are closest to living tissue by providing a self-repairing network with the maintenance of blood supply and also the ability to respond to external stimuli such as stress or strain. As a native tissue, bones possess these abilities and thus the ideal graft material will have to meet or exceed the same to be adopted widely (Vallet-Regí and María González-Calbet 2004). To meet the goals of tissue reconstruction the candidate materials need to meet several criteria that have been covered earlier. With respect to bony reconstruction, there is a need to have ~60% pores with a size ranging from approx. 150 to 400 μm with around ~20% smaller than approx. 20 μm (Hollister 2005; Karageorgiou and Kaplan 2005; Shao et al. 2015). One key factor here is the ability of a porous scaffold to modulate its resorption in line with de novo tissue formation. This is essential as tissue forms and creates its support structures there is a gradual transition of the load to the developed tissue, this happens over a few months to about 3 years. In this route of creation of tissues cells and signals have played a yeoman role. BMP has been delivered via carrier units to local sites to expedite repair or in severe cases fix non-unions. A combination of bone morphogenetic protein loaded on coral or ceramic dishes was identified to be the best delivery vehicle (Gao et al. 1996). The superior tissue ingrowth and lack of fibrous tissue intercalation were observed by early 1997 in hydroxyapatite–collagen–BMP composites (Asahina et al. 1997). Several more materials were attempted over time ranging (Sakou 1998) from poly (lactic acid)–poly(lactic-co-glycolic acid) copolymers (Miyamoto et al. 1993), the pore size of delivery scaffolds (Tsuruga et al. 1997), the ability for ectopic bone regeneration (Whang et al. 1998), and polyhydroxyalkanoates (Croteau et al. 1999). BMP loaded HAp (Tsuruga et al. 1997; Koempel et al. 1998; Takahashi et al. 1999) and in time tested combinations as in collagen–HAp (Asahina et al. 1997) along with

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Table 3.3 Tracks major developments as well as commercial acceptance of graft materials. This provides information on the biomedical acceptance and marketability of the products Year 1920

1934 1936 1965 1969 1970 1971 1973 1975–1979 1975–1982

Major development Use of an aqueous slurry of “triple calcium phosphate”a to stimulate bone growth Use of tricalcium phosphate, MCP, and DCP slurries to stimulate bone growth Polyphosphates discovered in yeast Apatite precursor phase, Posner cluster Synthesis of dense HAp for prosthetic applications Importance of macropores for bone regeneration Implantation of “degradable” tricalcium phosphate ceramic in rats CaP-mediated transfection Clinical study with β-TCP and HA

1980–1987

First commercial CaP products: “Synthograft/Synthos” (β-TCP, 1975), Ceros Hap (HA, 1980), Durapatite (HA, 1981), ProOsteon (HA, 1981), Calcitite (HA, 1982), Alveograf (HA, 1982), Ceros TCP (β-TCP, 1982), BioBase (α-TCP, 1982) Description of the hydraulic properties of α-TCP CaP coatings

1982–1987

CaP cements (CPCs)

1985–1990 1985

CaP used as carriers for drug delivery Importance of micropores for bone regeneration Injectable/non-setting pastes (“putties”) Osteoinductivity Bone augmentation Production of HapWhiskers by hydrothermal synthesis Clinical study with CPC, commercial launch of Norian SRS and BoneSource Production of CaP scaffolds by rapid prototyping Si-substituted HA Polymer-induced liquid precursor (PILP)

1976

1987–1999 1990–1991 1992–1999 1994 1994–1995 1997 1999 2000

References Albee (1920)

Park (2009) Macfarlane (1936) Eanes et al. (1965) Levitt et al. (1969) Hulbert et al. (1970), Klawitter and Hulbert (1971) Bhaskar et al. (1971a, b) Graham and van der Eb (1973) Denissen and de Groot (1979), Roberts and David Brilliant (1975) –

Monma and Kanazawa (1976) Ducheyne et al. (1980), De Groot et al. (1987) Nicholson (2020), Pietrzak (2008), Lemons (1987) Otsuka et al. (1990) Klein et al. (1985) Malard et al. (1999), Dupraz et al. (1999) Yamasaki (1990), Ripamonti (1991) Bohner (2007), Bai et al. (1999) Yoshimura et al. (1994) Kamerer et al. (1994), Constantz et al. (1995) Levy et al. (1997) Gibson et al. (1999) Gower and Odom (2000) (continued)

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Table 3.3 (continued) Year 2001–2004

Major development Biomimetic CaP scaffolds, macroporous CPC

2002–2008

β-TCP synthesis by precipitation in hydrothermal conditions or in organic liquids Micronization/amorphization by milling Ready-to-use CPCs, dual-paste CPCs

2003 2003 2003–2004 2004–2006 2005 2005–2013

Flame-synthesized CaP nanoparticles Re-discovery of the importance of micropores for bone formation

2005–2007

3D printing of CaP scaffolds

2008

Nano-particulate apatite paste as bone substitute New ca–mg phosphate phase diagram ACP found in evolving bone Use of Ca and phosphate ions as drugs (bioinorganics) Validation of the PILP model Protein-free template mineralization Covalent functionalization of CaP nanoparticles Detailed description of ACP formation in vitro

2008 2008 2010–2011 2010 2012 2012 2013 a

Custom-made CaP nanoparticle for gene delivery (transfection) Hydrated layer on apatite crystals

References Almirall et al. (2004), Bohner (2001), Takagi and Chow (2001), Barralet et al. (2002) Bow et al. (2004), Toyama et al. (2002), Tao et al. (2008) Gbureck et al. (2003) Takagi et al. (2003), LeMaitre et al. (2008) Roy et al. (2003), Schmidt et al. (2004) Cazalbou et al. (2004), Jäger et al. (2006) Loher et al. (2005) Malmström et al. (2009), Polak et al. (2013), Woodard et al. (2007), Lan Levengood et al. (2010), Mayr et al. (2013), Bernstein et al. (2013) Gbureck et al. (2007a, b), Seitz et al. (2005), Leukers et al. (2005), Gbureck et al. (2007a, b) Kilian et al. (2008) Carrodeguas et al. (2008) Mahamid et al. (2008) Habibovic and Barralet (2011), Habibovic et al. (2010) Nudelman et al. (2010) Wang et al. (2012) Kozlova et al. (2011) Habraken et al. (2013)

Most likely an apatite powder with CDHA composition

ceramic composites of HAp/TCP origin (Ono et al. 1995) have proven to be useful in limited use cases. They have accelerated bone formation but with concerns regarding resorption and also of inflammation when collagen is combined with apathies used as carriers for osteoinduction factors (Doll et al. 1990). Loading characteristics have also changed over time with the use of microspheres of nHAp to adsorb and deliver bioactive BMP-2, with exceptional improvement reported in a rat femoral defect model (Zhou et al. 2018). Use of composite calcium phosphates with α-TCP and tetracalcium phosphate powders used as precursors to prepare HAp spheres with crystals of needle-like morphology which loaded with BMP was tested in a rabbit

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vertical guided bone regeneration model with encouraging enhancement of bone regeneration (Baek et al. 2016). Composites of HAp and polylactide fibers loaded with BMP-2 were demonstrated to generate more new bone at weeks 4 and 12 indicating compatibility of the material in various configurations (Xu et al. 2020). A comparative study by Alam et al. (2001) of five different ratios of hydroxyapatite to β-TCP along with varying doses of (rh)BMP-2 in a cranial defect in adult male Wistar rats was conducted. Evaluation after 8 weeks indicated that bone, bone marrow formation, and degree of resorption of ceramics particles were increased in samples of 25% HAp and 75% TCP. Material properties were amplified with the growth factor adsorbed onto the material surface. Carbonate derive apatite is one of the burgeoning interest, combining it with collagen and basic fibroblast growth factor and recombinant human BMP-2 and testing in a rabbit model indicated combination scaffolds with signals to provide superior activity (Salim and Ariani 2015). Influence of materials in the signaling pathways was explored by several workers (Barrère et al. 2006). Use of hydroxyapatite nanomaterials and detailed studies indicated that HAp nanoparticles of all sizes could enhance differentiation of hMSCs towards osteoblastic lineage with increased weightage for 50 and 100 nm sized materials. This was reflected in the increased ALP activity, ALP staining, immunofluorescent staining for osteopontin, and real-time PCR analysis (Yang et al. 2018). This effect on the ability to guide cell development has been utilized in apatite scaffolds for cell delivery of various origin in bone defect repair. The use of HAp nanoparticles in combination with poly (3-hydroxybutyrate-co-3-hydroxyvalerate) was able to improve ALP activity, osteocalcin levels, and exhibit significant effects on the repair of critical bone defects in a rabbit model (Lü et al. 2013). Synthetic origin as well as natural origin apatite has been tried for fine-tuning application areas. Apatite derived from fish scales via thermal decomposition has indicated cytocompatibility and mechanical utility based on the processing routes (Mondal et al. 2016). Porous scaffolds based on the sponge replication process were applied to develop scaffolds with higher biological activity owing to higher protein absorption, with the response of SaOS2 cells evaluated using multiple techniques (Tripathi and Basu 2012). There has been a trend as various cell types were used to populate scaffold systems and elicit the desired response. Emphasis has been on mesenchymal stem cells with their versatile nature ensuring osteogenic ability in high demand. Currently, adipose-derived stem cells (Rao et al. 2013), embryonic stem cells (Mahmood et al. 2012), umbilical cord blood endothelial cells (Baba et al. 2013), fallopian tubes (Jazedje et al. 2012), and dental pulp cells (Gamie et al. 2012) are in use based on the target niche. Murine models were used to demonstrate the osteogenic ability of a porous HAp scaffold system, this was pre-seeded and successfully carried out ectopic bone formation (Yoshikawa and Myoui 2005). Preconditioning of cells was highlighted in a study by Chai et al. (2012) wherein osteogenic differentiation of hMSC on a CaP scaffold influenced downstream events. This brought to light the need to completely understand all parameters in vitro culture, proliferation, collagen production, and osteoclast action before implantation in existing in vivo models.

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Clinical Vignettes

Clinical trials of cell-loaded calcium phosphate ceramics are few with human subjects. Quarto et al. (2001) reported on the management of large (4–7) cm bone defects of the long bones (n ¼ 3) where conventional interventions had failed to have the desired result. Custom designed scaffolds loaded with in vitro expanded autologous bone marrow stromal cells were used to bridge the defects (Fig. 3.7). In all radiographical imaging over a span of 2 months indicated integration at the interfaces and abundant callus formation (Quarto et al. 2001). Vacanti et al. went on to report the replacement of the distal phalanx of the thumb with a natural coral that was treated (porous HA; ProOsteon) in vitro by seeding with autologous periosteal cells. This restored functionality and resulted in the resumption of normal biomechanics of a normal thumb without the complications associated with bone grafting (Vacanti et al. 2001). Treatment of bone tumors post-removal by using scaffolds containing mesenchymal stem cells differentiated into osteogenic lineage on hydroxyapatite scaffolds to fill the defects was attempted. The study reported no adverse findings with rapid integration as assessed by radiography (Morishita et al. 2006). Cell-seeded apatite-based grafts have been demonstrated to be effective than the autograft, allograft, or cell-seeded allograft in ectopic bone formation (Eniwumide et al. 2007). A combination of dental follicle cells and apatite ceramics may be applied to restore periodontal defects (Zuolin et al. 2010). Bone repair in vivo with expansive bone regeneration results has been observed by combining human periodontal ligament stem cells on an apatite coated polymeric scaffold (Ge et al. 2012). The technique has been successfully used in alleviating postseptic gap nonunion of 4 cm using a customized hydroxyapatite tricalcium phosphate tricalcium silicate composite loaded with autologous bone marrow-derived stem cells primed for osteogenic differentiation. Union was recorded at 3 months, with improved joint mobility at 3 years with radiographical evidence of graft incorporation (Ge et al. 2012). There is a widespread use of the techniques mentioned above in high-performance veterinary orthopedics as well as to aid animal bone healing in defect regions (Franch et al. 2006; Vertenten et al. 2010). Three-dimensional printing (3D) is an additive manufacturing process that generates materials from 3D model data via varying technologies (Fig. 3.8).

Fig. 3.7 (a) Monolithic hydroxyapatite cylinder with a central canal similar to materials developed for first in the class clinical trial of bone tissue engineering in India. (b) Cross-section of the cylinder demonstrating central canal—for nail or support structure placement as required

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Fig. 3.8 Composite view of 3D printed ceramic constructs: (a) Optical images—ceramic bodies post-sintering are imaged, (b) X-ray evaluation of internal structures using a digital X-ray analysis unit, AGFA CR 10X courtesy DIMT, BMT Wing, and (c) scanning electron micrographs of crosssection revealing porosity and strand placement. At 100X porosity between strands and inside strands is clear. Post-sintered ceramic body has been cracked in LN2 and visualized using an FEI QUANTA 200

Ceramics, polymers, composites, and metals are all accessible via this technique (Hwang et al. 2015; Ladd et al. 2013; Ligon et al. 2017; Sun et al. 2013). A widely adopted technique for developing apatite scaffolds has been the printing of the 3D scaffold followed by sintering to drive out binder materials and confer required properties (Shao et al. 2016). A bioglass/β-TCP green body was prepared with a dextrin binder and post-printing with high-temperature sintering was carried out by Seidenstuecker et al. (2019). Cryogenic processes have been adopted to generate biomimetic hierarchical and interconnected porous apatite structures (Song et al. 2019a, b). Non-sintered low-temperature approaches allow for the loading of the scaffolds with cells and labile signals. Fabrication of nanobiphasic calcium with polyvinyl alcohol combined with enriched fibrin has been reported (Song et al. 2018). This enhances cell response as well as amplified bone regeneration in a segmental bone defect model in rabbits. Incorporation of BMP-2 via PCL emulsion technique on a 3D printed HAp scaffold for promoting the slower release of BMP-2 has been reported (Kim et al. 2018). The clinical application of this method has been reported with the use of bioactive calcium phosphate-based reconstruction of cranial defects (Engstrand et al. 2015). This covers a small trial wherein successful closure was achieved, with the necessity of a wider trial highlighted. Small scale studies indicate its utility currently in acute cases and also in the future in cases of trauma or high-velocity injuries requiring reconstruction (Engstrand et al. 2014). Reconstruction has also been attempted using granular beta-tricalcium phosphate materials in conjunction with adipose-derived stem cells with mixed results (Thesleff et al. 2017).

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Summary and Outlook

The evolution of this field will be led strongly by the basic nature of the materials under this class and their inherent versatility. Modifications to bulk properties, surface, or anchorage addition and development of innovative composites will lead this field in the future. Maturation of the field will be led by improving composite materials that lean heavily on cell signaling and component finesse rather than opting for the bluntness of a cell-loaded system. Further development of this field will be from the fruits of well-designed clinical and pre-clinical trials that strive to bring out the function-based material selection that is direly required at this point. This can also be tied to the tendency to lump all apatite-based materials under a common head. Nano-level substitutions, ceramic–ceramic composites all bank on the characteristics thus noted to achieve functional excellence. A systemic approach that provides detailed information on materials that are passed into trials on a common platform allowing for exhaustive analysis on the biological response will serve to catalyze positive changes for the future.

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growth, structure and orientation of bone apatite. Nat Mater 11(8):724–733. https://doi.org/10. 1038/nmat3362 Whang K, Tsai DC, Nam EK, Aitken M, Sprague SM, Patel PK, Healy KE (1998) Ectopic bone formation via RhBMP-2 delivery from porous bioabsorbable polymer scaffolds. J Biomed Mater Res 42(4):491–499. https://doi.org/10.1002/(SICI)1097-4636(19981215)42:43.0.CO;2-F Whited BM, Skrtic D, Love BJ, Goldstein AS (2006) Osteoblast response to zirconia-hybridized pyrophosphate-stabilized amorphous calcium phosphate. J Biomed Mater Res 76(3):596–604. https://doi.org/10.1002/jbm.a.30573 Williams D (2003) Revisiting the definition of biocompatibility. Med Device Technol 14(8):10–13 Woodard JR, Hilldore AJ, Lan SK, Park CJ, Morgan AW, Eurell JAC, Clark SG, Wheeler MB, Jamison RD, Wagoner Johnson AJ (2007) The mechanical properties and osteoconductivity of hydroxyapatite bone scaffolds with multi-scale porosity. Biomaterials 28(1):45–54. https://doi. org/10.1016/j.biomaterials.2006.08.021 Xu T, Sheng L, He L, Weng J, Duan K (2020) Enhanced osteogenesis of hydroxyapatite scaffolds by coating with BMP-2-loaded short polylactide nanofiber: a new drug loading method for porous scaffolds. Regen Biomater 7(1):91–98. https://doi.org/10.1093/rb/rbz040 Yamasaki H (1990) Heterotopic bone formation around porous hydroxyapatite ceramics in the subcutis of dogs. Jpn J Oral Biol 32(2):190–192. https://doi.org/10.2330/joralbiosci1965.32.190 Yang Y, Kim K-H, Ong JL (2005) A review on calcium phosphate coatings produced using a sputtering process—an alternative to plasma spraying. Biomaterials 26(3):327–337. https://doi. org/10.1016/j.biomaterials.2004.02.029 Yang X, Li Y, Liu X, Zhang R, Feng Q (2018) In vitro uptake of hydroxyapatite nanoparticles and their effect on osteogenic differentiation of human mesenchymal stem cells. Stem Cells Int 2018:2036176. https://doi.org/10.1155/2018/2036176 Yoshikawa H, Myoui A (2005) Bone tissue engineering with porous hydroxyapatite ceramics. J Artif Organs 8(3):131–136. https://doi.org/10.1007/s10047-005-0292-1 Yoshimura M, Suda H, Okamoto K, Ioku K (1994) Hydrothermal synthesis of biocompatible whiskers. J Mater Sci 29(13):3399–3402. https://doi.org/10.1007/BF00352039 Zhou P, Wu J, Xia Y, Yuan Y, Zhang H, Xu S, Lin K (2018) Loading BMP-2 on nanostructured hydroxyapatite microspheres for rapid bone regeneration. Int J Nanomed 13:4083–4092. https:// doi.org/10.2147/IJN.S158280 Zuolin J, Hong Q, Jiali T (2010) Dental follicle cells combined with beta-tricalcium phosphate ceramic: a novel available therapeutic strategy to restore periodontal defects. Med Hypotheses 75(6):669–670. https://doi.org/10.1016/j.mehy.2010.08.015

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Polymeric Biomaterials in Tissue Engineering: Retrospect and Prospects Lynda Velutheril Thomas

Abstract

Tissue engineering advancements have seen a multitude of findings in several disciplines, including cell biology, imaging, characterization of cell–material and cell–cell interactions, and also novel biomaterial research. The main aim of tissue engineering, however, remains as a tool to restore, maintain, or improve defective tissue functions. The paradigm of this concept is threefold: (1) Isolation of cells, (2) Seeding of cells into the appropriate 3D scaffolds, and (3) Providing the appropriate growth factors and physical and mechanical conditions in-vitro thereby mimicking the native conditions conducive for cell and tissue growth. The development of the 3D scaffold or matrix is by far the most challenging aspect wherein the choice of the scaffold material, its biocompatibility, cell– material interactions, its biodegradation and bioresorption properties, all play a major role. Polymers have been a mainstay as scaffold material for such applications. Both synthetic and natural polymers have been used as matrices for cell and tissue growth. The main aim in development of polymeric scaffold for tissue engineering is that it should resemble the properties of the tissues native extracellular matrix. A lot of advancements have been made in the last 10 years in the area of polymers used for tissue engineering applications and this chapter aims to provide a comprehensive coverage of the field. Keywords

Polymeric biomaterials · Scaffold · Extracellular matrix · Scaffold fabrication · Synthetic polymer · Natural polymer L. V. Thomas (*) Division of Tissue engineering and Regeneration Technologies, Biomedical Technology Wing, Sree Chitra Tirunal Institute for Medical Sciences and Technology, Trivandrum, Kerala, India e-mail: [email protected] # The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 B. Bhaskar et al. (eds.), Biomaterials in Tissue Engineering and Regenerative Medicine, https://doi.org/10.1007/978-981-16-0002-9_4

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Introduction

The use of polymeric biomaterials in medical applications has been expanding rapidly since the last two decades in various application regimes that include tissue engineering as well. Furthermore, a lot of progress has been made in associated areas of technology related to polymeric biomaterial development for tissue engineering which includes microfabrication techniques, surface modifications, drug or growth factor delivery systems, nanotechnology, etc. The tissue engineering paradigm mainly involves—cells, scaffold, and growth factors or cues for maintaining cell function and growth. The scaffold is the framework provided for the tissue formed to attain their 3-dimensional shape ideally mimicking the native extracellular matrix. The unique features of the ECM and its role in tissue development and growth are discussed in this chapter. The basic requirement for a polymeric biomaterial as a scaffold for tissue engineering is that the biodegradation pattern meets the time point of tissue regeneration. Consequently, majority of the polymeric scaffolds used in tissue engineering are biodegradable. Such biodegradable polymeric biomaterials can be classified according to their origin as natural and synthetic; which has been elaborated. In developing scaffold structures for tissue engineering, a wide array of parameters needs to be looked into depending on the site of implantation, the cell and tissue milieu in the native site, the mechanical strength required, the degradation profile and its by-products, the ease of removal from the site on degradation, etc., which will be unique for each tissue under consideration. Furthermore, the fabrication methodology involved in the preparation of 3D scaffolds has also flourished, from microporous scaffold development to nano porous scaffolds and from self-assembled scaffold structures to more controlled and reproducible rapid prototyping techniques of scaffold development. The various properties of the polymeric scaffolds used as tissue engineering scaffolds and their prospects and retrospects have been explored in this chapter; which will enable us to understand the contribution of each polymer towards the regeneration of tissues via the tissue engineering approach.

4.2

Extracellular Matrix—the Framework Enabling Tissue Growth

The major constituents of tissue architecture are cells and extra cellular matrix (ECM) that controls and regulates several tissue functions. The ECM is a collection of different macromolecules that provides the framework for cells to adhere, grow, and proliferate. The main composition of ECM includes water, minerals, structural proteins, specialized proteins, and proteoglycans of which the fibrous proteins and the proteoglycans play a major role (Fig. 4.1). The ECM is a dynamic structural matrix that is constantly being remodeled, synthesized, and modified by the cellular components that they support (Teti 1992; Kleinman et al. 2003). Cell adhesion, proliferation, differentiation, and apoptosis are all controlled by the cell nucleus– ECM interaction. Hence recreating the ECM structure has been considered as the

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Fig. 4.1 Schematic representation of the basic composition of the extracellular matrix

first step when regenerating a tissue. Recreation of the structural variability of ECM using various polymers has been attempted. The major challenge is to initiate the cell matrix interactions through these polymeric matrices which has been attempted by surface modifications using cell signaling molecules, growth factors, etc. Moreover, the cell adhesion, infiltration, and growth also depend on many of the physicochemical properties of the matrices including pore sizes, porosity, pore tortuosity, bioactivity, stiffness, etc. This is where a lot of research has been performed in the area of natural and synthetic polymer fabrication to create ideal scaffolds for tissue engineering. The in vivo degradation pattern of the polymers used as matrices also plays a major role when regeneration of a tissue is of concern as the rate of degradation should ideally be in par with the formation of new tissue without eliciting any inflammation or immune response due to degradation by-products. This is where the type of polymers selected, the fabrication methodology, the surface modification and the mechanical stability, all determine the suitability of the scaffold for specific tissue repair and it all depends on the anatomy and function of the three-dimensional tissues.

4.3

Polymeric Materials as Ideal Scaffold

The scaffold is the major structural component in tissue engineering that provides the base mechanical and structural properties of the native tissue. The concept of an ideal scaffold has been explored and laid out by researchers. Any structure that is intended to be used as a matrix for tissue regeneration should consist of the following properties (Hutmacher 2001; Yang et al. 2001) (Fig. 4.2). 1. Be non-thrombogenic. 2. Allow ease of implantation with proper tissue integration would be required.

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Optimum porosity enabling cellular ingrowth

Desireable tensile strength

Sterilizable Non Thrombogenic with proper tissue integration

Appropriate level of biostability with good chemical and mechanical stability

Low rate of infection when implanted

Biocompatible and biodegradable

Fig. 4.2 Schematic showing the different characteristic requirements of an ideal scaffold (Some images adapted from Thomas et al. 2012)

3. 4. 5. 6.

Be biocompatible, without eliciting any immune response and be non-cytotoxic. Should be easily sterilized with minimal effects on the developed scaffold. Once implanted, the scaffold should have low infection rate. Have the appropriate tensile strength matching the host tissue. The mechanical properties and the degradation profile should be in line with the kinetics of the developing tissue which will ensure proper tissue integration. 7. The scaffold should have appropriate values of fatigue endurance and biostability and should retain its chemical and mechanical properties such as compliance and elasticity during use. 8. Should possess the appropriate level of porosity enabling cellular ingrowth with efficient diffusion and transport of nutrients and waste.

There are different types of polymers that have been used as scaffolds for tissue regeneration. The properties of the polymers are mainly defined by its composition, molecular weight, structure, etc., and can be categorized broadly as naturally occurring polymers and synthetic polymers. Historically, the biomaterials that were used clinically were mostly based on the natural polymers. Their ideal properties include their natural structural entity that mimics the extracellular environment, their biodegradability with nontoxic by-products, and enhanced cell responsiveness and good tissue integration. Some examples of natural biopolymers used as scaffolds for tissue regeneration include polysaccharides (like chitosan, cellulose, dextran, gellan gum, xanthan gum, pullulan, etc.); proteins (collagen, gelatine, silk fibroin, keratin, fibrin, etc.); and polynucleotides (like DNA, RNA). The major drawbacks of the use of

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such natural polymers as scaffold materials include their immunogenicity, reproducibility in terms of structure and functional properties, and low mechanical strength. To improve upon the properties of these natural biopolymers several modification techniques like conjugation and blending with other polymers have been performed. Some of the natural polymers used in the area of tissue engineering are discussed in detail in further sections. Synthetic polymers have been used in several medical devices. The synthetic biodegradable polymers have found a lot of advantage in the area of tissue engineering as robust scaffold systems with good mechanical properties similar to or even better than the native tissues, better control over the degradation pattern, easy process ability through several fabrication routes owing to their thermal transitional properties, etc. They are also seen as more cost effective with longer shelf life. Synthetic polymers present a bigger class of biodegradable polymers which can be produced under controlled conditions thereby maintaining its reproducibility in function. Some examples include PCL, PLA, PLGA, PGA, PHA, etc. However, some of the advantages of synthetic polymers used as scaffolds for tissue engineering applications include their limited cell responsive surfaces, greater degradation time with toxic by-products in some polymers, and poor bioactivity. Some examples of such synthetic polymers and their use in the tissue engineering arena have also been discussed in detail in the next sections. The combination of degradable polymers with inorganic and bioactive materials, blend systems of natural and synthetic polymers, chemical modifications of natural and synthetic polymers has also been explored as a strategy to obtain better mechanical and biological performance.

4.4

Natural and Synthetic Polymers as Scaffolds

In terms of bioabsorbability and biostability, polymeric biomaterials can be divided in to two groups • Biodegradable polymer • Non-biodegradable polymers Biodegradable polymers which are the most preferred candidate as a scaffold for tissue engineering can be further classified in terms of their origin as natural and synthetic. Non-biodegradable polymers are not usually applied in the area of tissue engineering as it do not meet the criteria of an ideal scaffold and do not degrade in the physiological environment. Non-biodegradable polymers include poly (amido amine), polysulfones, certain polyurethanes, poly (ethylene oxide), polyphosphazenes, polyamides, poly (ethylene imine), etc. Most of these polymers have very good mechanical strength and have good processability. However, their use is limited owing to their insufficient biodegradability. These issues have been overcome to some extent by blending these polymers with biodegradable synthetic

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and natural polymers, by preparing composite structures that can enable controlled degradation, or by modifying the polymers by introducing more labile and degradable moieties as easy degradation sites. In this chapter we will focus more on the biodegradable natural and synthetic polymers which play a major role in the scaffold regime in tissue engineering.

4.5

Natural Biodegradable Polymers

Some of the natural polymers used as scaffolds for tissue engineering have been discussed below.

4.5.1

Collagen

Collagen, being one of the major extracellular proteins, is considered to be an ideal biomaterial to be used as a scaffold for tissue engineering. Collagen matrix gives the structural framework for many connective tissues and also has integrin like sequences within its network that helps in cell–matrix interaction and supports cell adhesion, migration, proliferation, differentiation, and growth. The shape and structural properties of a native collagen molecule are established as a right-handed bundle of three parallel, left-handed polyproline II-type triple-helical α-domains. There are twenty-seven types of collagen that have been identified with Collagen type I being the most explored as a biopolymer for scaffold development and other biomedical applications (Yang et al. 2017). The low mechanical strength of collagen is being compensated by the incorporation of biodegradable synthetic polymers like PGA, PLA, P(LLA-CL), etc. (Torikai et al. 2008; Park et al. 2009). Collagen forms the major class of proteins found in the extracellular matrix of animals which is structurally composed of three polypeptide chains. Three separate alpha polypeptide chains are present in the triple helical domain structure of collagen, which is also called tropocollagen. This tropocollagen is involved in the collagen formation in the extracellular matrix (Gelse et al. 2003). The collagen biopolymer used as scaffolds is normally extracted from the tendons, skin, bone, and cartilage of animals. Collagen has a high biodegradation profile owing to the fast degradation by enzymes, good biocompatibility, with minimal effects on the immune system and good cell adhesion and growth (Lynn et al. 2004; Malafaya et al. 2007). It finds many applications in medical science including delivery systems, burn and wound healing (Liu et al. 2019), and tissue engineering (Cen et al. 2008). Collagen scaffolds have been widely explored for various tissue engineering applications (Chan et al. 2016; Irawan et al. 2018).

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Gelatin

Due to the batch to batch variability in the quality of extracted collagen, types, the problems related to antigenicity and its high cost, sources similar to collagen are being considered. This is where gelatin, which is the denatured form of collagen and contains many amino acids like glycine, proline, glutamic acid, hydroxyl proline, arginine, alanine, aspartic acid, etc., finds immense potential as a scaffold material. Furthermore, Gelatin evokes considerable interest on account of its biodegradability and amenability to modifications. It is basically found in two forms—type A, which is prepared by acid treatment of collagen, and type B, obtained by alkaline treatment. The Bloom value or gel strength of gelatin depends on the amount of intact collagen chains and the overall average chain length of the natural polymer. In comparison to collagen, it is also less immunogenic and cost effective. Gelatin has found wide range of applications in the pharmaceutical industry as well as in the biomedical field mainly in different forms like hard and soft capsules, microspheres, as sealants for vascular prostheses, as wound dressing and absorbent pad for surgical use, as scaffolds for tissue engineering of bone, skin, cartilage, etc. Gelatin has also been used after chemical modification or after blending with other biopolymers as scaffolds for tissue engineering (Liu et al. 2005; Witte and Kao 2005; Zhao et al. 2006). Its main limitation is the lack of mechanical property and fast biodegradation potential which in turn limits its potential applications as a biomaterial for many long-term applications. Crosslinking of gelatin has been found to improve both the thermal and the mechanical stability of the biopolymer and also its biodegradation profile (Marois et al. 1995; Bigi et al. 2001). Chemical modification of natural polymers by grafting has also been attracting interest over the years for getting better properties in terms of mechanical strength and controlled rate of biodegradation (Thomas and Nair 2012). Gelatin vinyl acetate porous scaffolds exhibited interconnected porous structure, the cell adhesion and colonization, extracellular matrix production covered throughout the scaffold, and cell alignment across walls of tube were clearly demonstrated by SEM images (Fig. 4.3). Mao et al. (2003) proposed the use of a novel hybrid of Chitosan–Gelatin as a matrix or scaffold for tissue engineering. Cross linked hydrogels of Gelatin and PVA have also been prepared which could be used as a promising scaffold structure. Sakai et al. (2007) synthesized a gelatin–agarose conjugate which was found to have suitable characteristics of a tissue engineered scaffold. Electrospun membranes of collagen, gelatin (denatured collagen), and solubilized alpha elastin also find potential application in the area of tissue engineering. These electrospun engineered protein scaffolds were seen to support attachment and growth of human embryonic palatal mesenchymal (HEPM) cells on performing cell culture experiments (Li et al. 2005). Another study by Chong et a. (2007) showed that PCL-gelatin nanofibrous scaffolds could be prepared which finds potential application in tissue engineering of skin (Chong et al. 2007).

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Fig. 4.3 The porous nature of freeze dried gelatin vinyl acetate copolymer (GeVAc) as evidenced through SEM images (a and b) showing the interconnected pore structure; (c) freeze dried tubular porous GeVAc construct; (d) Rat aortic smooth muscle cells seeded construct after 6-week culture completely covered with RASMC and its ECM; (e) magnified image of vessel wall (static culture); and (f) magnified image of the cells that are aligned in the vessel wall on application of mechanical stimulation. The GeVAc is a promising scaffold material finding immense potential in the area of blood vessel regeneration (Adapted from Thomas et al. 2012)

4.5.3

Chitosan

Chitosan is a cationic polysaccharide which is sourced from chitin with basic units of glucosamine and N-acetyl-D-glucosamine. Chitin is a biopolymer which is obtained

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from the shell of crustaceans like shrimp and crabs, from squids and even from the cell walls of fungi (Khor and Biomaterials 2003). Chitosan which is the deacetylated form of chitin has found extensive application as a biomaterial owing to its solubility and cationic nature. It has also found application as a scaffold for tissue engineering mainly in the area of nerve, liver, and cartilage tissue, due to the similarity in structure to glycosaminoglycans present in the extracellular matrix. However, some of the drawbacks are the low mechanical strength and the solubility which affects the adhesion and interaction with cells. Hence, chitosan has been modified by blending and copolymerizing with other polymers to improve upon its strength, cell adhesion, and cytocompatibility. Combination by blending of chitosan with synthetic polymers like poly (vinyl alcohol), poly (ethylene glycol), poly (vinyl pyrrolidone), or natural polymers such as collagen, hyaluronic acid, gelatin, etc., has already been produced (Shakir et al. 2015; Kanimozhi et al. 2016; Casimiro et al. 2018; Ranganathan et al. 2019). Arca and Şenel (2008) emphasize on the use of chitosan-based biomaterials as scaffolds for blood vessel regeneration therapy.

4.5.4

Alginate

Alginates are natural biopolymers obtained from seaweeds and algae. Since these linear polysaccharides are obtained from marine sources, extensive purification is required to classify them safe to use in biomedical applications and also to prevent immune responses after implantation (Willerth and Sakiyama-Elbert 2007). However, these biopolymers have been shown to be cytocompatible and used extensively as a cell encapsulation agent for enhancing cell survival and growth. They have been used as a scaffold in liver, nerve, heart, and cartilage tissue engineering. As they are mainly used in the form of hydrogels, their major drawback is the low mechanical strength and poor cell adhesion. Hence, these polymers have been used in combination with other polymers to overcome these drawbacks (Kirdponpattara et al. 2015; Seok et al. 2019; Datta et al. 2020). Mohan and Nair (2005) prepared highly porous 3D scaffolds from sodium alginate, which exhibited good cytocompatibility. The scaffold was prepared by the process of freeze drying. The developed scaffold showed good porosity and swelling properties enabling cellular ingrowth and nutrient supply. Alginates have also found potential application in recent years as a bioink for tissue engineering owing to their viscoelastic properties (Axpe and Oyen 2016; Baena et al. 2019; Zhang et al. 2019). Naghieh et al.(2019) developed alginate scaffolds that were fabricated using an indirect-bioprinting process and characterized the potential of these scaffolds for nerve tissue engineering application and they observed better cell functionality on the scaffolds fabricated with a lower concentration of alginate compared to a higher concentration.

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Fibrin

Fibrin is a natural matrix and was first used as a sealant for wound healing and tissue repair. Fibrin is a biopolymer that is derived from the monomer fibrinogen. The fibrinogen molecule is composed of six disulfide bonds that help to bridge two sets of three polypeptide chains named Aα, Bβ, and γ. The fibrin biopolymer is formed after thrombin-mediated cleavage of fibrin peptides from their respective α and β chains enabling conformational changes and exposure of polymerization sites. The fibrin monomer formed in the process undergoes self-association to form insoluble fibrin. Fibrin clots provide a 3 dimensional framework for cell adhesion, proliferation and migration of cells, which is a major prerequisite in wound healing with remodeling and resorption through normal fibrinolytic processes. Apart from wound repair applications, it has also found advantages in the area of drug delivery systems, cell and growth factor delivery systems, and as a three-dimensional scaffold for cell growth and differentiation. The major limitation in the use of fibrin scaffold for tissue engineering application has been the poor mechanical strength and hence it has been incorporated with various biodegradable polymers like poly (L/D-lactide) and poly caprolactone for enhancing the mechanical strength (Osathanon et al. 2008; Pankajakshan et al. 2008; Tschoeke et al. 2009). Shaikh et al. (2008) reviews the use of fibrin as scaffold for blood vessel tissue engineering. Ahmed et al. (2008), in their review paper, highlight on the manipulation of fibrin for tissue engineering applications wherein the different forms of scaffolds including fibrin hydrogels, fibrin glue, and fibrin microbeads (FMBs) and their usability in different tissues are discussed. Furthermore, Puente and Ludena (2014) also explore the methods of development of alginate based systems and analyzed the commercial and autologous fibrins that are available, for their application potential in tissue engineering.

4.5.6

Hyaluronic Acid

Scaffolds based on hyaluronic acid have also found applications in the area of tissue regeneration. Hyaluronic acid is a naturally occurring non-adhesive glycosaminoglycan polymer associated with various cellular processes involved in wound healing, such as angiogenesis and found mostly in connective, epithelial, and neural tissue. Scaffolds made of hyaluronan gel crosslinked with divinyl sulfone promoted elastogenesis when neonatal rat aortic SMCs were cultured (Ramamurthi and Vesely 2005). Hyaff-11, a commercial hyaluronic acid-based biomaterial has found success as a scaffold for vascular tissue regeneration (Zavan et al. 2008). Although HA scaffolds have found application in both hard and soft tissue regeneration, most of the scaffolds are in the form of hydrogels. This is due to its good swelling characteristics and cellular encapsulation capability. Hemshekhar et al. (2016) have extensively reviewed on the pharmacological and pathophysiological properties of native and modified HA and its major clinical uses. The review also highlights the therapeutic applications of HA-based bioscaffolds in organ-specific

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tissue engineering and regenerative medicine. Pandit et al. (2019) present an overview of oxidized hyaluronic acid (OHA) based hydrogels as scaffolds and their potential application in the field of TE. Collins and Birkinshaw (2013) have extensively reviewed on the mechanical, biological function and degradation of HA and the different fabrication methodologies used for preparing HA scaffolds for tissue engineering.

4.5.7

Silk

Silks are naturally occurring polymeric proteins, extracted from Lepidoptera larvae (such as silkworms), some arachnids (spiders, mites, and some scorpions), and a few flies. It has excellent biocompatibility, good thermo-mechanical stability, and can be easily tailored using a wide plethora of fabrication technique and its biodegradation can also be tuned. The protein component of silk with a rich source of cell responsive peptides has been the main reason for its popularity as a scaffold for tissue engineering. Silk is composed of two major protein components—sericin which is a watersoluble protein that has adhesive properties and presents issues of antigenicity and fibroin, the fibrous part of silk which has a rich source of cell responsive peptides and has good mechanical stability. Silk can be tailored and fabricated to 3D scaffolds using different fabrication techniques to obtain scaffolds in the form of hydrogels, foams, nanofibers, and films for different tissue engineering applications. Correia et al. (2012) developed silk based scaffolds using different solvents and different pore sizes and explored using human adipose-derived stem cells (hASC) an attractive cell source for engineering autologous bone grafts. Bhattacharjee et al. (2017) have reviewed on the use of silk in bone tissue engineering. Silk composites have also been used to generate tissue specific properties. Electrospun nanofibrous silk fibroin (SF)/carboxymethyl cellulose (CMC)/nano-bioglass (nBG) composite scaffold with the appropriate ratios have been used as scaffolds for bone tissue engineering and fabricated by free liquid surface electrospinning technique (Singh and Pramanik 2018). Farokhi et al. (2018) reviewed on blending silk fibroin/hydroxyapatite and developing bone constructs for tissue engineering. Teimouri et al. (2014) developed novel composite scaffolds consisting of silk fibroin and forsterite powder by a freeze-drying method for bone tissue engineering application.

4.6

Synthetic Biodegradable Polymers

Synthetic biodegradable polymers have found vast applications in the medical implant industry. Many biodegradable synthetic polymers have been approved for medical use by the Food and Drug Administration (FDA), e.g. Poly(caprolactone) (PCL), polyL-lactide (PLLA), poly(lactide-co-glycolide) (PLGA), poly(vinyl alcohol) (PVA), poly(ethylene glycol) (PEG), poly (glycolide) (PGA), etc. Synthetic polymers do not possess structural surface characteristics and groups which are familiar to cells and enable cell attachment. These polymers can, however, be

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tailored to produce 3D scaffolds having a wide range of mechanical properties and degradation rates. Here, some of these synthetic biodegradable polymers used as scaffold material in tissue engineering application are described.

4.6.1

Poly Lactic Acid (PLA)

Poly lactic acid (PLA) is a polyester prepared using two monomers, namely lactic acid and the cyclic lactide monomer. Polycondensation reaction of lactic acid in the L and D stereoisomer forms is performed to obtain PLA polymer. PLA can also be obtained from lactide monomers by the ring opening polymerization reaction with metal catalyst like stannous octoate in solution or in suspension. L-lactic acid which is the degradation product of PLA is non-toxic and is usually involved in the metabolic pathway of all animals and microorganism. This non-toxic nature of polylactides has been made use of in resorbable medical sutures which have been in market for over three decades (Onose et al. 2008). Some of the applications in which PLA has been used include sutures, drug delivery vehicles, prosthetics, vascular grafts, etc. Sculptra™ is a commercialized FDA approved biomaterial based on PLA which is used as an injectable device that is used currently to treat facial atrophy (Blasi 2019). The degradation rate of PLLA is considered to be relatively slow. It is also highly crystalline in nature and hence these materials will be hard and brittle in nature. Due to the chirality of lactic acid, PLA exists in three enantiomeric states, L-lactide, D-lactide, and meso-lactide of which the first two enantiomers are most commonly used. However poly (D,L-lactic acid) PDLLA is amorphous due to the disruption of stereo-regularity owing to the random distribution of PLLA and PDLA. Both the polymers PLLA and PDLA have comparable mechanical properties (tensile strength (4–8 GPa), elongation at break (1–8%), and tensile strength values (40–70 MPa)). The degradation rate is also influenced by the chiral state of lactic acid in the polymer and the highly crystalline PLLA is seen to degrade completely in 2–5 years whereas the PDLLA is less stable and the degradation rate ranges from less than 2 months to 1 year which is proven by in-vivo studies. Since this polymer is biocompatible and biodegradable with good mechanical strength, it also finds application as a scaffold for tissue engineering. PLLA is often blended with other polymers to improve processability to obtain a 3D scaffold. Ju et al. developed an efficient and eco-friendly processing technique with supercritical carbon dioxide foaming to prepare porous PLLA/poly (ethylene glycol) (PEG) (95/5 wt%) scaffolds for bone tissue engineering (Ju et al. 2019). To test the efficacy of this scaffold, a rabbit model with bone defects was used and it was found that the obtained porous scaffold supported bone tissue engineering. Hu et al. (2010) studied the effect of nanofibrous poly-L-lactide (PLLA) scaffolds on phenotype control of human aortic smooth muscle cells (HASMCs) for vascular tissue engineering and found that these scaffolds preferentially supported contractile phenotype of HASMCs under the in-vitro culture conditions with in-vivo subcutaneous implantation studies confirming HASMCs differentiation in the implants. Jun Negishi and team investigated a vacuum pressure impregnation (VPI) method for creating a

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composite of natural polymer collagen with PLA (Negishi et al. 2019). PLA based biomaterials can be fabricated using different fabrication technologies like stretch blow molding, film casting, thermoforming, extrusion, fiber spinning, electrospinning, melt electrospinning, injection molding, foaming, and micro- and nano-fabrication techniques and can be formed into various shapes and sizes for use as scaffold structures in tissue engineering (Santoro et al. 2016; Grémare et al. 2018; Zhou et al. 2018).

4.6.2

Poly (glycolic acid) (PGA)

PGA is highly crystalline polymer with high melting point in the range of 225–230  C and low solubility in organic solvents. Since the polymer has high crystallinity the polymer is blended with other polymers to improve on the processability. Breuer et al. (2008) developed an autologous vascular graft where they used non-woven PGA meshes in the tubular form after coating with a 10% solution of 50:50 L-lactide and caprolactone and then cultured it with autologous bone marrow derived mononuclear cells (Breuer et al. 2008). Iwasaki et al. (2008) prepared tubular constructs using PGA and PCL seeded with SMCs and a PGA sheet alone seeded with fibroblasts. After implantation, the ester bonds underwent degradation and these by-products were seen to be absorbed by the body, however, differences in pH were observed around the site of implantation.

4.6.3

Poly (lactic-co-glycolic acid) (PLGA)

Poly (lactic-co-glycolic acid) (PLGA) is prepared by the ring opening copolymerization reaction of the two monomers—(1,4-dioxane-2,5-diones) of glycolic acid and lactic acid, in both random and block copolymerization using metal catalysts at high temperatures (130–220  C), including aluminum isopropoxide, tin (II) 2-ethylhexanoate, or tin (II) alkoxides. The degradation rate of the scaffolds can be tailored and controlled by varying the glycolide to lactide ratio during the polymerization reaction (Park 1995). PLGA can be solubilized in a wide range of common solvents, including tetrahydrofuran, acetone, or ethyl acetate and chlorinated solvents and it can be tailored and fabricated to different sizes and shapes using different techniques like solvent casting, 3D printing, phase separation, freeze drying, electrospinning, etc. The structure can also support the incorporation of a wide range of biomolecules into the polymer and thus finds potential application in the area of tissue engineering. Furthermore the PLGA is available as D-, L-, and  D, L-isomers and the Tg values are greater than 37 C and hence is glassy in nature. As the lactide content decreases the Tg value also decreases. Since PLGA has ester linkages the major degradation mechanism is via hydrolysis. The biodegradation pattern of PLGA is observed from one to six months depending on the ratio of lactide to glycolide (Holy et al. 1999). The degradation by-products include the release of lactic acid and glycolic acid which results in the generation of acidic environment

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around the tissue scaffold which is a major concern when used as a scaffold for tissue engineering. Nevertheless, this copolymer has found a wide range of applications in both native and modified forms as films, porous scaffolds, hydrogels, etc., and has been modified by blending with other biocompatible polymers (Han et al. 2011; Pan and Ding 2012; Gentile et al. 2014). The major application using this polymer has been in the area of bone tissue engineering. Jazi (2017) prepared biodegradable PLGA nanocomposite scaffolds incorporated with nano-crystalline zeolite powder (7 (wt%)) by electrospinning and showed that these scaffolds had potential application in the area of bone tissue engineering (Jazi 2017). Yun et al. (2009) studied the osteogenic differentiation of various stem cells on electrospun PLGA/nano-HA taken in a 5:1 blending ratio. They used primary human adipose tissue-derived stem cells (hADSCs) and human bone marrow cells (MSCs) in their study. To improve upon the processability of PLGA, blending of PLGA with a plasticizer, such as poly (ethylene glycol) (PEG) is reported which helped to reduce the Tg to 37  C (Dhillon et al. 2011). The PLGA/PEG particle on mixing with a carrier solution makes the polymer soft enough to be molded to the required shape and then subsequently harden into a scaffold at 37  C. Liang et al. (2018) examined the osteochondral regeneration potential of a composite with one layer made of biodegradable polymer poly(d,l-lactide-co-glycolide) (PLGA) and another layer made of a PLGA-hydroxyapatite (HAp) matrix. Kim et al. (2020) reported that 50% polyoxalate POX/PLGA film can be applied in different tissue engineering fields including bone tissue engineering and drug delivery applications. Porous scaffolds via 3D printing for bone tissue engineering applications have been developed using a 10:1 weight ratio of poly lactic-co-glycolic acid (PLGA)/TiO2 composite and found that these scaffolds significantly improved osteoblast proliferation compared to pure PLGA with significantly higher ALP activity and calcium secretion (Rasoulianboroujeni et al. 2019). Ju et al. (2019) developed gelatin microspheres loaded poly(lactic-co-glycolic) acid (PLGA) scaffolds (PLGA/GMs scaffold) for enhancing osteogenesis. A sustained release property of recombinant human bone morphogenetic protein-2 (BMP-2) was also achieved in BMP-2-releasing PLGA/ GMs scaffolds that were developed. Sahoo et al. (2010) coated bioactive bFGFreleasing ultrafine PLGA fibers over knitted microfibrous silk scaffolds. Sustained release of bFGF was observed which mimicked the function of ECM. This helped to initiate stimulation of proliferation of mesenchymal progenitor cells (MPC), and subsequently, their tenogeneic differentiation. Ferreira et al. (2019) prepared nontoxic, hydrophilic, and porous polymeric conduits based on poly (lactic-co-glycolic acid) (PLGA), polycaprolactone (PCL), and polypyrrole fibers (PPy) for regenerating peripheral nerves. A bilayered poly (lactic-co-glycolic acid) (PLGA) scaffold has also been fabricated having small (200–300 μm) and large (200–500 μm) pores by salt leaching technique and evaluated for chondrocyte differentiation, cartilage formation, and endochondral ossification. The scaffold surface was also modified with tyramine and implanted in vivo into porcine osteochondral defects to promote scaffold integration (Lin et al. 2019).

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Poly(caprolactone) (PCL)

Poly(caprolactone) is one of the most promising polymeric biomaterials that has been used as scaffold for regenerative medicine. PCL is an aliphatic polyester synthesized via the ring opening polymerization of ε-caprolactone. The degradation of PCL is very slow owing to the five –CH2 groups on the aliphatic repeat chains and the degradation is mainly through hydrolytic cleavage of ester bonds and, therefore, has received a lot of focus as a material for use as an implantable biomaterial (Schnell et al. 2007). PCL is a Food and drug Administration (FDA) approved biomaterial used in medical applications as a drug delivery device, sutures, and adhesion barrier. A recent advance saw the emergence of 3D printed PCL scaffolds as a bone void filler for craniofacial application which obtained the 510 (K) FDA clearance in 2006. PCL has found a range of applications as scaffolds for tissue engineering application which spans a range of tissues from bone and cartilage to soft tissues like blood vessel and skin owing to the ease of fabrication. PCL has a low glass transition temperature of 60  C and exists in rubbery state at room temperature. It also has a high thermal stability with a decomposition temperature of 350  C. Hence this polymer can be tailored to different 3D scaffolds using different fabrication techniques like solvent casting, electrospinning phase separation, FDM, 3D printing, and extrusion methods. However, one of the drawbacks is the issue of hydrophobicity. This may be addressed by surface modification like functionalization and synthesis of blend polymerization with hydrophilic polymers and formation of block copolymers. Dong et al. (2017) developed a 3D printed PCL scaffold incorporating cell seeded chitosan hydrogel for bone tissue engineering applications. Christiani et al. (2019) developed multi-layer PCL scaffolds by depositing PCL struts via 3D printing in opposing angular orientations of 30  , replicating the angle-ply arrangement of the native annulus fibrous tissue. Another research group studied the process of supercritical foaming of polycaprolactone and polycaprolactone/graphene composite in a carbon dioxide atmosphere for bone tissue engineering applications (Evlashin et al. 2019). Pektok et al. (2008) produced biodegradable small diameter scaffolds made of PCL nanofibers using electrospinning and reported better healing characteristics and more stable neointima formation compared with ePTFE grafts.

4.6.5

Poly Vinyl Alcohol (PVA)

Poly(vinyl alcohol) (PVA) is an FDA approved material that has been used in several biomedical applications [Li et al. 2012]. PVA finds a lot of applications as drug delivery devices, soft contact lenses, membranes for bioseparation and hemodialysis (Burczak et al. 1994; Yon et al. 1994). PVA is synthesized by the polymerization of vinyl acetate monomers into poly (vinyl acetate) which involves the hydrolysis of the acetyl groups into poly (vinyl alcohol). Because of the extensive hydroxyl moieties the material is hydrophilic. They are also cytocompatible and can be easily tailored using different fabrication techniques. They have also found potential

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applications as tissue engineering matrices especially for the soft tissue segments. Moreover, this material has found to have excellent mechanical properties and is flexible. The semi permeable matrices allow for transport of oxygen and nutrients which is a prerequisite for cell survival and growth (Saavedra et al. 2003). Matrices of PVA in the form of hydrogels are most commonly used. Crosslinked PVA hydrogel matrices using photocrosslinking strategies have also been developed in recent years (Schmedlen et al. 2002). One such photocrosslinked scaffold poly (lactic acid)-g-PVA hydrogel for tissue engineering of heart valves was developed (Nuttelman et al. 2002). Lee et al. (2005) developed a PVA chondroitin sulfate hydrogel as a matrix for tissue engineering. PVA has also been used as blend systems with other polymers for improved mechanical stability and cytocompatibility (Oh et al. 2003). A semi-IPN blend of PCL/PVA was developed by Mohan and Nair (2008) for cartilage tissue engineering applications. Electrospun membranes based on PVA have also been attempted. Asran et al. (2010) developed electrospun polyvinyl alcohol and poly(hydroxyl butyrate) scaffolds for skin tissue engineering. Blend systems with natural polymers to improve upon cell–material interaction have also been developed like the electrospun PVA/Gelatin composite system for bone tissue engineering (Linh and Lee 2012). Since PPVA does not contain any functional groups that aid in cell attachment apart from the hydroxyl groups (-OH) groups, modification using bioconjugation techniques with specific peptides to enhance biomimeticity of PVA has also been explored. Citric acid modified PVA scaffolds using the technologies of freeze drying and electrospinning have also been developed and their performance evaluated as construct for vascular tissue engineering (Thomas et al. 2009; Thomas and Nair 2019). The molecular weight and the degree of hydrolysis are the major contributory factors that govern its use as a scaffold for tissue engineering application owing to its influence on the solubility in water and its molecular crystallinity.

4.6.6

Poly-b-hydroxybutyrate

During the past few years, poly(hydroxyl alkanoates) have gained a lot of attention (Valappil et al. 2006). Poly-β-hydroxybutyrate (PHB) is the most studied polymer in the poly (hydroxyl alkanoates) family that has found immense application potential as scaffold for tissue engineering. PHB has been previously used as a matrix that promotes wound healing by enabling cellular growth (Ljungberg et al. 1999). It is a linear homopolymer with head-to-tail arrangement of (R)–hydroxybutyric acid monomer. Poly (3-hydroxybutyrate-co-3-hydroxyvalerate) has found extensive applications as scaffold in bone tissue engineering applications. This polymer is highly biodegradable, biocompatible and has good thermal processing properties. This material after implantation degrades slowly at body temperature with non-toxic degradation product that is excreted out through urine (Mosahebi et al. 2001). They also have good mechanical properties and are highly crystalline. These polymers have also been blended and modified with other synthetic and natural polymers to get improved results. Misra et al. (2010) developed multifunctional scaffolds of poly

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(3hydroxybutyrate) (P(3HB) by incorporating Bioglass BG with different concentrations of Vitamin E along with carbon nanotubes for tissue engineering application. PVA/PHB blend nanofibers showed maximum adhesion and proliferation on pure PHB. Its copolymers with varying ratios of hydroxyvalerate (PHBV) have also been used (Asran et al. 2010). These copolymers of hydroxyl butyrate with hydroxyl valeric acid are less crystalline, more flexible, and more readily processable than PHB itself. Kuppan et al. (2014) developed poly(3-hydroxybutyrate-co-3hydroxyvalerate) (PHBV) fibers via electrospinning and the process parameters were optimized to obtain defect-free fibers. The physico-chemical properties, cell adhesion, proliferation, and gene expression of human skin fibroblast cells were evaluated and compared with 2-D PHBV films. Another group prepared electrospun poly(hydroxybutyrate)/chitosan blend scaffolds for cartilage tissue engineering (Sadeghi et al. 2016). Pramanik et al. (2019) developed poly(hydroxyl butyrateco-hydroxyvalerate) copolymer modified graphite oxide based 3D scaffold for tissue engineering application. Fu et al. worked on tissue engineering the cardiovascular structures where they cultured human pediatric aortic cells on P4HB-coated PGA scaffolds for 7 and 28 days (Fu et al. 2004). A trileaflet heart valve replacement in sheep was done using the composite scaffold of PGA non-woven mesh dip coated in a 1% solution of P4HB.

4.6.7

Polyethylene Glycol-Based Polymers

Synthetic hydrogels based on poly (ethylene glycol) (PEG) or polyethylene oxide (PEO) are gaining popularity as a scaffold material in spite of their inert surface characteristics, mainly due to the flexibility in design and tailorable structure through various conjugation techniques. Poly (ethylene glycol) (PEG) is prepared as an oligomer or polymer of ethylene oxide monomer through anionic or cationic polymerization depending on the acidic or basic catalyst that is used in the process. High molecular weight poly(ethylene glycols) are prepared via suspension polymerization using organometallic compounds as catalysts. PEG and PEO do not present any functional moieties in its structure and hence is bioinert which will help minimize the immune response after implantation. PEG hydrogels can be functionalized and crosslinked using various crosslinking mechanisms and also it is possible to modify the chemistry of the hydrogel to include cell adhesive peptides and other moieties using several bioconjugation protocols (Hamley 2014; Zhang et al. 2014). Bioactive molecules such as cell adhesion ligands and peptides, growth factors, and proteolytic degradation sites can be incorporated into PEG hydrogels which will help enhance cell adhesion, proliferation, migration, and extracellular matrix production. They also provide a bioinert surface with reduced non-specific protein adhesion and the chemistry of the PEG macromers which can be modified to incorporate various cell adhesive ligands and differentiation cues. This modification is possible through many different synthetic strategies. PEG is non-degradable due to its inert nature and hence undergoes limited metabolism in the body and is eliminated as whole polymer chains through the kidneys (30 kDa). Hence high

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molecular weight PEGs are not preferred, and molecular weight less than 50 kDa are typically considered in tissue engineering applications to ensure complete elimination from the body (Tessmar and Göpferich 2007). Additionally, when PEG is modified through synthetic routes care should be taken to synthesize degradable polymer segments that would efficiently release the low molecular weight PEG blocks (Lin and Anseth 2009). The most common approach to make PEG hydrogels is photopolymerization of PEG macromer solutions into solid hydrogels at physiological temperature and pH. Ali et al. (2013) showed that covalent attachment of several cell binding motifs based on laminin like RGDS, YIGSR, etc., on the PEG hydrogels has found to have a profound effect in enhancing the vascularization in the construct. Moreover, to enhance extracellular matrix formation elastin and collagen peptide fragments have also been incorporated into the PEG chains. Poly (ethylene glycol) (PEG) hydrogels were explored for their potential as encapsulation matrices for osteoblasts to assess their applicability in promoting bone tissue engineering (Burdick and Anseth 2002). Dong et al. (2019) developed a strategy for improving this property of porcine small intestinal submucosa (SIS) which was recrosslinked by a four-arm polyethylene glycol (fa-PEG) with succinimidyl glutarate-terminated branches into a three-dimensional (3D) bioactive sponge (SIS-PEG), which possessed porous 3D frameworks to mimic the structure of skin. Asadi et al. (2019) developed a hydrogel based on gelatin/polycaprolactone–polyethylene glycol and loaded with TGFβ1 loaded nanoparticles (Gel/PCEC–TGFβ1) for cartilage tissue engineering and the differentiation of human adipose stem cells to chondrogenic lineage was studied. Kumar et al. (2019) reported on the synthesis of nHA-PEG and nBG-PEG scaffolds using the space-holder method for hard-tissue engineering applications. Another novel photocurable hydrogel made of acrylated poly(ethylene glycol)-co-poly (xylitol sebacate) (PEXS-A) has been synthesized for tissue engineering and used for 3D printing applications (Wang et al. 2019). Burke et al. (2019) examined the stability of a range of poly(ethylene glycol dimethacrylate) (PEGDMA) hydrogels over a 28-day period in simulated physiological solution for tissue engineering application. Bryant and Anseth (2002) developed hydrogels by copolymerizing a degradable macromer, poly (lactic acid)-b-poly (ethylene glycol)-b-poly (lactic acid) end capped with acrylate groups (PEG-LA-DA) with a non-degradable macromer, poly(ethylene glycol) dimethacrylate (PEGDM) for cartilage tissue engineering.

4.7

Polymer Scaffold Fabrication Techniques

One of the challenges in tissue engineering is to find a more suitable method for the fabrication of scaffolds of defined architecture to guide cell growth and development. Ideally, a biomaterial scaffold should have well-controlled micro architectures, including well controlled reproducibility, biocompatibility, pore sizes and porosity, thermal and biochemical stability. These physical factors are also dependant on the efficient nutrient supply and vascularization of the cells in the

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Fig. 4.4 Schematic showing the evolution of different scaffold fabrication technologies which over the years has led to more precision and control over the process of scaffold preparation

implant. Many fabrication techniques are adopted to prepare 3D scaffolds offering cell in-growth and structural support for tissue regeneration (Fig. 4.4).

4.7.1

Conventional (Traditional) Manufacturing Techniques

Conventional scaffold fabrication techniques include phase separation, freeze drying and particulate leaching, fiber meshes and bonding, gas foaming, (Hutmacher 2001; Yang et al. 2001). A major aspect of these constructs will be the replication of in-vivo geometry and dimensional size scale that will aid in the maintenance of an in-vivolike cell phenotype. The fiber bonding technique as the name suggests involves bonding of fiber meshes to form 3D scaffolds and highly relies on the choice of solvent used, polymer immiscibility, and melting temperatures to achieve the desired shape (Mikos et al. 1993). Solvent casting is used in combination with particulate leaching to develop scaffolds with controlled pore sizes. The disadvantage lies in the fact that porosity is achieved by varying the amount and size of salt particles during evaporation and there are chances of residual salt particles within the system. Moreover, in many of these techniques, micro architecture is achieved by altering solute or solvent concentration, thus reproducible features in the micro- and nanometer range may not be obtained. A modification to the above technique involved creating multifunctional 3D scaffolds by combining centrifugation to fiber bonding. These structures display a pore gradient network, which is useful to study the specific affinity of cells to different pore sizes and ultimately allow the creation of hierarchical architectures.

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In another technique of fabrication known as the gas foaming technique, the polymer is saturated with carbon dioxide (CO2) at critical pressures to allow high solubility of the gas in the polymer. When the gas pressure is brought back to the atmosphere pressure, there is a decrease in the solubility of the CO2 in the polymer, resulting in the formation of pores of variable sizes. A similar technique is that of phase separation, where a polymer solution is quickly cooled at low temperature to generate a liquid–liquid phase separation. On quenching this polymer solution, a two-phase solid is formed. The process of sublimation is used to remove the solvent to generate the porous scaffold. Freeze drying is a similar process where the polymer solution is directly frozen at a controlled rate and then dried via the sublimation process to yield porous scaffolds. These fabrication techniques have been used for the preparation of 3D scaffolds for possible tissue-engineering applications. However, there are several drawbacks to these techniques wherein a control on the various parameters affecting tissue formation needs to be considered. Even though a control of pore sizes is possible through these conventional techniques, interconnectivity of pores following a tortuous pathway may not be possible for ensuring effective nutrient supply and release of biological signals for cell signaling. Pore tortuosity, which can be defined as the ratio of the actual length of the arbitrary pathway that a molecule has to cover to pass through a pore to the shortest linear distance, enables the cell to cell contact by maintaining the interconnectivity (Melchels et al. 2010). Even though pore tortuosity is achieved in scaffolds fabricated using conventional techniques, cell viability is seen only up to depths of 0.5–1 mm due to the lack of oxygen from the outside to the center of the scaffold (Liu et al. 2007). Furthermore, growth factor delivery through these scaffolds is also affected by the processing conditions. In most of the techniques that employs a solvent based system, the solvents used can also effect the stability of the growth factor delivery systems due to a change in the pH, which will lead to loss of activity of the factor. In particulate leaching a further problem inducing toxicity is the efficiency of the agent incorporation, as the washing steps may also remove the bioactive agents that have been loaded.

4.7.2

Nano Fabrication-Based Techniques

The cell adhesion, proliferation, morphology, and differentiation are also influenced the topography of the material or the scaffold used. The influence of different roughness factors on the adhesion and growth of various cell types has been studied and these factors have been seen to influence the cell behavior (Desai 2000; Dalby et al. 2003). Rough surfaces with a dimension in the range of 1–50 nm have shown to offer better early adhesion of cells, with reduced fibrous encapsulation, and enhanced integration of implants with host tissue compared to corresponding smooth surfaces. Most of the findings on surface topography for cell attachment involve scaffold materials with a topography having roughness features greater than 0.5 nm. A rough surface with topography well within 10 nm may help to preferentially adsorb small biomolecules and ions. Electrospun nanofibrillar scaffolds with fiber

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diameters ranging from 50-500 nm have been seen to be effective for cell attachment and function. The optimal substrate is one that has similar topographical characteristics like that of the natural cell–substrate interface. The natural substrate provided throughout the vertebrate body for cell attachment is the basement membrane which is a porous, fibrous meshwork of extracellular matrix (ECM) proteins and proteoglycans. The cell–surface interaction influences and modulates cell processes such as adhesion, proliferation, migration, differentiation, and cell morphology which also has an influence on the gene expression and controls cell cycle activity. Hence recreating such a nano featured environment has been considered as one of the conditions for the proper growth of cells and subsequent tissues (Desai 2000). Extracellular matrix is basically a nano featured environment with a complex mixture of pores, ridges, and fibers on which the cells are attached and layered. Engineering nano-fibrous scaffolds with specific fiber orientation is, therefore, a major prerequisite for the success of tissue engineering. Currently, the development of tissue scaffolds using nanotechnology related techniques includes electrospinning, self-assembly, phase separation, of which the process of electrospinning has been the most widely used technique used by research groups in the fabrication of nanostructured scaffolds (Engel et al. 2008; Ma 2008). The advantages offered by the use of such nanotechnologies to create scaffolds in comparison with the conventional fabrication techniques is that these techniques can provide more uniform fabrication of ultra-fine fibers having controlled orientation in arrangement and pore geometries, high surface area, and high aspect ratio which are necessary for better cellular growth functions in-vitro and in-vivo, because they directly influence the cell adhesion, cell expression, and transportation of oxygen and nutrients to the cells.

4.7.3

Additive Manufacturing-Based Techniques

In the conventional fabrication techniques as well as the nano-fabrication based technique, the control over the regulation and positioning of pores and their relative density and size cannot be regulated. Moreover, the seeding of cells manually is not efficient enough to control the distribution of cells within the scaffold, is user dependant, and is not economically viable when considering the commercialization aspects. This is where the use of additive manufacturing helps to address these issues (Ngo et al. 2018). Furthermore, the use of this automated technology can also help to facilitate production of tissue engineered constructs using good manufacturing practices (GMP) with the appropriate quality control. The work flow of fabricating a scaffold via additive manufacturing techniques includes generation of the 3D model of the tissue using 3D computer software (Fig. 4.5). The files are then converted to the corresponding .stl file and these models are then sliced into thin layers which are then reproduced into the 3D scaffold using the additive manufacturing unit. There are different types of additive manufacturing techniques

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STEP 1:

STEP 2:

Generaon of the 3D model of the ssue using a 3D computer soware

The files are converted to the corresponding .stl file

STEP 3:

STEP 4:

Models are then sliced into thin layers

Reproduced using addive manufacturing techniques

Fig. 4.5 Work flow of the steps involved in additive manufacturing

employed in the area of tissue engineering. These include fused deposition modeling, selective laser sintering, stereo lithography, and 3D printing. Stereolithography is one of the oldest techniques that dates back to 1986 where Charles W. Hull used to develop scaffolds where liquid polymer material was cured with ultra-violet light layer by layer in three axes to form a solid 3D structure. The process is limited to the use of photopolymerizable materials. Selective laser sintering technique uses infrared laser beam to sinter powder on a powder bed which may be a polymer, wax or even metal. In polymers, the temperature of the laser is adjusted to meet the glass transition temperature of the polymers used to initiate effective bonding and fusing of the polymer powder particles in a layer by layer manner to obtain the 3 D solid. This method is limited by the power of the laser and the thermal efficiency as well as the glass transition and melting temperatures of the polymer to be used. Since polymer powder is used in the process, the pore size will also be very small and hence larger pores are difficult to create. In fused deposition modeling, a fiber of polymer material is extruded and builds up in layers to form the scaffold. Zein et al. (2002) studied the fabrication of PCL scaffolds via FDM technique and their physical properties were evaluated. Some disadvantages of this process include the high temperature required for fusing or melting the polymers and the need for support structures to create irregular shaped scaffolds. 3D printing technology for tissue engineering applications dates back to 1990, where the developed scaffolds used were mostly of synthetic origin. It was not until the last decade that this technique saw the advancements in the use of both natural and synthetic polymers with or without cells which came to be known as 3D bioprinting. The print solution used for such printing of tissues is known as bioink (Zhang et al. 2015). Bioinks can be cells and scaffold based or scaffold free types. 3D bioprinting helps to develop structures that are similar to the native ECM which can be used to study cell–cell and cell–matrix interactions (Gu et al. 2016). The advantages of this technique include precise control of the cell density and its

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distribution, high-resolution cell deposition that is viable, controllable to scale, and cost-efficient. The basic concept of this technology is the building up of layered structures of cell-laden building blocks or cell aggregates using several types of 3D printing techniques. There are three primary technologies, namely, laser-assisted bioprinting, inkjet bioprinting, and extrusion bioprinting each having their own requirements and benefits (Kačarević et al. 2018). Polymeric materials encompass a wide variety of materials that can range from being soft to hard, or synthetic to natural, however, the most commonly used polymers for 3D printing tissue engineering include PCL, PEEK, PLA, poly(lactic-co-glycolic acid) (PLGA), and chitosan. In the last 10 years hydrogels have been commonly used as bioinks that are biocompatible and have high liquid holding capacity and can also be loaded with cells (Bishop et al. 2017). Natural hydrogels are most commonly used both in their native or modified forms. Examples include chitosan, alginate, gelatin, agarose, hyaluronic acid, gelatin methacrylate, hyaluronic acid methacrylate, etc. Synthetic hydrogels include the acrylates of polyethylene glycol and which are photocrosslinkable and can be further modified to conjugate peptides, Pluronic (PF127) or poloxamer (block copolymer consisting of a central poly (propyleneoxide) (PPO) block flanked by poly(ethylene oxide)(PEO) blocks), PEG-Fibrinogen with unmodified PF127, diacrylated Pluronic F127(PF127-DA) with PF127, etc. The major challenges in the development of hydrogels as bioinks encapsulating cells are the print temperature, the extrusion pressure applied, the solvent used, and the pH of the developed gel which may affect the cell growth after printing.

4.8

Conclusion and Perspectives

The use of appropriate 3-dimensional scaffold structures for cell and tissue growth is one of the major requirements and has been a major challenge in the tissue engineering paradigm. In recent years, both synthetic and natural polymers have found application as scaffold structures. Modifications to introduce cell responsiveness and improvement in mechanical strength have led to several modification and synthesis procedures. Furthermore, the fabrication of three-dimensional framework structures with the appropriate pore structures started with the conventional techniques like fiber bonding, solvent casting, freeze drying to more complex nano-techniques like phase separation and electrospinning and then moving to more advanced technologies of freeform fabrication techniques like 3D printing and stereolithography techniques. However, there are several advantages and disadvantages in the use of each technique depending on the polymer that has been selected as scaffold. This is where a proper understanding of the tissue to be regenerated, composition of the ECM of the intended tissue, the rate of remodeling need to be considered while selecting the right polymer, be it synthetic or natural in origin. With the advent of new technologies, tissue engineering strategies have diversified from just seeding of cells onto scaffolds to generating 4D and 5D printed tissue constructs, which can respond to stimuli, controlled biomimeticity, and use of

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sophisticated bioreactors. The emergence of new polymers and the advances in the modifications to the natural and synthetic polymers have opened up more avenues for developing smart biomaterials for tissue regeneration. In fact, a number of tissue engineering technologies have advanced to human clinical trials and commercialized. The rapid growth in research in organ-on-chip field has also shown optimum characteristics for generating engineered organs to study pathophysiology or even drug testing. However, a lot of research is still warranted to study the functional characteristics of the regenerated tissue and the integration and new tissue formation within the defect site. Hence there is a long way to go before these varied strategies can follow common functional guidelines for specific tissues especially in scaffold selection and use.

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Composite Biomaterials in Tissue Engineering: Retrospective and Prospects Charu Khanna, Mahesh Kumar Sah, and Bableen Flora

Abstract

The last two decades have seen the advancement in tissue engineering strategies to provide viable alternatives to native tissues, implants, and prostheses. Progress made in biomaterials development assisted with macro-, micro-, and nanotechnologies contributed to mimic the native tissue and its microenvironment for in situ regeneration and further its complete replacement with functional tissues. A key component of this strategy i.e., biomaterials development, requires a range of properties to support regeneration of specific tissues which is mostly unachievable with a sole component and thus needs to be bio-composite of two or more components in specific ratio, form, and distribution. This chapter is dedicated to bio-composites development for tissue engineering applications and will be focusing on the properties required and strategies employed for the development of different types of bio-composites for hard and soft tissue regeneration. The modulation of material properties by compositing with different biomaterials and approaches affecting its functionality and efficiency for tissue regeneration is discussed. This chapter is also reporting the recent advancements as has happened in the terrain of bio-composites for tissue regeneration and the challenges encountered to achieve the benchmark success.

C. Khanna School of Pharmaceutical Sciences, Lovely Professional University, Phagwara, Punjab, India M. K. Sah (*) Department of Biotechnology, Dr. B. R. Ambedkar National Institute of Technology, Jalandhar, Punjab, India e-mail: [email protected] B. Flora Department of Biotechnology, Lovely Professional University, Phagwara, Punjab, India # The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 B. Bhaskar et al. (eds.), Biomaterials in Tissue Engineering and Regenerative Medicine, https://doi.org/10.1007/978-981-16-0002-9_5

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Keywords

Bio-composites · Tissue engineering · Surface modification · Nanofabrication · Reinforcement

5.1

Introduction

Tissue engineering, a multi-disciplinary field, involves the use of composite biomaterials which are referred to as the materials being used in medical devices (Gurman et al. 2012). A great interest has been developed globally in this field owing to the scientific, industrial, and medical perspective of such materials. Immense success of such biomaterials can be witnessed for hard tissues such as in treating fractures and in medical issues involving dentistry. Still comprehensive researches in this sphere are targeting the advancements of techniques, involving such bio-based materials that can improve the tissue conditions such as those assisting in the regeneration of damaged neural tissue. Thereby, they have potential to improve the life of the concerned patients (Jammalamadaka and Tappa 2018). Structurally, the chemically different constituents of composite materials possess discontinuous phases which get embedded in a continuous phase, where discontinuous phase being harder and stronger in nature (Iftekhar 2004). For instance, medically, the musculoskeletal framework involves the massive integration of structural tissues such as bones, ligaments, tendons, cartilages, skeletal muscles, and peripheral and spinal nerves, the bones and tendons illustrate the example of biological composite. A large number of bio-composites, in spite of being distinct chemically and morphologically, are being studied for their potential in tissue engineering, and their use significantly depends upon the nature of biomaterial and the healing tissue involved (Fig. 5.1) (Iftekhar 2004). Hence, these composites have been classified in multiple categories realizing their specific properties which ensures their applications. The scaffolds, constituted by such biomaterials, have provided a platform where they may mimic a healthy living tissue enabling the proliferation of the healthy cells

Fig. 5.1 Scanning electron micrograph of (a) eggshell membrane, a natural polymer composite with globular proteins tightly embedded in collagen matrix, and (b) silk fibroin/synthesized nanohydroxyapatite (nHAP) developed composite for bone tissue engineering (unpublished data)

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leading to repair of impaired body component and enhancing the growth of healthy tissues. Still, the biodegradability along with bio-compatible make up of such functional scaffolds needs to be ensured (Vats et al. 2003). Further, the proteinaceous biomaterials are of another interest for their application in biomedical devices (Crnković et al. 2018). The advancements related to biosynthesis of non-canonical amino acids containing proteins clearly depict the processes where the chemical functionaries of such proteins be customized for enhancing their characters and subsequently, applications in the field of biomaterials (Crnković et al. 2018). Environmental concerns towards using of synthetic plastic-based polymers compelled the researchers to analyse the potentials of natural polymers such as bacterial cellulose and cotton fibres. The modifications of such natural materials with other additional materials are being studied to enable them to behave as bio-composites for tissue engineering. To exemplify, the functionalities of bacterial cellulose when modified with xylans has been studied for its potential as bio-based material (Santos et al. 2017). Similarly, cotton fibres along with metal-organic framework presents its ability as substrate for waste water filtration, photo catalytic property, as a decontaminant and as a source of degradation of micro pollutant (Schelling et al. 2018). Such studies illustrate the extensive progress that has been made towards the developments of bio-based composites as intended towards strengthening the applications of tissue engineering. More recently, the nano-based biomaterials, such as carbon nanotubes, have also shown some promising platform for successful tissue repairing. The nano dimensions and surface chemistry of such particles differ from their actual natural state but still the bio-compatibility is the main issue of concern with such materials (Catalán and Norppa 2017). Although, much work is being carried out in developing newer bio-based materials, still in depth understanding of mechanical attributes, along with their effects towards cellular activities such as proliferation, growth, and differentiation is need of the time. Hence, a remarkable success may happen for biomaterial perspectives, if the concerned issues are addressed for composites involved, so that they can provide ample safer solutions for recovery of injured tissues and uplift the physical miseries in the life of a patient.

5.2

Bio-Composite Components: Classes and Desirable Properties

Composite materials are recognized as those which are prepared by amalgamating two or more constituents possessing discrete morphology, composition, and physical properties to fabricate a material with specified physical, mechanical, and chemical traits. The primacy of such materials has been realized as they present better properties when compared to the individual component along with flexible design (Salernitano and Migliaresi 2003). Such composite materials have significant applications as these curative strategies may imitate the natural processes enabling the tissue repair along with regeneration leading to success or failure of the involved techniques (Gurman et al. 2012; Correia et al. 2014).

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The development of neo-tissues, by the involved cells, are not only influenced by the biological environment, but also by the characteristics of the biomaterial such as its bio-compatibility, porosity (micro/macro), biodegradability, and interconnectivity along with the design and build of scaffolds intended to fit in the injured body space. Hence, the researchers working with such materials must understand the basis of choice of composites under study, along with their advantages and limitations (Cui et al. 2016). This section provides a review to categorize different bio-based materials along with scrutinizing certain key factors that may provide fundamentals of selection criteria of these materials. On the basis of tissue response, the biomaterial may be classified to be bio inert (possess minimum host tissue interaction), bioactive (host tissue interaction), and biosorbable (provide framework for some new tissue to grow while itself gets absorbed). Realizing the structural bonding or chemical makeup and properties, the biomaterials may be classified to be metals, ceramics, polymers, and composites (the combination of first three) (Fig. 5.2) (Ramakrishna et al. 2016). Based upon the source, composites may be classified to be natural and synthetic. A substantial number of natural materials employed in formulating tissue engineered products includes alginate, collagen, porcine, bovine, stem cells, and metals as obtained from algae, animal tissues, and nature. Currently, several natural products are still under study such as fibres from different parts of various plants (Cocos nucifera, Luffa cylindrical, musa indica, Phoenix dactylifera, and Ceiba pentranda) and animals such as chicken feathers (Pickering et al. 2016). Synthetic entities including glass ceramics, silicones, and polyesters contribute significantly to application in the

Synthetic

Natural

Source

Biomaterials Composites Metals

Inert

Structure

Tissue response

Bioactive

Polymers Biosorbable Ceramics Fig. 5.2 Different classes of biomaterials that could be considered for the development of bio-composites based on the tissue of interest and associated cellular microenvironment

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different fields such as dentistry, orthopaedics, and cardiac so as to repair or to regenerate the injured tissues and hence acknowledged as healers. The engineering performance of different composite materials are subjected to different factors such as the discrete components; form, quantity, and arrangement of the structural components; and the behavioural interactions of the constituents. The structure of composite biomaterials possesses a bulk phase which constitutes matrix, a continuous phase; a discontinuous and dispersed phase, known as reinforcement; and a third phase present between the two, renowned as interphase. As per the components used in the matrix phase, the composites are classified to be ceramicmatrix, polymer matrix, and metal matrix composites. With respect to reinforcement, the composites may be categorized as particle reinforced composite and fibre reinforced composites (Iftekhar 2004; Chandramohan and Marimuthu 2011). The larger surface area of the matrix accepts the load, which is then transferred to the reinforcement phase. This phase being different ensures mechanical changes in the formed composite material suitably for its stiffness, strength, toughness along with fatigue resistance. Hence, such substances may be produced with the desired properties for their application in the body (Iftekhar 2004). The different constituents, as used in the formation of composite materials, have been enlisted in Table 5.1. There are certain key factors, playing an integral role in appropriate growth and regeneration of impaired biological tissues. Composite materials may behave distinctly with different biological system such as tissue processes in healing wounds, stem cells in bioreactors, and target cells in gene therapy. Thus, bio-compatibility may be better recognized as the characteristic of material-biological host system (Cabral et al. 2014). The composites must promote the cellular adhesion; provide pathways for vascularization, exhibit non-allergic, non-toxic, non-pyrogenic, and non-carcinogenic response along with ability to promote the biological tissues for Table 5.1 Examples of components used in formulating composite biomaterials (Iftekhar 2004) Bio-composite phase Matrix

Category Thermosets Thermo-plastics Inorganic Resorbable polymers

Particles Fibres

Inorganic Organic Polymers Resorbable polymers Inorganic

Examples Epoxy; polyesters; polyacrylates; polymetacrylates; silicones Polycarbonates; polyesters; polysulfones; polyolefins Calcium carbonate ceramics; glass ceramics; carbon; titanium; steel; hydroxyapatite Polylactide; polydioxaone; polyglycolide; chitosan; alginate; collagen Alumnina; glass Poly-methacrylate Polyolefins; aromatic polyamides; polyesters; PTFE; UHMWPE Polylactide; silk; collagen Glass; carbon; hydroxyapatite; tricalcium phosphate

PTFE polytetrafluoroethylene, UHMWPE ultrahigh molecular weight polyethylene

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the implant integration (Velu et al. 2020). For an instance, PTFE polymers are generally considered to be bio-compatible but in situations where cell adhesion to the polymer is expected, it cannot be used due to hydrophobic nature of the material. Similarly, the host response with other materials such as titanium and silicon materials must be ensured before use (Cabral et al. 2014). Biodegradable and bioresorbable properties have attracted the significant interest of researchers to provide essential mechanical function during tissue engineering processes. Various composites in the form of implants, prosthesis, scaffolds or as drug delivery agents are currently being used successfully in pharmaceutical and biomedical industry. Polymers, such as PLA and PGA, are utilized to manufacture sutures, fixation plates, and interference screws and for craniomaxillofacial fixations. The biodegradable materials must degrade effectively along with assisting the healing and regeneration of the concerned tissue as in case of scaffolds. The primary parameters while selecting a biodegradable material to be ensured are biosafety, age of the patient, risk of infection, fracture types, and physical condition, etc. (Prakasam et al. 2017). Along with, there are certain other factors which are influential towards the biodegradation as highlighted in Fig. 5.3. Bioactive ceramics including calcium phosphatebased hydroxyapatite, brushite, and tricalcium phosphate are extensively being studied as replacement materials in the bone tissue owing to their osteoconductive and resorption attributes. In case of metals, various alloys such as Mg-Ca and Mg-

Pore size, porosity

Surface roughness Surface area to vol ratio

Size/shape

Implantation site

Enzyme concentrations

Additives/ Impurities

Scaffold

Fabrication method

Biodegradation parameters

pH/ ionic strength

Cell type/ density

In vivo factors

In vitro factors Mechanical loads

Mechanical loads Tissue modellind and remodelling

Biological medium composition

Fig. 5.3 Factors affecting biodegradation of composites

Incubation temperature

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Mn-Zn have been studied for their lower in vivo corrosion rate. Titanium alloys exhibit easier bonding with the bone along with corrosion resistance and lower modulus, leading to better integration and hence being used as ventricular devices, staples, and screws in spinal surgery, implantable drug pumps, as pacemakers and as dental implants (Prakasam et al. 2017). Ideally, the mechanical integrity of the developed implantable construct must match with that of the patient’s body tissue under treatment which may be hard or soft in nature. Varied factors that illustrate the mechanical attributes of the composites include compressive modulus, compressive strength, fracture toughness, bending strength, and Vickers hardness. Bio inert ceramics including ZrO2 along with Al2O3 possess high mechanical strength and durability and are used in forming artificial femoral head and acetabular cup for hip prosthesis. The mechanical integrity exhibited by hydroxyapatite is also a promising substituent for bones along with non-resorbable in nature. Studies have also revealed the enhancement of mechanical properties when the organic-inorganic hybrid composites are conjugated at molecular levels such as in case of Poly dimethyl siloxane and PCL/silica hybrids (Kaur et al. 2019). The porosity of the composite materials especially scaffolds play an integral role for enhanced cellular differentiation, their proliferation, and the migration (Bakhshandeh et al. 2011). Ideal porous scaffold must possess specified pore size, higher porosity along with significant surface to volume ratio which ensures the proper diffusion of substances such as nutrients and drugs. Various polymers such as polyglycolide (PGA), polylactic acid (PLA), polycaprolactone (PCL), and poly ethylene glycol (PEG), are widely used clinically such as for skin tissue engineering (Chaudhari et al. 2016). The surface morphology plays an integral role, as it is the platform of cellular interactions which are controlled by surface topography, surface energy, surface chemistry, and surface functionality (Cabral et al. 2014). For example, the wet and rough surface of scaffolds with bioactive glass nanoparticles accelerates blood clotting time and thus exhibit better suitability for tissue engineering (Rai et al. 2010). Several factors are thus responsible for the ideal functioning of the bio-composites in tissue engineering. Focussed to cater the need for inducing desired properties in the available composites till date, the strategies which enable these materials to be completely functional must be studied for effective repair and regeneration of the impaired tissues.

5.3

Strategies of Bio-Composite Development

Currently, the TE approach using composite biomaterials includes the strategies involving (a) Scaffolds: the three-dimensional porous supports; (b) Signalling molecules such as growth factors/bioactive agents; (c) Cells with undifferentiated or differentiated characters (Vats et al. 2003; Correia et al. 2014). Hence, scaffolds synthesized from composites are not only focussed for their support property but also for delivering required therapeutic factors including proteins, drugs, and growth factors. To qualify the composites for ensuring cell viability along with cell functionality, as in scaffolds, different characters are usually considered which includes

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Fig. 5.4 Strategies for the development of bio-composite

surface characters such as reactivity, chemistry, charge, and roughness; rigidity and contact angle. Such properties are responsible for evaluating the cell-cell-biomaterial composite and subsequently cell–cell interaction, thereby supporting cells to survive, differentiate, enhancing cellular adhesions, fastening cellular responses, deliver therapeutics along with ensuring biodegradability, bio-compatibility, adaptability, directional stability, mechanical strength, and serializability. Various pitfalls should be eliminated to synthesize novel bio-composite. Numerous challenges inclusive of bio-compatibility, biodegradation, mechanical strength, and topographic attributes of the material should be considered. However, to achieve a functional novel bio-composite, these interim barriers need to be strengthened. Different methods have been implicated to improvise the material compatibility, degradability, and mechanical properties of the chosen material. Different strategies are employed for the development of bio-composite from different classes of biomaterials selected as per the desirable properties to be repaired or regenerated. Among these, the present scenario employs conventional blending, advanced fabrication methods such as co-electrospinning and 3-D printing, reinforcement methods, and surface modifications (Fig. 5.4).

5.3.1

Conventional Blending and Mixing Technique

Biomaterials of different classes as per requirement are mixed and blended with homogenizer or even simple magnetic stirring in the form of liquid/liquid, solid/ liquid, and solid/solid form. The selection of form and parameters of homogenization is decided to achieve homogenous phase mixture and stability of components. The bio-composite of silk fibroin blended with PVA has been reported to develop

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three-dimensional porous scaffolds by salt leaching method by our group (Sah et al. 2017).

5.3.2

Advanced Bio-Fabrication Methods

Several fabrication techniques involving use of bio-composites to serve tissue repair can be witnessed now-a-days. The conventional techniques are generally inclusive of solvent casting, freeze-drying, melt moulding, particle leaching, and gas foaming. However, these methodologies exhibit certain vulnerabilities such as uncontrolled porosity, interconnectivity, and improper spatial arrangement that inhibit the adequate distribution of cells, and hence, limit the success of bio-composite. However, the rapidly growing technologies have revolutionized the field of biomaterials where the advanced methods have replaced the older methods exhibiting better results. Techniques such as additive manufacturing, also known as 3D bio printing, use computerized designing through different software to maximize the specific requirement of bio-composite and thereby, lowering the cons of older methodologies. Although 3D bio printing has eliminated the major concerns of conventional methodology but designing a medical device or biological tissue or organ is still accepted more challenging (Hoch et al. 2014; Moroni et al. 2018; Liu et al. 2020). To overcome this, the engineered structural designing with a bigger approach using computational tools is being highly looked upon. Additionally, there are heap of other additive methodologies which are used as advanced fabrication methods. Selective Laser Sintering (SLS), Stereolithography, Vat-photopolymerization, Fused Deposition or 3D fibre deposition, Powder based fusion process, and Spheroid based method have been reported as leading approaches (Li 2019; Moroni 2018). Although this progressive breakthrough of different methods became a boon, but still it has some of the technological gaps and challenges. Database update, Expensive biomaterials, cellular and functional moieties, material development, and standardization have been appraised as big pitfalls in this field (Starly 2015). In addition, a more improved technology recently developed that can transform shape either before cell deposition or after cell deposition leads to formation of highresolution dynamic shape of desired material known as 4D biofabrication (Ionov 2018).

5.3.2.1 Co-electrospinning Co-electrospinning, also known as blend electrospinning, has been extensively used for fabrication of composites that can be broadly practiced for tissue engineering or in drug delivery system. The matrix is normally provided with drug releases through the process of diffusion. Usually biodegradable polymeric matrix has been reported in case of co-electrospinning (Boroojen et al. 2019; Ivanoska-Dacikj et al. 2019; Yahia et al. 2019; Yongcong et al. 2019). Coaxial electrospinning, the advanced version of electrospinning, is of remarkable interest to researchers nowadays. Implicating this technique, drug releasing nano fibres were synthesized using two polymeric liquids, which were simultaneously suspended through two needles,

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external and internal, which led to formation of a core structure (Vysloužilová et al. 2017). It has also been scrutinized that core or hollow shell fibres could be fabricated from non-electrospinnable materials such as polymers employing this technique (Moghe and Gupta 2008; Agarwal et al. 2009). Fibres encapsulated with nanoparticles and ceramics can be prepared using coaxial electrospinning technique (Moghe and Gupta 2008; Agarwal et al. 2009). In a research, PCL reinforced with starch has been fabricated through co-electrospinning technique and the fibres obtained can be used in wound healing applications (Komur et al. 2017). Successful synthesis of nanofibers using co-electrospinning has been studied where core and shell were composed of PLA and PVA, respectively. This fabrication exhibited enhanced adhesion and proliferation on human embryonic kidney cells, clearly illustrating the suitable properties such as surface wetting, mechanical, and cytocompatible features (Alharbi et al. 2018). Non-toxic nanofiber scaffold as a source of drug carrier was prepared by coaxial electrospinning using sodium carboxymethyl cellulose, methyl acrylate and poly (ethylene oxide). Tetracycline hydrochloride drug was integrated in the core and subsequently, the performance of this model was analysed. Sustained release of the drug along with potential antimicrobial activity demonstrated the scaffold for its utility in health industry (Esmaeili and Haseli 2017). In addition, PLA as core has been incorporated with polyacrylonitrile/cellulose nanocrystals and polyacrylonitrile/chitin nanocrystals as shells, respectively, and the enhanced tensile strength, water permeability, and antimicrobial activity have been reported (Jalvo et al. 2017).

5.3.2.2 Bioprinting Three-dimensional printing is another exemplary technique where bio-composite entities have been composed to form such functional structures which could address the suitability of tissues for transplantation. The bio printers, as used in this strategy, are categorized based on the method of printing and instrumentation. Using bio printing, the composite mixture/or blend of materials can be printed in three dimensions as required or different biomaterials can be localized in a threedimensional printed scaffold based on the requirement of tissues to be regenerated. Currently, the main techniques of bio printing include Inkjet 3D bioprinting, Extrusion 3D bioprinters, and Laser-assisted 3D bioprinting. Materials as PVA, poly-DLLactide (PDLLA), citric acid, and water are generally used in Inkjet 3D bio printing which is well known for its speed and accuracy. However, the challenge remains with the higher temperature near the printing head, which results in failure the process (Bishop et al. 2017; Tappa and Jammalamadaka 2018). Polymers including both naturally and synthetically derived can be easily printed using Extrusion 3D bio printers, still the temperature, nozzle diameter, and speed of printing are some vital modalities of such bio printers. Laser-assisted 3D bio printing has been reported to be more successful than Inkjet and Extrusion bio printing due to high cell viability, but it is an expensive process. It is nozzle-free and there is a ribbon which is supported by titanium or gold layer and is suspended which contains bio ink (Murphy and Atala 2014; Seol et al. 2014; Starly 2015). Figure 5.5 illustrates

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Fig. 5.5 Different methods of three-dimensional composite scaffolds printings (a) inkjet, (b) extrusion, and (c) laser-assisted Bioprinter for tissue engineering applications

these 3D printing methodologies for the development of composite scaffolds for tissue engineering applications. In addition to above bioprinters, numerous other technological methods have been reported for bio printing including selective laser sintering (SLS), fuse deposition modelling, two photon polymerization, and electrospinning (Moroni et al. 2018). The wood fibre reinforced bio-composites have been fabricated by fused deposition modelling showed the relation between mechanical properties and the orientation at which printing is done. However, the porosity reported to be improved but the cohesion of material got reduced. Moreover, the tensile strength also reported to be lowered but the hygroscopic moieties have been improved which can be a blessing to manufacture programmable moisture actuated functionality (Le Duigou et al. 2016). Bioinks used in bioprinters comprised of hydrogels with are naturally derived or synthetically made and another multimaterial bioink have also been reported for selective bioprinters (Adepu et al. 2017). Materials used for bone implant should be osteoconductive as well as osteointegrative in addition to bio-compatibility and biodegradability. Different methods have been implemented to fabricate bone scaffolds including in situ forming implants or hydrogels. An outbreak of new technology called “additive manufacturing” found to be promising when targeting patient specific implants. There are more diverse methods to produce scaffolds inclusive of stereolithography, selective laser sintering, and fused deposition modeling. Chen and co-workers successfully prepared two types of scaffold inclusive of Tricalcium phosphates (TCP) and Titanium, respectively, through computer aided designing and

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manufactured through lithographic based ceramic manufacturing. The scaffolds have been tested in rabbit for calvarium defect and reported both the scaffold showed similar microstructure and bone regeneration behaviour (Chen et al. 2013).

5.3.2.3 Reinforcement Methods To achieve bio mimic structural and functional properties of the bio-composites, several fillers have been incorporated in specific matrix forming binary or ternary composites that can be witnessed in diverse applications. The main objective of reinforcement of specific material used in fabrication of composites aims to improve physiochemical and mechanical traits of the blend. To achieve such targets, numerous types of material are available (Fig. 5.6) for synthesizing reinforced bio-composites that can demonstrate successful applications in tissue repair. Materials have been classified based on their source either obtained naturally or synthesized. Polysaccharides, proteins have been obtained from plant or animal sources. On the other hand, metals, ceramics, polymers, and composite material slips into man-made category. Composite material comprises of blending of two or more material with improved properties and functionality focussing the target application. Several blends have been made through mixing of two or more materials that have resulted in enhanced bio-compatibility or physiochemical traits and showed successful applications in tissue engineering. Further, bio-composite material can be classified on two criteria. Firstly, the classification is done based on matrix material. These include Metal matrix, Ceramics matrix, and organic matrix which is further sub-divided into polymeric matrix and carbon matrix composites. Whereas, the reinforcement material has been used for synthesis of composite, correspondingly, it is also used to classify composites. Fibre reinforcement, Nanoparticles reinforcement are some of the filler-based material. Biodegradable metal-based composites have been synthesized through different methods including powder metallurgy, casting, pressing, fibre enhancement, and many others (Yang et al. 2018). Table 5.2 summarizes common examples of reported bio-composites by reinforcement in different matrices through various fabrication methods for specific

Materials

Natural

Polysaccharride based

Protein based

Synthec

Metals

Ceramics

Polymers

Composites

Fig. 5.6 Different biomaterials classes utilized for the development of reinforcement-based bio-composite materials for tissue engineering application

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Table 5.2 List of bio-composites, reinforced in different matrix through various fabrication methods and its applications Fabrication method Electrospinning

S. No. 1

Bio-composite PCL/gelatin/ chitosan

Matrix PCL

2

MgO/silk/ Polycaprolactone PHB/Kenaf fibres

Silk/ PCL PHB

Electrospinning

PVA/collagen/ HA Collagen/PVA

PVA

Electrospinning

Collagen

6

PVA/nHA/iron oxide NPs

PVA

7

PLLA/nHA

PLLA

Solvent casting and freezedrying Ultrasonic dispersion and freeze-thaw Microwave

3

4 5

Extrusion

PLLA poly (L-lactide), nHA/HA hydroxyapatite, polyhydroxybutyrate, PCL polycaprolactone

Applications Skin and other tissue engineering Bone regeneration Packaging

Bone regeneration Osteochondral regeneration Cartilage tissue engineering Bone tissue engineering

PVA

poly

vinyl

References Gautam et al. (2014) Xing et al. (2020) Kuciel and Liber-Knec (2011) Song et al. (2012) Iqbal et al. (2019) Huang et al. (2018) Singh et al. (2018) alcohol,

PHB

tissues. Different material blended using various fabrications techniques in order to improve bio-compatibility, degradability or enhancement of physical or mechanical properties that can be further used as replacement for damaged cell or tissues. Material can be incorporated in different forms as nanoparticles, hydrogels, and moreover, it has been testified that multi walled carbon nanotubes can be used as filler to regenerate the lost part. In a study, hydroxyapatite has been reinforced with carbon nanotube (CNT), improving their mechanical, physical, and biological properties of hydroxyapatite (Mukherjee et al. 2014).

5.3.3

Nano-Particle Reinforced Bio-Composites

Nano materials, the fourth-generation composite materials, have emerged as a significant tissue engineering strategy being comprehensively studied for their success in treatment of different biological conditions. Nanocomposites are defined as, such matrix possessing distinct configurationally properties along with fillers sizing less than 100 nm. The higher surface area, robustness and reactivity of nano materials help enhancing the different attributes (physical, mechanical, optical, and chemical) of the resulting bio-composites and are subsequently being investigated for the biomedical implementations (Table 5.3) (Velu et al. 2020). Studies to improve the thermal and mechanical attributes of biodegradable polymers reinforced with nano-particles such as nano-clay, nano-apatite along and

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Table 5.3 Applications of different biomaterials with nano-size approach Category Metals

Ceramic

Nanofillers Gold nanoparticles Silver nanoparticles

Magnetic nanoparticles (Fe2O3) Nanohydroxyapatite powder

Titanium oxide nanoparticles ZnO nanoparticles

Carbon

Carbon nanotubes

Matrix Extracellular

Application Musculoskeletal tissue engineering Antimicrobials in dress wounds Enhanced modulus, electrical conduction, sustained ag ions release, non-toxic with human cells Magnetic and antibacterial

Reference Smith et al. (2017) Bhowmick and Koul (2016) Kumar et al. (2016)

Chitosan

Controlled pore sized scaffolds with higher strength for load bearing joints

Cellulose

Scaffolds for artificial bone tissue

PCL

Implants enhancing cell attachment Devices such as surgeon’s gloves and water bags with antibacterial property

Chen et al. (2013), Tomoaia et al. (2013), Kim et al. (2015) Gouma et al. (2012); Lee et al. (2015) Kiran et al. (2018) Li et al. (2019)

PVA based hydrogel PCL reinforced with grapheme oxide Polyurethane backbones

Rubber matrix reinforced with cellulose fibres Polyurethanes PCL Collagen

Graphene

Cellulose

Nanocrystals

Chitosan and hydroxyapatite Polypyrrole

Cement based materials

Nanofibrils Bacterial cellulose

Reinforced fibre-cement composites

Enhanced osteoblast adhesion to the scaffold Coating for implants along with improved bio-compatibility – Bone and cartilage regeneration Vascular implants

Das et al. (2013)

Jell et al. (2008) Pan et al. (2012) Tan et al. (2010) Im et al. (2012) Kumar et al. (2015) Kurtis (2015) Sarker et al. (2018) Schumann et al. (2009), Picheth et al. (2017)

carbon nanotubes have been conducted where enhancement of Young modulus along with glass transition temperature of the polymers such as epoxy (Šupová et al. 2011) has been reported. As compared to traditional materials, the carbon

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nanotubes have been known to be the stronger material. These tubes have been categorized into three types, viz. zigzag, armchair, and chiral, each possessing its individual fracture, electrical and mechanical properties. Further, the geometry of this nanotube permits the cell delivery of the hydrogen along with drug for fuel cells, as biomedical materials for microenvironmental sensing, formation of scaffolds, for improved cellular tackling and for delivering transfect ion agents (Harrison and Atala 2007). Nanomaterials, although being novel materials and possessing some significant properties, might pose some risk to the human body. Nanotubes have been reported to be cytotoxic in some studies along with poor biodegradability. Incubation of nanotubes with alveolar macrophages exhibited cytotoxicity post 6 h of exposure (Li et al. 2012). Such issues hinder the applications of such materials as implants in the body and must be studied exhaustively to ensure their use for tissue engineering. Different conventional strategies undertaken to prepare the nano-composites includes sol-gel, hot pressing, freeze-drying, casting and electro spinning techniques which enhances the functionality of the formed product by incorporating the required moieties such as porosity and strength (Ravichandran et al. 2015; Rashti et al. 2016). Sol-gel technique facilitates the homogenous distribution of nanoparticles in the polymer followed by its jellification to the solid form. The nanoparticles of Fe3O4 were dispersed in chloroform, thereafter, the colloidal suspension of hyperbranched polyurethane formed and subsequently the polymerisation was done. The structure thus formed exhibited promising material for fabrication of biomedical devices as it demonstrated improved shape memory behaviour, thermomechanical properties, antibacterial potential, cytocompatibility, biodegradability, and non-immunological attribute (Das et al. 2013). Another study reported the enhanced mechanical and bio-compatible characters of polyurethane with doping of silica nanoparticles by sol-gel technique (Rashti et al. 2016). The potential of cardiac patch formulated with nanofiber using electrospinning has been analysed for its cardiac tissue regenerative behaviour in myocardial infarction (Ravichandran et al. 2015). More recently, the additive manufacturing with a unique ability to fabricate the complex three-dimensional structures along with constrained geometries are being widely applied for the rapid prototype modelling (Velu et al. 2020). The fabrications of newer nano-composite materials using additive manufacturing techniques are being looked to open up the opportunities where the possibilities of formulating bio-compatible products such as functional implant devices, prosthetics, and drug delivery along with tissue engineering with on ground application may be revolutionized. During the fabrication of bio medical implants, the polymer matrices, generally composed of natural origin such as gelatine, collagen, and polypeptides, are mixed with nanofillers which enable the matrix with specific required attributes by acting as molecular bridges (Baiguera et al. 2014). More recently, a fundamental method for fabrication of unsaturated polyester based nanocomposites reinforced with graphene oxide demonstrated an effective peptide bonding when embedded with polyhedral oligomeric silsesquioxanes (POSS). This highly performing nanomaterial structure is proposed to be the composite of choice for electro-technical applications (Divakaran et al. 2020). Hence, such advancements

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in the field of nano-composite materials may significantly help in development of medical grade materials possessing features such as bio-compatibility, functionality, serializability, and manufacturability (Suzuki and Ikada 2011; Levón 2017).

5.3.4

Surface Modifications

Numerous methodologies have been reported to improve surface characteristics including surface wetting, adhesion, porosity, and surface tension (Fig. 5.7). In this chapter, only chemical modification required for bio-composites will be discussed. The chemical modification has been done through addition of some chemical group that eliminate the unwanted part to improve the properties of bio-composite. Alkali treatment (Chung et al. 2018; Ng et al. 2018), Acetylation, Silane treatment, Benzoylation, methyl acrylate found to be popular chemical modifications (Peças et al. 2018). In the last decade, significant works has been conducted on natural fibres and polymeric composites including their characterization, fabrication methods, and mechanical properties along with potential applications. In this chapter, the methods of reinforcement of natural and synthetic fibres or their bio printing have been highlighted along with their perspective applications. Natural fibres possess unique properties and can be obtained by multiple resources using different methodologies. Aspen wood derived cellulose fibres have Mechanical polishing

Photolithography Surface roughness

Plasma etching/spraying

Physical Modification

Sandblasting

Dipping methods Steam treatment

Topography

Surface Modification of Composites

UV radiation Oxidation

Thermal treatment

Surface Coatings

Electrophoretic deposition Sol-gel coating Grafting

Chemical Modification Solvent casting

Vapor deposition

Acetylation

Alkali hydrolysis Glycidyl group addition

Silanization

Fig. 5.7 Different techniques employed for the modification of biomaterials surface properties

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been reported to be surface modified using Maleated Polyethylene (MAPE) as coupling agent. It has been analysed on the basis of morphology, rheology, and mechanical behaviour and it has been apprised that on modification, the tensile properties of fibre matrix coupled with MAPE has been enhanced by 29% as compared to non-modified fibres used for automotive, construction, and packaging applications (Chimeni et al. 2019). Polymeric (rHDPE) composite material reinforced using rice husks (RH) has been prepared via extrusion and compression moulding and subsequently, modified using alkali, acid and Ultraviolet-ozonolysis (UV/O3) treatments. Such modifications of the resulting bio-composite enhanced the tensile strength by 5% (Rajendran Royan et al. 2018). In another study, investigation was targeted to enhance the adhesion during reinforcement of PLA with two different husks, namely, rice and Einkorn wheat husk already treated with alkali and silane. Interestingly, the results demonstrated the lowered moisture sensitivity and increase of energy surface of the husks pre-treated with alkali (which was higher for wheat husk). Thereafter, enhancement of bending modulus along with stress was also featured for bio-composite composed by PLA reinforced with treated husk as discussed above (Tran et al. 2014). Similarly, surface modifications were carried with milkweed fibre using 5% NaOH (at varying time intervals), and subsequently treated with boiling water and some detergent (Sayanjali Jasbi et al. 2017). Thereafter, bio-composite was formulated using polyvinyl acetate (PVA) matrix and milkweed fibre as filler. This investigation reported the enhanced adhesion and tensile properties resulting due to increased surface roughness making strong mechanical interlocking, being highest when treated for 60 minutes. Unfortunately, the tensile and flexural strength declined with 90 min exposure, due to fibre damage because of longer exposure of to the alkali (Sayanjali Jasbi et al. 2017). In addition, Kenaf fibres have been prepared in polypropylene matrix via compression moulding (Asumani and Paskaramoorthy 2020). In another study, Polylactic acid (PLA) was reinforced by microcrystalline cellulose (MCC) for 3D printing has been surface modified using titanate coupling agent to improve the compatibility of composite (Murphy and Atala 2014).

5.3.5

Surface Effects and Characterization

Bio-composite basically comprises of a matrix and a filler or reinforcing material, which may be present on the surface or interface, required to be characterized for further processing. In fact, the surface of the material is mainly responsible for biological and environmental interactions as it is the most exposed part to the tissues. Hence, the surface characteristics need to be well scrutinized by thorough study of matrix and filler along with the interactions existing between them. The geometrical arrangement of the different phases commutes vital challenges in the field of tissue engineering. Several attempts have been made to treat the surface differently from the bulk portion (Agarwal and García 2019). These include the spatial arrangement, fabricating methods, and addition of fillers to enhance the efficacy of the bio-composite surface thereby projecting a promising outcome of bio substitute

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(Agarwal and García 2019). In real practice, one of the material members in bio-composite governs the surface, showing reactivity and bio-compatibility, but in due course of time, the bulk portion becomes the surface in dynamic, which affects the functionality of the incorporated biomaterial. Hence, the selection of materials used as raw material for synthesis of bio-composite is quite a challenging task. Furthermore, the complete analysis of surface characterisation, with or without modifications, using different methods is another crucial task faced by the researchers and is of high concern (Agarwal and García 2019). To cater this issue, the electron spectroscopy for chemical analysis (ESCA), Auger electron spectroscopy (Auger ES), and scanning tunnelling microscopy (STM) are the available techniques illustrating the advancements of technology and precision of the findings. In recent years, tissue engineering involving composite biomaterials have changed the approaches of treatment for the health industry. The gaps, where the conservative treatments failed to manage or restore the complete functionality of the lesions, have significantly been replenished by the strategies involving the concepts of tissue engineering (Davis and Leach 2008). To fulfil the demands of the health industry, retrospection of the success of composites till today has been scrutinized and subsequently, attempts to understand the challenges for their actual on ground applications along with analysing the strategies available to solve the problems so as to ensure their prospects to provide alternatives in managing ailments has been targeted in this study.

5.4

Retrospectives of Composite Biomaterials in Tissue Engineering

The establishments of tissue engineering have prospered towards a golden platform to serve the need of health issues related to developing artificial tissues and organ regeneration. The recent approaches, by way of advanced tissue engineering techniques, implicate the employment of cellular implants and using 3D scaffolds and other matrices to deliver tissue growth promoting factors. The basic principle of such techniques focuses on improving or restoring the tissue functions of the body by establishing novel bio-compatible substitute products or by reconstructing the impaired tissues (Chaudhari et al. 2016). Currently, composite biomaterials are being widely used in formulating products for damaged body tissues. Polymers, ceramics or other composites are being immensely applied in formations of scaffolds, the decellularized framework, wherein the seeding cells may be seeded resulting to the construct of living tissue which is either better or equal to the normal tissue to be replaced in terms of function, structure, and mechanics (Vats et al. 2003; Ganesh 2019). The applications of such composite materials have been recognized in various tissues including dentals, bone, cartilages hard tissues and muscle, skin soft tissues, and organs. This section intends to discuss the applications of composite biomaterials in various tissues/organ impaired conditions and highlight the studies done so far in different models to understand the challenges which needs to meet for success of such materials. Table 5.4 depicts the characteristics and tissue engineering applications of promising composite biomaterials components.

Ceramics

Bio-glass and glass ceramics Calcium phosphate ceramics

Pyrolytic carbon

Zirconia (ZrO2)

Alumina (Al2O3)

Hydroxyapatite

Nitinol (Ni-Ti alloys)

Porosity, better influence on bone tissue regeneration

Higher strength, good fatigue resistance, corrosion resistance, strong osteointegration capacity Corrosion resistance, shape memory effect, pseudo elastic property Hardness, brittleness, higher strength, stiffness, corrosion resistance, low density, electrical and thermal insulator Corrosion resistance, wear resistance, good bio-compatibility, no cytotoxicity Higher mechanical strength, fracture toughness Higher compatibility with bone and other tissues, similar bone mechanical properties, low tensile strength, brittle Porosity and bioactivity

Titanium based alloys

Cobalt based alloys

Orthopaedics

Bone defects

Total hip replacement ball heads, medical devices Orthopaedic implants

Dental implants, hip prosthesis

Dentals, orthopaedics, medical sensors

Dental, orthopaedic, cardiovascular uses

(continued)

Roopavath et al. (2019)

Rath et al. (2014)

Thamaraiselvi and Rajeswari (2004) Thamaraiselvi and Rajeswari (2004) More et al. (2013)

Fernandes et al. (2011) Roopavath et al. (2019)

Kirmanidou et al. (2016)

Narayan (2012)

References Szala and Łukasik (2018)

Biocomposite Metal/ alloy Applications Vascular stents and electrodes Cardiac pacing systems Eg. ATMF138/139, F1314, F1586, F2229 Neurosurgical and vascular implant fabrication Fracture fixation implants Eg. CoCrMo alloys Elgiloy (ASTMF-1058) ASTMF-563 Orthopaedic and dental implants E.g. Ti6

Table 5.4 Characteristics and tissue engineering applications of composite biomaterials components

Characteristics Reasonable strength Fatigue resistance Pitting corrosion resistance Higher mechanical properties Higher spring back property

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Type Austenitic stainless steel

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Synthetic polymers

Biocomposite Natural polymers

Polycaprolactone

Poly L- lactic acid

PLGA

Polyethylene glycol derivatives

Gellan gum

Hyaluronic acid

Chitosan

Alginate

Silk fibroin

Type Collagen

Table 5.4 (continued)

Slow biodegradable, structural flexibility, non-toxic metabolism of its degraded products

Soft, elastic, thermo-reversible, non-toxic, structure similarity to glycosaminoglycan cartilage Swelling under biological conditions, higher permeability to gases, nutrients and metabolites, good bio-compatibility Bio-compatible, biodegradable, mechanical strength, Biodegradable

Characteristics Protein in abundance, low antigenicity, higher mechanical strength, bio-compatible Good mechanical properties, bio-compatibility for cell attachment, establishes adequate 3D porous structure along with mechanical support required for tissue generation (bone and cartilage) Bio-compatible, non-toxic, biodegradable, hydrogel formation Biodegradable, antibacterial, bio-compatible, wound healing, mucoadhesive Hydrogel formation

Scaffolds for cartilage and bone matrix

Bone tissue regeneration, scaffold formation

Scaffold formation for tissue repair

Drug delivery, scaffold fabrication

Degradation product, hyaluronan, may be used as scaffold for cartilage repair Treatment of cartilage defects

Scaffolds for tissue engineering

Scaffolds for cartilage repair

Scaffolds for soft tissue repair

Applications Scaffolds for soft tissue repair

Spadaccio et al. (2009), Deplaine et al. (2013) Jeong et al. (2012)

Nagura et al. (2007)

Oliveira et al. (2010b), Oliveira et al. (2010a) Hui et al. (2013)

Nettles et al. (2004)

Oliveira et al. (2006)

Wang et al. (2011)

Augst et al. (2008)

References Agung et al. (2004)

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Polytetrafluoroethylene

Silicone rubbers Ultra high molecular weight polyethylene (UHMWPE) Polyethylenterephthalate

Poly methyl methacrylate (PMMA)

Manipulation capability

Artificial vascular grafts, implantable sutures, mesh, heart valves Artificial vascular grafts, catheters

Blood pumps, blood reservoirs, dentures, bone cement, ocular lens, membrane for blood dialyser Breast implants As bearing surface in joint arthroplasty

Xue and Greisler (2003) Glickman et al. (2001)

Kappel et al. (2014) Sobieraj and Rimnac (2009)

Oshihara et al. (2017)

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Composite Biomaterials for Hard Tissue Regeneration

5.4.1.1 Bone Tissue Regeneration The mechanical properties of bones vary with the biological locations due to natural adjustments for their composition, porosity, and crystallinity towards biological and biomechanical environment (Pazzaglia et al. 2009). The conventional treatment strategies involved to treat bone defects was the use of allograft or autograft, which due to certain drawbacks such as limited availability, high cost, donor site morbidity, and disease transmission with such treatment strategies impelled the increased use of economically viable bio-composite materials such as bone graft substitute (BGS) which may be capable of supporting the skeletal structures while the healing occurs (Sah et al. 2017). Composite biomaterials are being extensively employed in managing orthopaedic conditions including joint replacements, grafting, and cementing of bones and bone fixing plates. The commonly used materials include 316 L stainless steel, Co-Cr alloys, and Ti-Al-4 V titanium alloys, as they provide ample stiffness and tensile strength. Fibre composites such as carbon fibre, liquid crystalline polymer, polysulphone, and polyetherimide exhibit the advantage of flexibility, adaptability, radiolucency, and non-corrosiveness. The hitch remains to their appropriate fabrication, durability, and bio-compatibility (from carbon debris), although polishing along with coating of such materials with hydroxyapatite has presented some solution partly (Iftekhar 2004). In cases where fracture fixation is demanded, use of bioresorbable and osteoconductive bio-composites, such as tricalcium phosphates and gelatine, helps promoting bone healing along with refrainment from second surgery to remove implants, as was a requisite done. The mini-plates and mini-screws, where matrix is composed of polyL-Lactic acid reinforced by raw hydroxyapatite bestows easier handling, shape modulation (as per site requirement), better bonding to the bone tissue, complete resorbable nature, higher stiffness, and bio-compatibility (Salernitano and Migliaresi 2003). Bone cements incorporating polymethyl methacrylate (PMMA); artificial tendons using polyesters along with hydrogels; artificial ligaments using ultra high molecular weight polyethylene (UHMWPE) or polyethylene terephthalate (PET) reinforced correspondingly by ethylene-butene copolymer or poly-2hydroxyethylmethacrilate (pHEMA); and artificial cartilages constituted with PET fibres and reinforced by PHEMA elucidates the progression in applications of bio-composites (Siraparapu et al. 2013). Calcium phosphates, in virtue of possessing similar properties to bones, have been studied in combinations with polymers, for their applications in bone tissue engineering. Bio-composites formed with PMMA along with tricalcium phosphate illustrated suitability of the material for selective laser sintering because of controlled mechanical attributes and porosity (Velu et al. 2016). Contemporary studies suggest some bio-composites which are looked forward as potentials for bone repair engineering. The composites derived from chitosan and natural hydroxyapatite have been shown to be suitable frame material for artificial bone implants owing to their good bio-compatibility and excellent osteoconductivity

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(Lowe et al. 2016). Similarly, the chitosan scaffolds embedded with hydroxyapatite nanoparticles exhibited well spread morphology along with improved cellular attachment and enhanced proliferation contrast to chitosan alone scaffolds (TheinHan and Misra 2009). The nanocomposites conceived using hydroxyapatite nanoparticles have shown bio-compatibility, thereby strengthening the mechanical assets of such bone grafting composite bio-entities (Dorozhkin 2011). The chitosangelatine based scaffolds possessing nano-hydroxyapatite particles also illustrated similar results (Dan et al. 2016). Naturally, the higher tensile strength along with fracture toughness of the bone is characterized due to presence of flexible and tough collagen fibres which are reinforced by hydroxyapatite crystals (Stock 2015). A dual network hydrogel fabricated from polyvinyl alcohol, an FDA approved composite material, and alginate was synthesized, optimized, and characterized using thermal and spectral analysis. The results presented the formation of tough hydrogels along with controlled swelling and enabling them suitable for developing the bone composite substitutes when incorporated with the ceramic fillers such as bioglass® 45S5 (Shankhwar et al. 2016). The fabrication of hydroxyapatite monoliths showed exceptional strength (compressive modulus: 3.2–4.4 GPa; Compressive strength: 142–265 MPa) which corresponded to the properties of the cortical bone and retained high porosity (>60%) as found in the cancellous bone (Fig. 5.8). Different level of biodegradation and bioactivity is required for different region of bone anatomy and accordingly biomaterials and bio-composites are developed with other controlled parameters of scaffold characteristics. This combination of the properties, thus, propose enabling the repair of larger bone defects structurally, as the higher strength would facilitate to bear the skeletal force, high porosity would enhance the blood vessel infiltration in the monolith, so as to initiate effective healing along with multiplication and differentiation of the human osteoblast like MG63 cells (Meredith 2009).

Fig. 5.8 The desired mechanical behaviour of bone tissues comprised of dense and spongy tissues channelled with vascularization for induced regeneration. The depicted graph is reproduced from thesis of Bio-composites for Bone Tissue Engineering Innovation Report (Rezwan et al. 2006)

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5.4.1.2 Dentistry Composite biomaterials have been exceptionally successful in sphere of dentals where the original tooth material needs to be treated by replenishing its crown, the dental cavities or replacing the entire tooth along with orthodontic arch wires. Different composite biomaterials used for such purpose includes resin monomers such as urethane di-methacrylate ester derivatives, viz. bis-GMA, or a methacrylate along with hard filler particles such as crystalline quartz, calcium silicate, calcium fluoride, silicon nitride whiskers and glass ceramics; and hybrid dental resins (Iftekhar 2004) (Fig. 5.9). The properties which qualify the use of such agents in dentistry includes hardness, resistance to fatigue, fracture or wear, dimensional stability, the ability to sustain varying thermal stress in mouth, retention in the caries, aesthetic satisfaction for colour and gloss match with other teeth. Dental braces are made up of composites including polyethylene matrix reinforced by ceramic hydroxyapatite particles, thus exhibiting isotropic characteristics and better enamel adhesion (Iftekhar 2004). Certain factors reported to degrade the resins and adhesives, as used in treating carious tooth, includes saliva, chewing force, thermal variations, chemical changes, and caries producing oral bacteria such as Streptococcus mutans (Gupta et al. 2012). Research findings exhibited the production of certain by-products which lead to compromise of dentin-resin interface leading to secondary caries (Bourbia et al.

Fig. 5.9 Application of bio-composite as (a) amalgam based tooth filling and (b) ceramic screws and metallic wires illustrated by X-ray of left shoulder joint fixation. Images received with the consent of patient

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2013). Use of silane for the interphase bonding and incorporating nano-microfillers has presented better results (Bayne 2005). Other area of clear need is the generation of crack resistant dental ceramic materials. Composites of alumina with zirconia have been explored to be crack resistant, but strong bonding and aesthetics is still a struggle with their use. Some dental biomaterials based on composites such as pHEMA (2-hydroxyethyl methacrylate), Bis-GMA (2, 20 -bis [4(methacryloxypropoxy)-phenyl]-propane), TEGDMA (triethylene glycol di-methacrylate), and UDMA (urethane di-methacrylate) have been found to be cytotoxic, thereby limiting their application (Gupta et al. 2012). More recently, in some studies, the concept of scaffolds, cells along with signalling molecules for repair or regeneration of dentals in animal model has been exhibited to be successful but still some concerns such as, regulatory issues in adult stem cell collection, need to be addressed for their applications in humans (Bayne 2005).

5.4.2

Composite Biomaterials in Soft Tissue Engineering

5.4.2.1 Vascular Grafting Vascular substitutes are one of the most common strategies involved in coronary and peripheral bypass surgeries in cases such as atherosclerosis. Vascular grafting includes a graft which enhances recellularization by the host’s endothelial cells along with anticoagulant treatment. Currently, synthetic grafts are often used clinically and are generally fabricated from the entities including polyethylene terephthalate, polytetrafluoroethylene, and polyurethane. However, thrombus formation and inflammatory responses are the major drawback along with inability of such materials to be grafted with required size of less than 6 mm, as required to replace radial or mammary artery and saphenous vein (Ravi and Chaikof 2010). More recently, employment of diverse biomaterials entities for fabrication of tissue engineered blood vessels has been performed and studied for their vascularisation effects in in vitro/in vivo analysis. Scaffolds were fabricated by ink-jet printing technique using alginate-collagen biopolymer in a study while using fibrinogen and thrombin with same technique and the in vivo and in vitro investigations of this study resulted in developments of functional vascular tissues (Sarker et al. 2018).. Improved biological functionality of vessels has been reported when scaffolds constituted by fugitive ink-GelMA, carbohydrate glass encapsulated in alginateagarose- Matrigel- fibrin- PEG based hydrogel and PVA-alginate biopolymers has been studied using extrusion based printing (Sarker et al. 2018). Similarly, another strategy where laser based printing was implemented to form scaffolds from GelMA, GelMA-poly (ethylene glycol) diacrylate, polyester urethane urea, polytetrahydrofuran ether-diacrylate, and ceramics (photosensitive and organically modified) exhibited interesting results where enhanced bio-compatibility and similar mechanical features as of body’s own capillaries could be observed (Meyer et al. 2012). Other investigations also reported enthusiastic results for vascular regeneration using other strategies where scaffolds were prepared implying micromodule assembly with collagen/fibronectin coated collagen modules, moulding technique

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with alginate-matrigel composites, and nanofabrication with random PCL/ collagen-PEO nanofibres (Sarker et al. 2018). Use of such bio-composites in vessel reconstruction has paved the path where research can significantly fabricate stable networking of biologically functional vessels capable to anastomosing with patient’s vasculature.

5.4.2.2 Cardiac Tissue Engineering Cardiac diseases, such as cardiac arrests, myocardial infarction (MI), angina, and atherosclerosis, accounts for major illness amongst population and, strategies of tissue engineering primarily focus on myocardium, valves, and coronary grafts. Currently, the mechanical heart valves are constructed from bio-composites such as titanium and carbon, sewn in place of original valve in the heart, along with anticlotting therapies suggested to the patient for life long (Harris et al. 2015). The cell therapy involving usage of endothelial progenitor cells, cardiac stem cells, skeletal myoblasts, mononuclear cells from bone marrow, and embryonic stem cells has been proposed to be effective for managing MI but still the limitations such as long-term safety, poor cell retention, and electromechanical integration are the hurdles which may be taken care by use of composite biomaterial scaffolds delivering such cells to the injury site. Studies involving use of patches made of collagen; scaffolds fabricated from collagen, chitosan, alginate; or hydrogels made of chitosan-collagen, alginate-chitosan, fibrin glue has shown marked repair of infarcts (Cui et al. 2016). The functional use of synthetic composites in managing MI has also been reported in some studies (Cui et al. 2016). Poly(ε- caprolactone), poly(L-lactic acid), along with collagen were used in preparation of nano-scale scaffold wherein rabbit cardiomyocytes were cultured and results comparable to native myocardium exhibited (Mukherjee et al. 2011). To PLGA nanoparticles, insulin like growth factor-1(IGF) was bounded, to observe the effects in rat model, and the results presented IGF-1 retention along with reduced cardiac cell apoptosis and enhanced LV function (Chang et al. 2013). Carbon nanotubes along with chitosan; and carbon nanofiber with gelatin hydrogel have also been investigated in rat model where inhibition on pathogenesis was observed (Cui et al. 2016). The patients with heart failure or conductive defects and blockage are generally being treated with implantation of artificial pacemakers which are usually modelled using titanium or its alloy along with lithium battery and encapsulated in a polymer such as polyurethane (Borcan et al. 2019). Comprehensive research in cardiac tissue engineering has opened up the scope where use of conductive biomaterials such as polyacetylene (an organic polymer with conductive property), polyazulene, polyaniline, polypyrrole, and polythiophene has been characterized in different studies (Cui et al. 2016). Exhibition of properties such as electrical conduction, easy synthesis, biodegradability, and bio-compatible nature ensures them to be the choice of material for formulating drug delivery device, biosensors, neural implants, along with scaffolds for tissue engineering. Studies exploring diverse potentials of such conductive, as in case of fabricating biological pacemakers, are on progress, and may replace the artificial pacemakers in the coming times (Cui et al. 2016).

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5.4.2.3 Contact Lens and Cornea Hydrogels synthesized by the co-polymerisation of various synthetic bio-composites are applied in making contact lens and cornea. The properties that satisfy the needs of such eye lens implant preparations include bio-compatibility, softness, oxygen diffusibility, high refractive index, transparency, and modulus. The hydrophobic monomers such as silicon possessing monomers, methyl methacrylate, perfluoro polyethers along with hydrophilic monomers such as methacrylic acid, dimethylacrlamide, and N-vinylpyrrolidone are potentially used in designing of contact lens (Patel and Mequanint 2011). The recent progress can be witnessed where artificial cornea has been engineered using PMMA and titanium and has provided vision to several patients with corneal opacity and graft failures (Sikora et al. 2019). 5.4.2.4 Neural Tissue Engineering Nervous tissue is an integral but highly complex tissue where large numbers of patients are affected by its functional disruption due to multiple causative factors such as accidents, wounds, and birth defects. Currently, no medical treatment has registered success in managing repair of injured CNS and presents a challenge for neurobiologists, although the proximal segments of PNS does exhibit healing by regeneration of affected nerve fibres. Comprehensive management therapies rely on stabilization and prevention approaches such as orthopaedic fixation of the unstable spine followed by rehabilitation and use of prosthetics. Current research in this area focuses on studies using different composite biomaterials so as to create resorbable, bio-compatible, easily available, oxygen porous, and functional artificial nerve grafts that may help in this tissue regeneration by supporting cellular growth due to possession of oriented substrate (Huang and Huang 2006). The oriented substrate helps cell adhesion, proliferation along with nervous impulse transmission (Ermis et al. 2018). Different studies suggest some composites, the combinations of which provide suitability for nerve regeneration. Such composites include the combination of aligned fibres of polycaprolactone (PCL)-gelatin, PCL-chitosan, PCL-collagen, PLLA, and PLGA by electrospun strategies. In a comparative study, the aligned fibres exhibited improved cellular orientation, enhanced neurite outgrowth along with contact guidance when compared to randomly oriented fibres (Cooper et al. 2011). Chitosan, collagen, gelatin, poly (organophosphazene), poly (glycolide-cotrimethylene carbonate), polyurethane, and poly glycosamineoglycan-co-collagen has been used to compose scaffolds and studied in rat model for neural tissue repair (Huang and Huang 2006). Similarly, scaffolds incorporating polyglycolic acid has been studied in monkey, rat, rabbit, and human model for the purpose of neural repair (Huang and Huang 2006). PGA mesh coated with collagen was studied in dog model and the results present the scope of applications of such biomaterials in managing nervous disorders (Huang and Huang 2006). Currently, a change in trend has headed where the nerve guidance channels (NGC) using composites with sophisticate bio-fabrication techniques are being looked upon for better nerve repair results (Papadimitriou et al. 2020). Contemporary researches have exhibited

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various strategies which are undertaken to develop ideal nerve guide from bio-composite materials. Such fabrication techniques include injection moulding, solid free form fabrication, magnetic polymer fibre alignment, phase separation, micro patterning, ink-jet polymer printing, lyophilizing, extrusion, electrospinning, film rolling, braiding, and salt leaching (Nguyen et al. 2015). Table 5.5 presents different composite materials which have been selected by the researchers to have an insight for possibilities for neural tissue regeneration. In the neural tissue engineering sphere, studies involving varieties of natural and synthetic composites have been exhibited in Table 5.5. Their role in repair and regeneration of this complex tissue has added knowledge for the scientific community. The innovation of nerve guidance conduits using composites have shown light for making treatment possible but still an ideal conduit possessing bio-compatible, bioresorbable characters along with properties of supporting neurite extension to enhance nerve regeneration and minimizing interactions between axon growth and myofibroblast is still to be developed. Moreover, very limited number of FDA approved conduits such as Nerve Cuff, Neuroflex™, Axoguard™, Neuragen™, and SaluTunnel™ are available in the market to be used for small gaps (3 mm or less), which emphasizes on the need to look for newer materials which can potentially be used in the patients requiring neural tissue repair and regenerates (Papadimitriou et al. 2020). Similarly, the advancements in the tissue engineering strategies led to a height where the patients are benefiting such as implantation of encapsulated pancreatic islets for diabetics and encapsulated hepatocytes for liver failure.

5.5

Bottlenecks of Composite Biomaterial Applications

Contemporary researches in this field of composite biomaterials are being vastly studied and have touched every part of the human body, but complete success for each of them is still a challenge due to certain drawbacks which hinders the formed product to be available for patient use. The biomaterial science furnishes composites made of metals and its alloys for taking care of skeletal system issues such as fractures and tooth problems, but challenges with metal implants such as untimely degradation of resorbable metals such as Mg and Fe, leading to early loss of mechanical strength; hypersensitivity and osteopenia with long presence of metals such as alloys of Co, Fe, Ti, and Cr; infection/inflammation leading to complication of the issue; fatigue and loosening of the load bearing implants; intrinsic brittleness; and cytotoxicity is a concern. Further, the scope for metallic scaffolds are visualized to be very low, as these cannot be loaded with bioactive cells or molecules, leaching of metallic ions may be carcinogenic, and corrosion behaviour affects the selection and application of the metals (Siraparapu et al. 2013; Prasad et al. 2017). The global challenge with respect to disposal of non-biodegradable polymers including plastics is well known (Testin and Vergano 1996). Hence, there is a need to develop and use such composite materials which are environmental conscious and fulfil the patient needs. To address this issue, biodegradable polymers have been

Compression/ injection moulding Fibre mesh/fibre bonding

Electrospinning

Moulding and texturing methods

Electrohydrodynamic techniques

PHEMA-MMA-EDMA-APS-SMBS

Hydrogel formation

PLGA-polyurethane

PLGA/PCL blend, PCL/EEP blend

PLLA/THF solution

PLGA-Polypyrrole

Poly-D, L-lactic acid

PLGA/Pluronic F127

Collagen-glycosaminoglycan

Freeze-drying

Promote cell proliferation, differentiation, exhibited degradability, conductivity rat sciatic nerve model Nanofiber to produce nerve grafts, rat sciatic nerve regeneration, tube stability, no inflammations Rat sciatic nerve model, effective regeneration and ample electrophysiological recovery Enhanced growth of PC12 and S42 cells required for peripheral tissue regeneration

Neuro2a cells showed growth on the conduit Rat sciatic nerve, regeneration of myelinated and unmyelinated axon over 10 mm 60% of regeneration of male rats sciatic nerve Significant regeneration of male rats sciatic nerve Improvement in Schwann cell alignment

Chitosan

(continued)

Kim et al. (2016)

Chew et al. (2007)

Sun et al. (2012)

Jing et al. (2018)

Miller et al. (2001)

Dalton and Shoichet (2001) Oh et al. (2008)

Chamberlain et al. (2000)

Ao et al. (2006)

Technique of Bio-fabrication of scaffold Conventional method Solvent casting/ particulate leaching Phase separation

Reference Kokai et al. (2009)

Table 5.5 Fabrication methods for bio-composites development in neural tissue engineering Prospects Maintains glucose permeability required for cell growth in scaffold

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Biomaterial used Polycaprolactone (PCL) PCL/PEO solution

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Solid freedom fabrication

3D bioprinting

Selective laser sintering Fused deposition modelling

Technique of Bio-fabrication of scaffold

Table 5.5 (continued)

Mouse bone marrow stem cellsSchwann cells-agarose Silicon tubes-hydrogel drops (neurotropic factors)-alginatemethacrylate gelatine PEGDA

Double layer polyurethane-collagen PCL

PLA-PGA-PLGA

PLGA microparticles

Polysialic acid-PEO along with PGA/ PLA copolymer

Biomaterial used PLLA/PLLA-fibronectin fibres, agarose/ methylcellulose gel

Prospects Infiltration of microglial, macrophages and astrocytes towards this implant in rat striatum model Cell proliferation observed along with generation of myelinated and non-myelinated axon and blood vessels Scaffolds exhibited to support as well growth of Schwann cell in rats in vitro 3D microstructures prepared and new method for creating biodegradable implantable devices 3D nerve conduit prepared Porous NGC supported proliferation and differentiation of PC12 cells in vitro Decent regeneration exhibited for rat sciatic nerve model Regeneration of injured complex nerve in rats Higher sciatic nerve regeneration observed in rat with single lumen conduits

Johnson et al. (2015)

Owens et al. (2013)

Cui et al. (2009) Vijayavenkataraman et al. (2019)

Valmikinathan et al. (2008) Yamada et al. (2008)

Jha et al. (2011)

Reference Rivet et al. (2015)

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introduced by the scientists where different polymers such as poly lactic acid, polyhydroxyalkanoates, cellulose acetate, co-polyester, polycaprolactone, poly ester amide, and polyglycolic acid are available, but their higher cost as compared to conventional plastics remains a concern. The bio-composites composed of natural ligno-cellulose fibres, viz. cotton, jute, hemp, coconut, and flax have also been investigated for their properties. Although some of the natural fibres did exhibited best mechanical strength but low thermal resistance, tensile and flexural strength, natural degradation, and shrinkage restrict their use in matrix formation (Mohanty et al. 2000). Ceramic and other such materials are relatively in lesser use than metals, alloys, and polymers because of their poor behaviour with tension stress and sensitivity towards presence of cracks. The scope of this category of biomaterials has highly been valued in dental and orthopaedic applications because of properties involving compressive strength, toughness, and flexural strength as owned by aforesaid materials. However, the challenge remains with the formation of scaffolds with such materials with appropriate porosity as required managing the injured tissue along with mechanical properties, such as with use of forsterite Mg2SiO4 ceramics for load bearing joints owning to its lower apatite formation and poor degradation ability (Ni et al. 2007). After successful makeup of the novel material using composites, their actual implication in the body needs to undergo and comply with certain tests such as cytotoxicity, mutagenicity/genotoxicity, local reactions, and sensitization which are complicated complex and a challenge in itself (AL-Oqla and Omari 2017). The material needs to comply with ISO 10993 series as developed for all the medical devices (Schmalz and Galler 2017) or Food and Drug Administration regulations such as 510 K (Jammalamadaka and Tappa 2018). Such regulations ensure the risk assessments for patient safety and environmental concerns (Schmalz and Galler 2017).

5.6

Prospects of Composite Biomaterials

With an endeavour to tackle the drawbacks exhibited by use of metals and its alloys, as discussed earlier, Bulk metallic glass and Shape metallic glass are the two new classes where metals have displayed better competence as compared to conventional ones. High toughness and strength, better elastic strain, low elastic modulus, resistance to corrosion, bioresorbable nature, and amorphous shape of BMG have perspectives and formation of scaffolds involving BMG with 3D printing along with its in vivo effects is still the area of future study. Shape memory alloys, such as Nitinol, are another class which possess the property of reversible phase transition, super-elasticity, and other desired mechanical characteristics (Prasad et al. 2017) which may resolve the problems associated with the use of metals and alloys for tissue engineering applications in near future. Iron oxides with magnetic property are also being addressed in the field of biomaterials as by use of magnetic nano-particles or as polymeric matrices such as hydrogels. The perspectives of these materials are

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being believed to be the future tools for drug delivery, availability of growth factors and magnetic scaffolds which can be targeted to the required size by the influence of external magnetic field (Gil and Mano 2014). Different natural polymers have gained the attention of scientific community for their usage in tissue engineering. Different polysaccharides including chitosan, cellulose, alginate, and hyaluronan are in applications, still the other less studied source like starch and nanofiber furnished from this polysaccharide (high porosity and surface area, similar to ECM component structure) needs to be investigated exhaustively. Formation of 3D scaffolds from such polysaccharides using recent technology such as topochemical engineering techniques and subsequently, their analysis at molecular levels to peep the interactions of materials, cells and biomolecules in the relevant physiological environment (Tchobanian et al. 2019) may contribute considerably in managing injured tissues. Products involving synthetic polymers such as PLA and PGA, as discussed earlier, are consistently found in the market. Presently, decellularised scaffolds, hydrogels, and self-assembled cellular sheets fabricated from different polymers are the interest of investigations, such as in cardiovascular tissue engineering. Hybrid scaffolds involving combinations of natural and synthetic polymers, including conductive polymers, are holding a great promise for development of wide sphere of functional scaffolds because of their capacity to modify in terms of mechanical, biochemical, and electrical potentials (Theus et al. 2019). To secure the usability of ceramics and bio active glass in regeneration therapy, different strategies are being adopted to modulate their properties. Modulation of ceramic surface by coating of polymers over it; changes in composition of sol-gel materials; processing of the powder based bioactive glass along with ceramic products using advanced technology, such as additive manufacturing, may help immensely in improving the functionality of such biomaterials (Kaur et al. 2019). Intensive research is being conducted where the scientists are immensely working in search of a perfect biomaterial that satisfies all the requirements for use in living body. Currently, computational science and approaches of omics in the field of biomaterials has significantly shown a path where the complexity of composites may be better understood along with acknowledging the interrelations between the properties of involved biomaterials to their effects on complex living systems. Such approaches may help to limit the “trial and error” methods and indeed would facilitate to develop innovative biomaterials for their relevant employment in tissue engineering system (Groen et al. 2016).

5.7

Conclusion

The accomplishment of functionality of tissue engineering techniques critically depends on the composite biomaterials, as selected for the repair and healing of the impaired tissues. These composite biomaterials, in virtue of their commendable properties, have shown a path to formulate products, the applications of which may be sensed by the shift in paradigm for approaches towards the management of

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different ailments especially in the dental and orthopaedics field. These composite materials have greatly helped in improving the lifestyle and life expectancy of such patients. Practically, after every 30 seconds, a death is recorded due to certain diseases, where a patient life could have been saved by the tissue replacement (Ganesh 2019). This emphasizes the need of immense research done till today for composite biomaterials, to get transformed into the shape where such materials, by way of tissue engineering processes, may safeguard the life of patients especially involving the soft body tissues. Comprehensive research indicating the understandings related to bioelectric signalling across the cells at molecular level have opened the platform for formulating the next phase generation of biomaterials such as implanting devices with biosensors. Still, different challenges such as bio-compatibility, non-degradability, leaching, and regulatory issues does create a wall but newer studies, such as smart composites and magnetic composites must be worked with a passion for patient reach outcomes. To achieve such heights, the collaboration of work study of medical scientists, engineers, and doctors is the demand of time. Nevertheless, such remarkable innovations are highly being looked upon by health industry which have shown promising aids for tissue repair and have heightened the probabilities of success in medical practices. Acknowledgement The authors acknowledge the support received from TEQIP Phase-III of Dr. B. R. Ambedkar National Institute of Technology, Jalandhar, Punjab, for the completion of this chapter. Conflict of Interest The authors declare that there is no conflict of interest regarding the publication of this article.

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Part II Trends in Biomaterials

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Trends in Bio-Derived Biomaterials in Tissue Engineering Dimple Chouhan, Sharbani Kaushik, and Deepika Arora

Abstract

Biomaterials have become indispensable for tissue engineering applications including the development of artificial grafts, implantable devices and drug delivery systems. Amongst all, the biomaterials derived from natural resources, generally termed as “bio-derived biomaterials” present a sustainable and greener route for developing scaffolds and implants for artificial grafts, with wide scope of processing them into tailor made supplies. Such materials are often preferred over synthetic counterparts owing to their physiological relevance, inherent cell– material interactions and biocompatible properties. The nature holds a great treasure of numerous such materials that have been extensively utilized for regenerative therapeutics since ages. The bio-derived biomaterials can be obtained from microorganisms, plants, marine creatures and animals. Being nature derived, these materials can mimic the structural and functional aspects

Dimple Chouhan, Sharbani Kaushik and Deepika Arora contributed equally with all other contributors. D. Chouhan (*) Department of Biosciences and Bioengineering, Indian Institute of Technology Guwahati, Guwahati, Assam, India e-mail: [email protected] S. Kaushik Department of Chemistry and Biochemistry, The Ohio State University, Columbus, OH, USA D. Arora Biosystems and Biomaterials Division, National Institute of Standards and Technology, Gaithersburg, MD, USA Skeletal Biology Section, National Institute of Dental and Craniofacial Research, National Institutes of Health, Department of Health and Human Services, Bethesda, MD, USA # The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 B. Bhaskar et al. (eds.), Biomaterials in Tissue Engineering and Regenerative Medicine, https://doi.org/10.1007/978-981-16-0002-9_6

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of the human tissues. The present book chapter gives a brief overview of the bio-derived biomaterials ranging from microbial derived biomaterials to animals/ plants derived proteins and polysaccharide-based biopolymers. Animal origin biomaterials such as collagen, gelatin, fibrin and hyaluronans have contributed significantly to the success of tissue engineering so far. Special coverage has been laid on decellularized extracellular matrix and its tissue regenerative properties highlighting the role of nature’s template in engineering bioactive constructs. Additionally, insect derived silk and chitosan-based materials are also briefly described along with a few polysaccharides such as alginates, agarose and carrageenan extracted form algae and marine seaweeds. Furthermore, microbial derived biomaterials have been discussed with a few representative model biopolymers, underlining their biosynthesis, purification and their biocompatible properties that make them versatile to aid tissue recovery and/or replace their functionality. These biomaterials provide an impressive 3-dimensional microenvironment to culture living cells while supporting guided differentiation, extracellular matrix secretion and tissue regeneration. With an aim to highlight the role of bio-derived biomaterials in tissue engineering applications and allied fields, the present book chapter provides an insight into their progress in healthcare market and future applications. Keywords

Natural biomaterials · Proteins and polysaccharides materials · Decellularized extracellular matrix · Plant derived scaffolds · Microbial biopolymers · Tissue engineering

Abbreviations 3D γ-PGA ε-PL BC CaCl2 Ca(PO4)2 ChitoMA CRG CS DECM ECM EDTA FDA FGF-2 GAGs GelMA

three-dimensional poly(γ-glutamic acid) poly(ε-L-lysine) bacterial cellulose calcium chloride calcium phosphate chitosan oligomer methacrylate carrageenan chondroitin sulphate decellularized extracellular matrix extracellular matrix ethylene diamine tetra acetic acid food and drug administration fibroblast growth factor glycosaminoglycans gelatin methacrylamide

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HA Hap LDV NMSF PEG PHA PHB PHBV P4HB RDT RGD rhBMP-2 SF SS TGF-β3 TiO2 VEGF

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hyaluronic acid hydroxyapatite Leu-Asp-Val non-mulberry silk fibroin polyethylene glycol polyhydroxyalkanoates polyhydroxy butyrate 3-hydroxybutyrate-co-3-hydroxyvalerate 4-hydroxybutyrate recombinant DNA technology Arg-Gly-Asp recombinant human bone morphogenetic protein 2 silk fibroin silk sericin transforming growth factor-beta3 titanium dioxide vascular endothelial growth factor

Introduction

Bio-derived biomaterials imply the materials derived from biological and natural sources like plants, animals, insects or microorganisms. These materials have been employed for tissue engineering and regenerative therapeutics owing to their biocompatible properties. The pressing health problem of organ failure and trauma cases demand tissue substitutes, which has been challenging due to unavailability of donor grafts and graft rejection. Biomaterials offer a potential solution to this unmet need by providing biomimetic structural and functional properties for the development of artificial tissue constructs. Tissue engineering aims at restoring/regenerating a diseased/injured tissue by implanting temporary or permanent tissue substitutes into the organism. It is achieved using scaffolding biomaterials that generate a suitable graft by supporting the tissue growth and biological functions (Badylak 2007). The developed construct can also be loaded with healthy cells that strategically improve the tissue functions (Howard et al. 2008). Herein, bio-derived biomaterials play a major role in the construction of a scaffold by providing an architectural support for tissue engineering applications. A tissue comprises of group of cells that self-assemble to establish their framework by secreting extracellular matrix (ECM) components. Tissue engineering deals with the fabrication of artificial grafts by using scaffolds made up of either a single type or combination of biomaterials. The scaffold is a porous three-dimensional (3D) biomaterial based construct that acts as an artificial matrix at the interface of biological systems that aids in cell proliferation and assembly of ECM secreted by cells (O’Brien 2011). The biomaterials used to fabricate the artificial constructs play an important role in deciding their structural and functional properties specific to the

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organ. Therefore, research in the fabrication and design of scaffold has flourished over the last few decades with advancement in the interlinked disciplines including biomaterials science, gene therapy, cell biology, biotic-abiotic surface characterizations and imaging (Chouhan et al. 2019b) (Nair and Laurencin 2007; O’Brien 2011). Biopolymers as biomaterials have received considerable attention over ceramics, alloys and metals owing to their versatility of mechanical and degradation modulating properties. Natural polymers (e.g. collagen, chitosan, hyaluronic acid, fibroin) are amongst the preferable clinically used materials due to their better biocompatibility and lower toxicity while synthetic polymers (e.g. polyvinyl alcohol, polylactic acid, aliphatic polyester polycaprolactone, etc.) have also been realized to be highly versatile, reproducible and workable. However, synthetic polymers suffer from the absence of cell recognition sites that result in sub-optimal cell adhesion, heightened hydrophobicity that leads to partial seeding of the scaffold, and toxic influence of the products generated under acidic environment (Nair and Laurencin 2007). Biodegradable polymers have the capability to dissolute into harmless products in vitro. Additionally, they can physiologically degrade in vivo without or with minimal metabolic transformations (Katti et al. 2002). Noteworthy, these biologically derived materials or bio-derived biomaterials possess inherent biodegradability and relevant biological properties that have spawned their enormous utilization in the field of tissue engineering. Remarkably, the natural biomaterials were the first biodegradable biomaterial to be used clinically (Nair and Laurencin 2007). Natural or bio-derived biomaterials are preferred candidates in tissue engineering due to a myriad of incentives such as ability to introduce receptor-binding ligands to cells, bioactivity and degradation on exposure to proteolytic enzymes (Nair and Laurencin 2007). These materials hold inherent ability to interact with host cells and the physiology of wounded tissues. Till date, numerous bio-derived biomaterials have been explored for healing damaged tissues and engineering artificial tissues. For example, decellularized ECM from animal origin and extracted biomaterials like collagen and fibrin have been utilized for various tissue engineering applications because they mimic the native ECM structure (Hong and Stegemann 2008). Likewise, biomaterials extracted from various plants, insects and microorganisms have exhibited promising results in wound healing, tissue engineering, drug delivery and regenerative medicine (Suarato et al. 2018). The expanding field of exploring and producing biomimetic materials aims at developing functional tissue engineered grafts and artificial organs. The present chapter intends to present a brief overview of the bio-derived natural biomaterials and highlight their properties suitable for tissue engineering applications.

6.2

Concept of Bio-Derived Biomaterials and their Applications in Tissue Engineering

A biomaterial is regarded as a biocompatible material that is suitable to be implanted in human body either temporarily or permanently to repair or replace the damaged tissue (Keane and Badylak 2014). The bio-derived materials are often processed to

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make them non-immunogenic so that the host system readily accepts the implanted grafts. For example, biomaterials extracted from animals or cadaveric source are processed through the steps of decellularization in order to remove the source cells and immunogenic factors from the ECM of the tissue (Badylak 2007). Biomaterials extracted from animal origin such as collagen, fibrin, hyaluronic acid (HA) possess cell conducive cues, which ultimately aid in the key events of tissue repair process. The inherent ability of such materials to recruit cells also helps in vascularization of the neo-tissue for a continuous supply of blood and nutrients (Badylak 2007). Apart from this, biomaterials like fibrin hold certain domains that are capable of sequestering growth factors and thereby assist in tissue regeneration at a faster rate (Hong and Stegemann 2008). The concept of applying natural biomaterials in tissue regeneration began with success of decellularized ECM, which further led to the exploration and isolation of individual ECM components. Most of the natural biomaterials are biodegradable in nature; hence, there is no necessity of a secondary surgery to remove the implanted graft. On the contrary, synthetic biomaterials have low or minimal biodegradability and the implanted material regularly fails to remodel along with the neo-tissue formation (Bhardwaj et al. 2018). Bio-derived materials are mostly formed of proteins or polysaccharides, which are bioresorbable in nature. Therefore, the biomaterial gets slowly degraded with time and their degraded products such as amino acids and sugars are automatically resorbed by the body. Biodegradability is an important concern while selecting a biomaterial (Bhardwaj et al. 2018). The process of tissue repair and regeneration demands a biomaterial that remodels continuously and thereby forms a neo-tissue at the site of damaged tissue. The bio-derived materials being biodegradable thus play an important role in the applications of tissue engineering and regenerative medicine. Bio-derived biomaterials have been extensively used in organ reconstruction, plastic surgeries, wound healing and various therapeutic applications like drug delivery and cancer treatment. A range of biomaterials extracted from plants, animals, insects and microorganisms have been explored for numerous tissue engineering applications (Mogoşanu and Grumezescu 2014). Scaffolds made up of biomaterials act as a platform for cells to reside and populate at the site of injury in order to reconstruct the damaged tissue. Herein, the properties of scaffolding materials play important roles such as integral stability, bioresorption ability, biocompatibility, mechanical strength and bioactivity. In this context, it is worth mentioning that the natural materials are advantageous over synthetic materials because bio-derived materials closely mimic the microstructures and properties of the ECM of human tissues. Plant derived scaffolds are also well-known for stimulating tissue regeneration process. Recently, plant based materials from decellularized fruits and vegetables have shown great success as a scaffolding material (Lee et al. 2019b). Apart from this, some algal polysaccharides have shown great potential as biomaterials. For instance, alginates are found in the cell walls of brown seaweeds. Several bacterial species are also capable of producing alginate as a capsular polysaccharide. Agars and carrageenans are obtained from galactan polysaccharide present in the red algae Rhodophyceae. These algal polysaccharides are exploited for medical applications owing to their impressive gelling, water retention, viscous and stabilizing properties (Mano et al. 2007).

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In addition, insect origin materials like silk and chitosan have also been proven to be outstanding biopolymers that could be made available at low cost with abundant resources (Chouhan and Mandal 2020; Kaushik et al. 2016). Furthermore, microbial derived biomaterials have been developed in recent times to produce the materials with desired properties. The advanced recombinant DNA technology (RDT) has benefitted the material research field with producing microbial derived biomaterials (Báez et al. 2005; Czaja et al. 2006). For example, cellulose derived from microbes could be extracted with a high yield with the help of biochemical plants. The cellulose biomaterial produced in such a way has been explored for tissue engineering and wound healing applications (Czaja et al. 2006). In addition, the protein materials produced from microbes could also be functionalized at the gene level to include a functional motif along with the biomaterial. For instance, artificial spider silk produced by microbes has been easily functionalized with cell binding motif, growth factor motif and antimicrobial peptide (Chouhan et al. 2018c). Such biomimetic materials can be easily produced with the help of RDT under microbial systems like bacterial, fungal and yeast systems. Herein, the whole idea of this chapter is to provide a wide picture of the bio-derived biomaterials that are being utilized for tissue regeneration therapies.

6.3

Decellularized Extracellular Matrix (DECM) as Biomaterials

The development of bioengineered organs using decellularized materials is a potential long-term goal to overcome the problem of donor shortage. Decellularized extracellular matrices (DECMs) have been recognized as a useful biomaterial that conserve a tissue’s native milieu, endorse cell adhesion/proliferation, and offer physiological cues to various cellular function (Badylak 2007; Hussey et al. 2018). In the tissue engineering field, the polymer biomaterials have always been a subject of great interest, as they offer important assortments in control of topography, morphology and chemistry as reasonable substitutes. The DECM materials have been credibly designed by removing cellular, nuclear matters and integrating the different functional groups. These groups were added into the molecular chain of the polymer to control physical, chemical and biological aspects and to imitate the tissues or organs characteristics. Basically, the ECM is the three-dimensional network of cellular macromolecules (collagen, enzymes, polysaccharides and glycoproteins), that provide essential structural framework and biochemical support to cells in a defined tissue architecture (Frantz et al. 2010). The understanding of tissue/organ microenvironment is important in tissue engineering, which often has been achieved by mimicking the configuration of the ECM. The primary purpose of construction of biomaterial scaffolds is to provide the structural stability to the cells in the 3D topography (O’Brien 2011). It not only facilitates spatial distribution but also provides a suitable native environment to cells to regenerate, as it maintains the cell–cell or cell–ECM interactions. To accomplish the effectivity of suitable scaffolds, various synthetic or natural polymers have been used, based on their vast diversity of properties and bioactivities. However, the potential immune response or toxicity associated with certain synthetic polymer combinations limit their use.

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The approach of using DECM based scaffolds has paved its way more lately, that aims to replicate the tissues or organs characteristics. Here the process of “decellularization”, i.e. removal of cellular components from a tissue is involved where the ECM remains preserved, which would further be employed as a scaffold to develop specific artificial grafts. Plenty of methods have been developed for generating DECM materials. Major categories include physical, chemical and biological. The ability to innately restructure the DECM systems residing their appropriate topographies is an imperative feature associated with DECM technology. It offers manifold advantages (a) fundamental insight into chemistry–structure– function associations, (b) direct utility of biopolymers without the risk of immune response, and (c) highly capable to control cell responses and functions (Andorko and Jewell 2017; Cramer and Badylak 2019). DECM as biomaterials can unswervingly influence the functional attributes of bioengineered tissues; therefore, widely recommended for their continuous exploration/usage in the tissue engineering and regenerative medicine field.

6.3.1

ECM and Decellularization

The term decellularization refers to the process employed in bioengineering field to segregate decellularized ECM as a template for developing an artificial organ and helps in tissue regeneration (Crapo et al. 2011). Typically, the removal of potential immunogens or antigens and the use of natural ECM fibre structure are the main aim so that the rate of cellular acceptance would be maximum. This physiologic framework allows cells to perform their functions as it has inherent signals for their attachment, proliferation, differentiation and migration (Midwood et al. 2004). Altogether, the DECM materials uphold the original composition of their native phenotype, accordingly provide a suitable intrinsic microenvironment for artificial tissue/organ development and function. At present, in the field of regenerative medicine the cohort of different DECMs and their integration with composed polymer scaffolds is gaining popularity. Decellularization of allogenic or autologous ECM and their modification recommended as the ideal materials/scaffolds for recellularization of autologous human stem cells potentially lead to the advanced personalized therapeutics and clinical approaches (Noor et al. 2019).

6.3.2

Methods of Decellularization

DECM materials are generally prepared by isolating them through different types of decellularization methods (Gilpin and Yang 2017). The properties and applications of DECM biomaterials are also dependent on the decellularization methods. Broadly, these methods can be categorized into chemical, biological, physical and their combinative approach. The biological methods for decellularization are primarily mild in their actions, i.e. removes cellular and nuclear components without effecting ECM composition and configuration. Acids (e.g. peracetic acid) and

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alkaline (e.g. sodium hydroxide) treatment solubilizes the cell membrane, cytoplasmic and nuclear material by utilizing their intrinsically charged properties (Gilpin and Yang 2017; Goissis et al. 2011). Detergents, (ionic/non-ionic/zwitterionic) are widely used/studied procedure of decellularization. They lyse the cells through disturbing the phospholipid cell membranes and remove cells and genetic material (Vavken et al. 2009). The physical methods in decellularization process are primarily intended to target the cells near to membranes and ECM. When these physical methods employed by combining with other biological and chemical methods, the enhanced efficiency of extraction of dECMs is recorded (Ott et al. 2008). Commonly used decellularization procedures are shown in Table 6.1. Although the basic purpose of decellularization process is to remove both cellular and nuclear components with minimal loss of protein configurations, but at the same time the choice of the modifications to be followed depends upon the subject of the study and target tissue. Limited potential for recellularization, loss of native conformation, damage to cell binding ligands during extraction processes are the crucial factors that could impede the success rate of any particular decellularization process (Fernández-Pérez and Ahearne 2019). Interestingly, their findings divulge some positive facts of extraction methods and their combinations (Fig. 6.1); however, in most of the cases these methods have been found to be limited in tissue-specific manner.

6.3.3

Regenerative Properties of DECM

The tissue regeneration process not only requires a structural platform for cells to reside in the construct, but also needs a favourable microenvironment to stimulate the residing cells towards reparative pathways. The bidirectional crosstalk between cells and tissue-specific ECM is well-known for the functioning of any healthy tissue (Hussey et al. 2018). The cells constantly monitor and control the ECM production by secreting proteolytic enzymes like matrix proteases, elastase, collagenase and many other enzymes that degrade the already deposited ECM fibres (Londono and Badylak 2015; Martin 1997). This prevents the hypersecretion of tissue matrix that could lead to tissue fibrosis. Further, abnormal degradation of ECM is also controlled at the cellular level by manipulating the secretion of degrading enzymes. Thus, the two-way crosstalk or dynamic reciprocity between cells and ECM is a perfect way to maintain a tissue both structurally and functionally. The DECM materials have been found to promote macrophage polarization towards regenerative M2 phenotype that aid in the constructive remodelling towards tissue regeneration (Badylak 2019; Zhu et al. 2018). Therefore, using DECM materials as implantable constructs provides a great advantage in the tissue regeneration process. The microenvironment of native ECM not only provides chemical or biological cues, but also delivers mechanical signals or mechanotransduction signals to the cells (Jansen et al. 2017). Native ECM also serves as reservoirs of bioactive factors such as growth factors, cytokines and chemokines, hence are able to stimulate cells (Wilgus 2012). The DECM materials might lack some of these bioactive

Chemical

Detergents

Non-ionic (e.g. sodium dodecyl sulphate)

Alkaline-acid treatment (e.g. peracetic acid, sodium hydroxide, calcium hydroxide, and ammonium)

Antibiotics (e.g. penicillin, streptomycin, amphotericin)

Nucleases (e.g. deoxyribonuclease and ribonuclease)

Efficiently remove cells and 90% genetic material preserve the

Work well with thinner tissues

Provides healthy environment

Minimizes immunological responses

Sometimes damaging to the structural and signalling proteins

Harsher Stiffness of ECM may increase

Sometimes creates hurdle for clinical biologic scaffolds

Alone may not completely remove all cellular debris

Disadvantages Collagenases sometimes break the peptide bonds in collagen

(continued)

O’;Neill et al. (2013), Zhou et al. (2010)

Gilpin and Yang (2017); Goissis et al. (2011)

Crapo et al. (2011); Gupta et al. (2017)

Crapo et al. (2011), Kim et al. (2019)

References Waldrop et al. (1980), Wicha et al. (1982)

Type Biological

Advantages Prevent ECM damage

Table 6.1 Decellularization methods used in tissue engineering Effects Act on proteins by hydrolysing peptide bonds. Extracellular deposition of collagen fibrils. Matrix remodelling Act directly on DNA and RNA chains, respectively, to hydrolyse phosphodiester bonds Promote the fragmentation of residual DNA, nucleotides Helps in detachment of cells from their ECM Microbial disinfection/ decontamination Solubilizes cytoplasmic components of cells and nuclear material by utilizing their intrinsically charged properties Helps in cell removal, prevent denaturation of the collagen matrix Removing the cellular material, alters microstructure

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Physical

Type

Induces agglutination of DNA when used without DNase, solubilize cell membrane, removes nucleic remnants and denatures proteins Contains non-ionic and ionic detergent properties Maintains ultrastructure and ECM proteins Cell lysis via osmotic shock preserve architecture Reduction in number of cell nuclei, reduced GAG Bind metallic ions, thereby disrupting cell adhesion to ECM Works better with hypertonic/hypertonic solutions Tissue lysis, maintains ECM proteins and mechanical properties, remnant DNA

Zwitterionic (e.g. CHAPS)

Effects

Ionic (e.g. Triton X100, sodium deoxycholate)

Freezing and thawing

Chelating agents (e.g. EDTA)

Hypertonic or hypotonic solutions

Methods

Table 6.1 (continued)

Retention of biochemical components and

Widely appreciated for supportive along with enzyme method

Supportive along with chemical methods

requires extensive washing Collagen depletion enhances stiffness

tissue ECM architecture Oftentimes utilized to remove the remnant SDS Less harsh than SDS, thus less damaging to the structural integrity of the tissue ECM ultrastructure preservation

Heath (2019), Yao et al. (2019)

Gilpin and Yang (2017), Roth et al. (2017)

Snap freezing; ice crystals disrupt ECM microstructures Fail to meet the

Crapo et al. (2011), Keane et al. (2015), Ott et al. (2008), Simsa et al. (2018) Elder et al. (2009), Xu et al. (2007)

Boccafoschi et al. (2017), Meezan et al. (1975); Vavken et al. (2009); Woods and Gratzer (2005)

References

Prolonged exposure perturbs mechanical properties of the scaffolds

Causes ECM-condensation alone may not have efficiency

Incomplete cell removal

Disadvantages

Advantages

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Mechanical agitation

Sonication

Mechanical force (High hydrostatic pressure)

Cell lysis (mostly on the surfaces), efficiently removes all cellular and nuclear materials, maintained ECM Formation of micropores, better penetration of decellularization material and better cell attachment after decellularization Removal of cellular components, retained ECM and biomechanical properties Cell lysis, helps in chemical exposure and removal of cellular material Maintains tissue functionality

Eliminates antigenic cellular components Retention of the vital extracellular matrix

biomechanical properties Non-toxic

Can disrupt ECM Requires excessive washing as it effects efficiency

Problems in recellularization

requirements for immunogenicity May damage ECM Excessive washing

Choi et al. (2011), Wilson et al. (2016)

Azhim et al. (2011), Forouzesh et al. (2019), Yusof et al. (2019)

Fu et al. (2014)

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Fig. 6.1 Impact of three different decellularization procedures: on ECM-derived hydrogels obtained from porcine corneas. (a) Main steps in the fabrication of cornea ECM-derived hydrogels. (b) histological examination of hydrogels, stained with haematoxylin and eosin, picro-sirius red and Alcian blue; black scale bar ¼ 100 μm, white scale bar ¼ 50 μm. Figures adapted from (FernándezPérez and Ahearne 2019) # 2019 Springer Nature

factors due to rigorous decellularization process. However, they hold capacity to sequester the bioactive factors due to their inherent affinity towards them (Wilgus 2012). Therefore, upon implantation, DECM materials are able to restore the growth factor pool at the injury site and accordingly guide the resident cells towards tissue repair pathway (Hussey et al. 2018). Unlike the DECM materials, synthetic materials lack this capacity to sequester the bioactive factors and control cellular behaviour.

6.3.4

Decellularized Material Systems: Applications in Tissue Engineering

The formulation of DECM from different tissues/organs is sometimes found more complex even as after decellularization as it encompasses far more residual factors

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than cell-derived ECMs (Caralt et al. 2015). The biochemical structure and remnants of DECM will influence the quality and composition of the final acellular matrix. The DECMs are designed to generate 2D or 3D templates that recapitulate the native microarchitecture and mechanical properties of the target tissue or organ ECM (Noor et al. 2019; Zhang et al. 2009). However, these articulated structures would sometimes work incompetently due to many unavoidable reasons like (a) difference in tissue origin, (b) age, (c) method of decellularization, (d) remnants deposit on DECMs, (f) interference with recellularization and (g) immunogenicity. A plethora of reports have been listed in the recent years where different approaches of decellularization and in vitro recellularization have been targeted, their data assessments, substantial aspects and an optimal decellularization characterization is still somewhat unclear (Caralt et al. 2015). However, with accrescent newer studies, the research of optimum conditions and practicability of decellularization procedures is enduring. The functionality of recellularized 2D cellular sheets, 3D tissue engineered biografts (in vitro models) or organoids must be optimized, through which the efficiency of the regenerated archetypical can be assessed. The process of recellularization which assimilates appropriate cell types, their seeding pattern, a physiological relevant culture microenvironment, would largely depend on the complexity of the derived DECMs. These biomimetic medical materials can be used as multiple platforms, through which they could provide a suitable prototype to tissue engineering based constructions. Based on the requirement of renewal or replacement, the decellularization procedure with the cell sheets, tissue or organ of interest could be chosen to derive suitable DECMs. Cell sheets are generally used to develop the monolayer biografts where majorly a single cell type is grown on the derived DECM platform. These types of constructs will be useful in both the cases: minor injuries regeneration (like skin patches) and in complex structural section regeneration (by sheet stacks, like periodontium, cartilage, etc.) (Cramer and Badylak 2019). In an elegant study, authors have shown that by using a decellularized periodontal ligament cell sheet (by using NH4OH/Triton X and DNase perfusion solutions), the periodontium regeneration was achieved efficiently (Farag et al. 2014). In another study, authors have explored the mechanically supportive properties of various DECM and found that PLGA/PLA mesh scaffold coated with cell-derived extracellular matrix (type I collagen 293 T-DK cells) can provide human umbilical cord blood-derived mesenchymal stem cells, a better microenvironment for osteogenesis (Noh et al. 2016). Various other studies related to the vascularized tissue, cartilage, bone and kidney regeneration are also in support of these stack models and showed promising results previously (Gong et al. 2011; Noor et al. 2019; Zhang et al. 2015). Furthermore, for the fabrication of multicellular complex organs (e.g. heart, liver) the seeding of multiple cell types in their sheltered pattern is a prerequisite. Because of this complexity, the wide-ranging decellularization procedure or combinative approach would be employed, so that the ultrastructure, biomechanical properties of ECM of specific organ can be maintained. So far, a variety of studies have put forward many bioengineered DECM based organ models that

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have subsequently been tested for their main characteristics and functions, either in vitro or in vivo or both (Caralt et al. 2015; Goecke et al. 2018; Pan et al. 2016). The other applications through which these decellularized materials can be used to nurture cells in artificial environment are (a) as a thin biopolymer layer coating on plastic/glass substrate or on other polymer layers to mimic ECM critical characteristics (Sullivan et al. 2014); (b) in the form of particulates as a vehicle for small molecule delivery (Edgar et al. 2018); (c) in the form of DECM based hydrogels that can be useful either injectable or in situ based studies (Bai et al. 2019) and (d) in the form of decellularized bioink for 3D printing (Noor et al. 2019). Recently in 2018, Landry MJ and his team have extensively reviewed the properties, ultimate challenges and critical aspects of the usage of DECM as a coating material. Emphasis was given to biomaterial’s assets (like their hydrophilic and soft nature) that could be used for glass/plastic coating purposes (Landry et al. 2018). Moreover, a layer-by-layer placing of soft ECM materials (water-soluble polymers) could be the right choice for coating purposes as they have the longlasting stability and water retention properties similar to those of natural constituent. More recently, in a well-designed study, authors have followed the strategy of a combinative approach, where the temperature-sensitive rat heart decellularized hydrogels were used for growing brown adipose-derived stem cells for cardiomyogenic regeneration. Interestingly, enhanced cardiomyogenic differentiation and myocardial repair with maintained chamber geometry were noticed in vitro and in vivo (Bai et al. 2019). Also, in another study Sawkins et al. emphasized to opt for the superlative decellularization method (depending on the site-specific homologous tissues or heterologous tissues) and necessity of the characterization of DECMs prior to use (Sawkins et al. 2013). While comparing different hydrogel compositions, authors have demonstrated that ECM hydrogels have been found to have significant potential for clinical delivery over carrier liquids. Moreover, with the advent of 3D printing know-how in biomedical sciences, the anticipation towards the manufacturing of DECM-derived functional parts with decent strength is now possible. This field has broad dimensions and applications as it involves advanced ECM biomaterials, live cells, controlled motor systems and computer-aided designs to have exact configuration and specific control over the prototyped structures. In a recent study, perfusable and vascularized thick cardiac patches were 3D bioprinted by using a bioink made up of DECM biomaterial obtained from the omental tissue of patients (Noor et al. 2019). This technology proved to be highly beneficial because it could lead to the generation of personalized bioink. Not only this, the research team also demonstrated 3D bioprinted fully grown cellularized human heart containing native architecture of the heart. Clinical claims of DECM based biomaterials are rising, howbeit, a majority of products are for the tissues that exhibit less complexity. Many FDA approved or IND submitted DECM based products and therapies have pointed to tissue rejuvenation and replacement. There are many commercial decellularized products that have already paved the way to market. For instance, AlloDerm, Fortiva, Biodesign Hernia graft, DermaSpan, Avalus, GraftJacket, Permacol and Oasis Ultra are some of the commercially available DECM based bioscaffolds for various tissue engineering applications

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(Cramer and Badylak 2019). Additionally, clinical trials have been carried out for more complex microstructures including fat grafting surgeries (Wang et al. 2013). These results illustrate the potential in the expansion of more DECM based conducts for a variety of tissue engineering applications. In the context of decellularized materials, it is worth mentioning that the decellularized materials obtained from plant tissues have gained much appreciation in recent times. The structural platform provided by plant tissues has been smartly used to grow animal tissues on their decellularized pre-formed construct. The wellknown example of cardiac tissue grown on decellularized spinach leaves is worth mentioning here (Gershlak et al. 2017). The vessels-like structures of the plant leaves were utilized as a natural platform to grow human endothelial cells, which ultimately generated a pre-vascularized scaffold for engineering artificial tissues. Perfusionbased decellularization applied on plant tissues led to the development of a perfect acellular construct, which could be again recellularized with human cells. Colonization of human endothelial cells towards the inner surfaces of plant vasculature was successfully demonstrated along with stem cell-derived cardiomyocytes on the outer surfaces of plant scaffolds. The study not only revealed a cost-effective plant based biomaterial but also provided a simple technology to obtain pre-vascularized scaffolds for tissue engineering applications (Gershlak et al. 2017). This successful study further led to the exploration of various other plant based decellularized tissues that could be used as a scaffold. Various fruits and vegetables are a source of cellulose biomaterials, which could be used as 3D constructs upon successful decellularization steps (Fig. 6.2). Plant derived porous scaffolds from apple, carrot, broccoli and other vegetables showed growth of human stem cells and generation of bone-like tissues under in vitro and in vivo conditions (Lee et al. 2019b).

6.4

Naturally Derived Biomaterials

Naturally available materials extracted from various sources of animals, plants, insects or microorganisms have inherent remarkable functional properties towards tissue regeneration. The diversity among numerous living creatures further becomes the source of variation in natural materials. For example, collagen and gelatin extracted from bovine, porcine and fish sources slightly differ in terms of functional properties (Guo et al. 2018; Ninan et al. 2014). Another example of such diversity is silk protein. Silk is isolated from silkworm cocoons or directly from silk glands (Janani et al. 2019). The raw material of silk is easily available in the sericulture farms. Likewise, chitin is extracted from the exoskeletons of insects, which is further processed to produce chitosan biomaterial (Croisier and Jérôme 2013). Silk isolated from mulberry and non-mulberry categories differs in terms of their amino acid sequence and functional properties (Chouhan et al. 2017; Chouhan et al. 2018b). Furthermore, plants generate a range of compounds like polysaccharides, biologically active proteins, antioxidant materials, antimicrobial compounds and many more, which have been extensively used since decades (Jovic et al. 2019). For instance, curcumin extracted from turmeric has been explored for cancer treatment,

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Fig. 6.2 Plant based biomaterials demonstrating decellularized plant derived scaffolds obtained from various fruits and vegetables that act as cellulose-based constructs with microporous structures. Figures adapted from (Lee et al. 2019b) # 2019 Springer Nature

wound healing and drug delivery applications (Ahangari et al. 2019). The most commonly used natural biomaterials are broadly categorized into polysaccharides and proteins (Fig. 6.3), which have been described in details. Other materials include bioceramics and biominerals that have also been described in the subsequent section of this chapter.

6.4.1

Proteins Based Bio-Derived Biomaterials

Proteins are macromolecules that are made up amino acids. The large sized proteins form structural biopolymers and act as a biomaterial that supports the framework of any tissue. Various animal and insect origin protein biopolymers have been established as biomaterials for tissue engineering applications. For example, collagen, gelatin, fibrin, keratin, serum albumin proteins extracted from animal sources have been extensively explored for developing artificial tissues and organs (DeFrates et al. 2018). Insect origin protein biopolymer includes silk proteins that are extracted from silk cocoons or silkworm glands. Such structural proteins have repetitive amino

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Fig. 6.3 The schematic representation of bio-derived biomaterials showing wellknown protein and polysaccharide biomaterials that have been utilized for tissue engineering and various biomedical applications

acid sequences and hold highly ordered secondary structures like β-sheets, triple helix or coiled coil structures (Janani et al. 2019). The common features of such proteins include structural hierarchy and self-assembly properties upon physical or chemical stimulation. Such proteins are helpful in developing a scaffold or a structural framework that supports cellular growth, proliferation and migration to gradually create a mature living tissue (Chouhan et al. 2019a; Chouhan et al. 2019c). Biomaterials that are completely made up of proteins are often biodegradable and bioresorbable. The degraded parts of protein biopolymers are amino acids, which are easily resorbed by the host system (DeFrates et al. 2018). Bioresorbable nature of protein biopolymers is also advantageous for tissue engineering as the slow degradation process happens with the remodelling of neo-tissue (Carmagnola et al. 2018). This eventually helps in the maturation of the newly formed tissue and functional restoration of the organ. Therefore, protein-based biomaterials have been widely explored in wound healing, tissue engineering, drug delivery and biosensor applications (DeFrates et al. 2018). Nanotechnology has also been greatly benefited by the protein-based biomaterials as the biodegradable nanoparticles are easy to uptake without inducing systemic toxicity in the host (Mehrotra et al. 2019). They are also easy to conjugate with drugs and DNA molecules, making it possible to deliver drugs and genes (DeFrates et al. 2018). Following are the protein biopolymers well known for tissue engineering applications.

6.4.1.1 Collagen Collagen is a naturally derived structural protein present in the ECM of tissues and is also abundantly present in the body in various forms (Chevallay and Herbage 2000). It is majorly produced by the fibroblast cell type. There are several types of collagen found in the animal tissues like skin, tendons, bones and ligaments. Collagen is triple helical polypeptide and it is rich in glycine amino acids, which enable stable structure formation of collagen fibres (Hu et al. 2012). Being a major component

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of native tissue, collagen is the most preferred natural biomaterial for tissue engineering applications. Development of artificial tissues began with the utilization of collagen biomaterials (Burke et al. 1981). Fabrication of collagen based ‘Integra’ and ‘Apligraf’ as skin substitutes are decent examples of acellular and cellular grafts made up of collagen, respectively (Bhardwaj et al. 2017). Collagen has been exploited in different forms such as hydrogel, sponge, film, sheet, lattice, nanofibrous mat or gels for the development of advanced bioengineered constructs (Copes et al. 2019) (Brauer et al. 2019). Regardless of the inherent biocompatibility property of collagen, the collagen based lattices were found to be mechanically weak as they suffer from rapid degradation (Meyer 2019). Therefore, many strategies have been developed to make it mechanically strong such as chemical crosslinking, enzymatic crosslinking and blending with other biomaterials (Chevallay and Herbage 2000; Meyer 2019). Collagen contains inherent cell binding motifs like ArgGly-Asp (RGD) that helps in better adhesion of cells, which ultimately helps to populate the scaffold with cells under in vitro conditions (Hong and Stegemann 2008). In addition, collagen material also provides necessary cues to the cells for growth, migration, proliferation and differentiation. The concept of pre-vascularized artificial tissues was also easily implemented due to the inherent biological cues provided by collagen biomaterial (Marino et al. 2014). With the advent of bioprinting technology, collagen has been widely used for bioprinting of artificial organs such as heart tissue and full-thickness skin, thereby showing potential for organ reconstruction (Biazar et al. 2018; Lee et al. 2019a).

6.4.1.2 Gelatin Gelatin is denatured and hydrolysed form of collagen obtained by hydrolysing the collagen protein fibrils through physical and chemical methods (Echave et al. 2017). It is mostly obtained from pig skin and is relatively cost-effective compared to collagen due to its denatured form. The physical method of acquiring gelatin includes thermal treatment of collagen at 40  C, by which the hydrogen bonds within the collagen fibrils are broken (Nikkhah et al. 2016). The chemical method includes hydrolysis of collagen under acidic or alkaline conditions that cleave the covalent bonds. The hydrolytic process performed under acidic environment (pH 4) produces gelatin type A; whereas hydrolysis under alkaline conditions produces gelatin type B (Nikkhah et al. 2016). The thermo-responsive property of gelatin is due to the denaturation process of collagen that brings the reversible gelation properties in gelatin. It readily gels at low temperature and melts at higher temperature into liquid state. Gelatin has been used in various formats and blends for tissue engineering applications owing to its biodegradability, biocompatibility and simple processing properties (Echave et al. 2017). Commercially available gelatin based sponge Gelfoam has shown potential grafting applications for soft tissues like heart (Li et al. 1999). In addition, a diverse range of charged bioactive molecules can be easily loaded to gelatin vehicles due to its tunable isoelectric point, which allows gelatin to form polyion complexes by electrostatic interactions (Foox and Zilberman 2015). A variety of growth factors have been loaded to gelatin based drug delivery vehicles either as single or multiple growth factors programmed release ferries

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(Yamamoto et al. 2001). Gelatin microparticles have also been utilized as implantable cell carriers for tissue regeneration (Contessi Negrini et al. 2020). The thermoresponsive and reversible gelation properties of gelatin provide shear-thinning properties to it, which make this biopolymer quite useful for 3D bioprinting applications. Gelatin and its chemically modified forms have been used as bioadhesives, bioink and pre-formed scaffolds (Echave et al. 2017; Guo et al. 2018; Hsu et al. 2019). For instance, the methacrylated modified form of gelatin— gelatin methacrylamide (GelMA) has been extensively used for tissue engineering applications (Hsu et al. 2019).

6.4.1.3 Fibrin Fibrin glue and fibronectin are the natural proteins of blood clot that serve as a supporting matrix over open wounds (Clark 2001). Fibrin biomaterial is developed by combining fibrinogen and thrombin in calcium chloride (CaCl2) solution. The fibronectin protein contains cell attracting sites due to the presence of RGD tripeptide sequences in the amino acid sequence (Clark 2001; Currie et al. 2001). The whole concept of using fibrin as a biomaterial has been adapted from the natural blood clot that acts as a provisional matrix over wounds. The fibronectin fibres in the blood clot recruit cells towards the wound and sequester growth factor for healing purpose (Currie et al. 2001). Hence, artificial constructs containing fibronectin fibres or fibrin protein have been developed in various forms such as suspension, gel, sheet or membrane for faster tissue regeneration (Clark et al. 2007; Currie et al. 2001). The fibrin protein also attracts blood platelets and thus has been utilized as glue for promoting haemostasis (Currie et al. 2001). Human plasma is rich in fibrin and thus has been used to extract fibrin protein for the fabrication of scaffolds as autologous biomaterial (Llames et al. 2004). Fibrin has been hugely utilized for developing tissue engineered grafts, especially dermal grafts. Fibrin based commercially available skin constructs like BioSeed-S, AcuDress, Allox and Cyzact are currently under investigation for skin tissue engineering applications on a larger scale (Shevchenko et al. 2010). Combination of fibrin glue and fibronectin has demonstrated enhanced migration of fibroblast and keratinocytes leading to an accelerated wound healing (Currie et al. 2001). The endogenous fibrin clots have been proven to sequester vascular endothelial growth factor (VEGF) and thereby promote angiogenesis by slowly releasing the growth factor during the initial phase of wound healing (Clark 2001). This property of sequestering growth factors by binding them and protecting them from proteases has been exploited to develop fibrin based artificial matrices as a sustainable GF delivery vehicle for enhancing the tissue repair rate (Wong et al. 2003). In addition, autologous cell delivery system using fibrin glue and cell–fibrin complex have been developed to deliver autologous mesenchymal stem cells directly at the injured site (Wu et al. 2012). 6.4.1.4 Silk Silk-based tissue engineered grafts have been increasingly investigated since last two decades due to its properties like biocompatibility, tunable degradability,

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tunable tensile properties, the potential of stimulating fibroblasts for ECM secretion and tissue repair (Holland et al. 2019). Silk cocoons contain two kinds of proteins— silk fibroin (SF) which is the fibrous protein component of cocoon and sericin protein which is the glue like protein responsible for sticking the silk fibres with each other (Altman et al. 2003). The well-known example of silk protein is SF isolated from Bombyx mori silk cocoons. The SF protein from B. mori silk fibres contains repeats of glycine-alanine [GAGAGS]n, which represent dominant β-sheet structures of the protein biopolymer (Fig. 6.4a) (Chouhan et al. 2019d; Janani et al. 2019). Both the silk fibroin and sericin are widely explored for the development of artificial tissue constructs. Various formats like scaffolds, nanofibrous mats, films, hydrogels and fibres can be easily generated from silk proteins (Chouhan et al. 2018a; Gilotra et al. 2018; Janani et al. 2019). Silk sericin (SS) has been considered as a biomaterial since decades, and it has wide applications in the fields of cosmetics, drug release, engineering of artificial tissues and wound healing (Fig. 6.4b) (Lamboni et al. 2015; Mehrotra et al. 2019). Being a biocompatible and non-immunogenic material, regenerated silk-based scaffolds have recently been approved with 510 (k) clearance by the U. S. A. Food and Drug Administration (FDA) for biomedical applications (Sofregen 2019). It is the first time that the scaffolds made up of a solubilized form of SF are approved for commercialization in the healthcare market. Such progress made by silk in the field of tissue regeneration demonstrates the potential of this biomaterial and silk-based products in the healthcare market in the near future. Apart from biocompatible nature of silk, the crystalline structures of this protein biopolymer provide high mechanical strength and structural properties that make this natural material unique (Bhunia and Mandal 2019). Isolation of SF from silk cocoons is performed by degumming the cocoon raw material and subsequently dissolving the fibres in ionic or organic solvents. The degumming process also yields in glue like sericin component of the silk cocoons, which can be processed separately to obtain pure form of sericin protein (Rockwood et al. 2011). The isolation procedures and protein sequences differ among various silkworm varieties. Therefore, isolation of SF from the silk glands of fifth instar silkworms is often performed on non-mulberry silk varieties (Nileback et al. 2017). Silk is broadly categorized under mulberry (domestic) and non-mulberry (wild) varieties. The non-mulberry silk fibroin (NMSF) protein is usually obtained directly from the silk glands because the fibres do not get easily dissolved in ionic liquids (Chouhan et al. 2017). NMSF holds unique properties in cell binding efficiency and cell culture. Their peptide sequence contains RGD tripeptide motifs, which is a wellexplored cell binding motif that gives an added advantage of this silk variety in tissue engineering applications (Gupta et al. 2015). Silk-based matrices have been studied for tissue engineering applications in the form of various constructs and formulations (Holland et al. 2019; Jewell et al. 2015). SF is considered an outstanding material for developing load bearing hard tissues like bone, meniscus and intervertebral disc when used in high concentration (Bhunia et al. 2018; Moses et al. 2018; Yao et al. 2016). In addition, SF also provides optimum mechanical properties and platform for mimicking the soft tissues like skin, pancreas, liver and cardiovascular grafts by

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Fig. 6.4 Schematic representation of (a) structure of silk cocoon fibre showing fibroin fibres wrapped with sericin coating. The fibroin strands are made up of numerous fibrils with silk-I and silk-II conformations and β-sheet structures. Figure adapted from (Janani et al. 2019) # 2019, American Chemical Society. (b) The sericin glue protein extracted after degumming of cocoon fibres holds various applications as shown in the diagram. Figure adapted from (Mehrotra et al. 2019) # 2019, American Chemical Society

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using low concentration protein (Chouhan et al. 2018b; Janani et al. 2018; Kumar et al. 2018; Mehrotra et al. 2017). Therefore, functional cellular tissues ranging from soft tissues to hard tissues have been successfully fabricated by taking the appropriate material format and protein concentration.

6.4.1.5 Keratin The keratin protein is broadly defined as filament-forming proteins that are found in corneous tissues like hair, horns, claws, nails or hooves (Bragulla and Homberger 2009). It is insoluble structural protein and is broadly categorized as hard and soft keratin proteins. They have dominant α-helical structures and cysteine rich non-helical domains that form disulphide bonds (Mogosanu et al. 2014). The cysteine rich domains are responsible for their toughness and durable structures. Extraction of keratin protein is performed by disrupting the disulphide bonds through oxidation process (Zhu et al. 2017). Keratin protein naturally holds cell binding property due to the presence of cell adhesion motifs like RGD and Leu-AspVal (LDV), thus making it a suitable biomaterial (Srinivasan et al. 2010). In addition, keratin is a biocompatible and biodegradable material, thereby showing potential in tissue engineering applications (Srinivasan et al. 2010). Various types of matrices have been successfully developed using keratin protein such as scaffold, hydrogel, thin film and nanofibrous mats that are utilized for tissue engineering applications (Bhardwaj et al. 2015; Srinivasan et al. 2010). In a study, haemostatic property of keratin-based hydrogel was well demonstrated under a liver injury model, thereby showing the applications of keratin for bioadhesives and acute trauma cases (Burnett et al. 2013). In another study, human hair keratin was modified by alkylation process to fabricate tunable hydrogels that supported the growth and proliferation of pre-osteoblasts. The hydrogel thus fabricated was further tuned to deliver multiple drug molecules including growth factors and antibiotics for possible drug delivery and tissue engineering applications (Han et al. 2015). The processed keratin, although soluble in nature, lacks high mechanical properties due to the rigorous isolation process. Therefore, it is often used in the composite form by adding a suitable biomaterial along with it. Blending of other biomaterials has been proven to improve the physico-chemical properties of the scaffolds (Bhardwaj et al. 2015). The composite scaffolds of silk-keratin blends demonstrated ideal physicochemical properties for soft tissue engineering applications with microporous structures and suitable mechanical strength (Bhardwaj et al. 2015). Similarly, composite porous scaffolds of keratin with gelatin and chitosan supported culture of fibroblasts and showed evidences of these constructs for tissue engineering applications (Balaji et al. 2012). Although keratin has been shown applications in soft tissue engineering especially for skin and wound healing, it has not been much explored for other broad area of research in tissue engineering, suggesting wide scope of this biomaterial for further applications in future.

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Polysaccharides Based Bio-Derived Biomaterials

Polysaccharides are natural biopolymers made up of simple sugars that hold essential properties to support tissue regeneration and have been used as biomaterials (Tiwari et al. 2018). Some of the polysaccharides hold exceptional structural properties and act as a pre-formed matrix. For instance, alginates can be directly used for wound management, haemostasis and tissue engineering applications (Aderibigbe and Buyana 2018). Polysaccharides can be easily derived from numerous natural sources like algae, animals, plants and microorganisms. For example, alginate biomaterial from algae, glycosaminoglycans (hyaluronans, chondroitin sulphate and heparin) from animals and pectin, gums derived from plants (Huang and Fu 2010). The well-known example of plant derived polysaccharide material is cotton, which has been used as a wound dressing material since ages (Sood et al. 2013). Insect derived chitin and chitosan are also great examples of bio-derived polysaccharide biomaterials that have been extensively applied for tissue engineering and regenerative medicine (Croisier and Jérôme 2013). Polysaccharides are easy to modify by simple manipulation of the intermolecular associations and chain conformation that lead to changes in their physico-chemical properties (Shelke et al. 2014). The intra/interchain hydrogen (H-) bonding is easy to reform in the polysaccharide materials due to the presence of hydroxyl groups. This enables insolubility in the materials upon drying and matrix fabrication under controlled conditions (Wasupalli and Verma 2018). In addition to free hydroxyl groups, some polysaccharides also hold amino or carboxylic groups that offer other chemical ways to alter their structure and physico-chemical properties (Shelke et al. 2014). Taking advantage of the chemically modified polysaccharide biomaterials, numerous types of porous scaffolds, foams, matrices, nanofibres, hydrogels, microgels and nanoparticles have been successfully developed. These materials have greatly benefitted the field of tissue engineering and regenerative medicine by developing wound dressing materials, artificial grafts, and drug delivery vehicles. Following are the well-known polysaccharide biomaterials that have been widely used for biomedical applications.

6.4.2.1 Glycosaminoglycans Glycosaminoglycans (GAGs) are components of the ECM of various tissues. There are three types of GAGs extensively utilized for tissue engineering applications— hyaluronic acid (HA), heparin and chondroitin sulphate (CS) (Rnjak-Kovacina et al. 2018). Hyaluronic acid is the most explored material among all GAGs because of its abundance in the ECM of skin, cartilage and other tissues (Price et al. 2007). It is highly hydrophilic material due to the presence of negatively charged polymer chains. It is a unique type of GAGs because it does not contain sulphate groups. It is composed of alternating α-1,4-D-glucuronic acid and β-1,3-N-acetyl-D-glucosamine units bonded by β(1 ! 3) linkages (Price et al. 2007). The HA contains both carboxylic and hydroxyl groups that allow chemical alterations of this material and also modulate the mechanical and degradation properties while retaining the bioactivity. For instance, enzymatic oxidative coupling was applied to modify the HA and

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develop HA-tyramine hydrogels (Kurisawa et al. 2005). The hydrogels thus generated, demonstrated, varied degradation patterns in presence and absence of hydrogen peroxide. Being a GAG, HA has high water absorption and water retention capacity that helps in cellular activities such as cell migration and proliferation. Being water soluble, it is usually crosslinked chemically or enzymatically to develop scaffolds or hydrogels (Chircov et al. 2018). Another advantage of this biomaterial is that it degrades into simple sugars that can be resorbed by the host system. HA based regenerative therapies have been mostly applied to treat osteoarthritis of knee joints and wound healing (Neuman et al. 2015; Wigren et al. 1978). Some of the commercially available HA based skin grafts are Hyalomatrix, Hyaff (Fidia Advanced Biopolymers) and Hyalograft 3D—dermal graft (Price et al. 2007; Shevchenko et al. 2010). Similar to HA, chondroitin and heparin have also been used for tissue engineering and regenerative therapies. Chondroitin sulphate is a sulphated and negatively charged glycosaminoglycan. It is composed of β-glucuronic acid and N-acetyl galactosamine molecules and acts as a connective tissue (Rnjak-Kovacina et al. 2018). It is majorly present in the fibrous connective tissue of articular cartilage and is produced by chondrocytes. It is mostly used for cartilage tissue engineering and cartilage repair (Henrotin et al. 2010). In a study, CS was ionically conjugated with transforming growth factor-beta3 (TGF-β3) for cartilage repair applications (Park et al. 2010). The scaffolds demonstrated culture of MSCs and chondrogenesis, giving clues towards cartilage tissue development. The scaffolds also showed longterm release of TGF-b3. Apart from cartilage regeneration, CS has also shown promising wound healing properties (Im et al. 2013). Another GAG, namely, heparin is also negatively charged and highly sulphated. It is composed of 2-Osulphated iduronic acid and 6-O-sulphated N-sulphated glucosamine (Liang and Kiick 2014). Heparin acts as a carrier of growth factors and positively charged molecules, thereby has been utilized for drug delivery applications (Liang and Kiick 2014). It is also a well-known anticoagulant as it supresses the thrombin formation. Therefore, heparin has been extensively used in the vascular tissue engineering applications. Heparin coated vascular grafts and stents are efficient in preventing the formation of thrombotic emboli (Zamani et al. 2017).

6.4.2.2 Alginates Alginates come in the category of linear unbranched polysaccharides. They are composed of 1,40 -linked β-D-mannuronic and α-L-guluronic acid (Lee and Mooney 2012). They are mostly derived from the brown seaweeds. Depending upon the source of alginate, there are variations in the proportions of the monomers. The gelation of the polymer depends on the ion binding (Shelke et al. 2014). Alginates hold high water absorption capacity and have been highly explored for wound healing and tissue engineering applications. They interact with Ca2+ to generate hydrogels that act as a well-formed matrix. Numerous alginate-based materials are commercially used as wound dressings, such as Nu-Derm (Johnson & Johnson) and AlgiSite (Smith & Nephew) (Muntimadugu et al. 2013). Alginates are also suitable for gene delivery applications as it readily forms nanoparticles with calcium

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carbonate that encapsulate DNA molecules (Muntimadugu et al. 2013). Range of alginate-based materials can be developed by modifying the molecular weight, block length and composition (George and Abraham 2006). Various formulations have easily been developed by slightly altering the alginate materials or by ionic crosslinkers (Iwamoto et al. 2005). Alginate-based hydrogels, sponges, fibrous matrices and injectable systems have shown potential in tissue engineering and drug delivery applications.

6.4.2.3 Agarose Agarose is a polysaccharide material consisting of galactose backbone (Zarrintaj et al. 2018). It is mostly obtained from red algae and seaweeds. In the non-sulphated fraction of agar, agarose, the monomers wrap together tightly to embrace a double helix capable of trapping water molecule inside its helical structure. Agarose is mostly used in gel form and can be easily tuned in terms of stiffness and mechanical properties (Zarrintaj et al. 2018). It also holds thermoelastic properties that allows temperature-controlled scaffold synthesis and gelation ability (Andersen et al. 2015). Agarose based matrices have been explored for the fabrication of insulin secreting islet cells encapsulated microgels targeting artificial pancreas development (Iwata et al. 1992). Agarose based matrices have shown promising results in culturing and delivery of stem cells, cardiomyocytes and chondrocytes (Mak et al. 2015). In a study, agarose microcapsules were developed to differentiate embryonic stem cells into dopaminergic neurons (Ando et al. 2007). The aim of the study was to develop a therapeutic solution for Parkinson’s disease by developing an artificial model of in vitro cultures neuronal tissue. 6.4.2.4 Carrageenan Carrageenan (CRG) polysaccharides are flexible linear polymers that form double helical structures at higher concentrations (Yegappan et al. 2018). CRGs are isolated from marine organisms like seaweeds. They hold various biomedical applications because it comes under sulphated polysaccharides and it mimics the GAGs. They are composed of D-galactose residues and 3,6-anhydro-galactose with ester sulphates (Campo et al. 2009). Depending on the extraction source, sulphate content and solubility, there are six forms of CRG, namely Kappa (κ)-, Iota (ι)-, Lambda (λ)-, Mu (μ)-, Nu (ν)- and Theta (θ)-CRG. Among all, κ, ι and λ – CRGs are widely used owing to their viscoelastic and gelling properties (Cunha and Grenha 2016). CRGs are also known to form hydrogels by interacting with K+ and Ca2+ ions (Chronakis et al. 2000). Owing to the thermoreversible properties of CRGs, they are used as an additive in food industry, gelling and stabilizing agent in engineering scaffolds. CGRs have essential biomaterial properties like biocompatibility, biodegradability and suitable mechanical strength (Wasupalli and Verma 2018). They form versatile gels which display thermo- (in presence of respective counterions) and stressresponsive (they thin under shear stress and recover their viscosity once the stress is removed) (Mano et al. 2007). CRGs based constructs have shown success in tissue engineering and drug delivery applications. Shear tinning nanoengineered gels were developed using k-CRG for bioprinting applications (Wilson et al. 2017). CRGs

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have also been used in combination with other biomaterials such as silk fibroin for bone tissue engineering (Nourmohammadi et al. 2017). The studies suggest potential of CRGs based constructs for various hard tissue engineering applications.

6.4.2.5 Chitosan Chitosan is an amino polysaccharide derived by deacetylation of chitin obtained from the exoskeletons of insects and shells of crustaceans (Croisier and Jérôme 2013). The native chitin material is not soluble in most of the solvents and thus it is further processed into chitosan material for easy handling material properties. They are composed of 2-amino-2-deoxy-Dglucopyranose and 2-acetamido-2-deoxyDglucopyranose units bonded by β(1 ! 4) glycosidic linkages (Croisier and Jérôme 2013; Shelke et al. 2014). Chitin is thoroughly processed through the steps of demineralization, deproteination and deacetylation to yield chitosan biomaterial. Chitosan has been utilized for numerous biomedical applications, such as wound healing, drug delivery and tissue engineering owing to its low toxicity, biocompatibility and biodegradability (Croisier and Jérôme 2013; Madihally and Matthew 1999). It also shares structural similarity with natural glycosaminoglycans, and thus is a preferred biomaterial for tissue engineering applications. It is easily moulded into a variety of formats with varied biological properties by simply changing the degree of acetylation (Liu et al. 2004). The free amino group of chitosan enables the material for chemical modifications and crosslinking properties. The easy processing and fabrication methods have led to the development of chitosan-based sponges, hydrogels, films, and nanofibres (Liu et al. 2004; Shi et al. 2006). Chitosan-based matrices have been widely explored for the development of wound dressings as it is a good haemostatic agent, and it also has bacteriostatic and fungistatic activities; thereby leading to enhanced wound healing rate (Carvalho and Mansur 2017; Mizuno et al. 2003). Being positively charged, it also facilitates easy incorporation of fibroblast growth factor and stimulation of ECM synthesis by triggering the proliferation rate of fibroblasts (Mizuno et al. 2003). A range of composite matrices have been developed by blending chitosan with other natural and synthetic polymers owing to the easy crosslinking efficiency of chitosan. For instance, GelMA was mixed with chitosan oligomer methacrylate (ChitoMA) to generate adaptable mesoporous hydrogel that helped in the development of a nerve tube (Hsu et al. 2019). Blended matrices with improved biostability and mechanical properties provide advantage for tissue engineering applications.

6.4.3

Other Bio-Derived Biomaterials

Other bio-derived biomaterials include materials obtained from bioceramics, corals and shells. The bioactive ceramic materials such as bioglass, hydroxyapatite, alumina dental implant and calcium phosphate Ca(PO4)2 materials have been highly explored for tissue engineering of load bearing tissues such as bone (Barua et al. 2018). Hydroxyapatite (HAp) is a crystalline compound naturally present in the bones, which is rich in calcium and phosphate ions (Zhou and Lee 2011). HAp has been used for bone tissue engineering since decades owing to its biocompatibility,

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non-immunogenicity and osteoinduction stimulating properties (Yoshikawa and Myoui 2005). It has also been used to coat bone replacements, fixation of prosthetic devices and as a filler in polymer matrices. Marine coral is another source of bio-derived biomaterial. The biocoral contains calcium carbonate mineral with other ionic compounds consisting of ions of strontium and magnesium (Giuliani and Manescu 2014). The biocorals are also mainly used for bone tissue engineering owing to its toughness properties and porous nature. Biocorals are also a source of producing hydroxyapatite by hydrothermal techniques (Balázsi et al. 2007). Biocorals hold osteoconductive and biocompatible properties that are suitable for artificially developing bone grafts (Guillemin et al. 1987). Similarly, sea shells, being composed of calcite microcrystals are also utilized for hard tissue engineering applications (Awang Junaidi et al. 2007). Plant derived materials like pullulan, dextran, pectins and gums have also shown potential applications related to tissue regeneration therapies (Jovic et al. 2019). In this context, it is worth mentioning that pectins and gums are hydrocolloid materials that are often used as moist wound dressings or gels for wound healing and burn management (Schoukens 2009). Mucilage or mucopolysaccharide biomaterials also come in the category of bio-derived polysaccharide biomaterials that are mostly used as bioadhesives or tissue sealants to treat trauma injuries (George and Suchithra 2019). Mucilage materials are composed of mannose and galactose sugar derivatives. Similarly, there are various bio-derived products extracted from plants, insects and animals that are being explored for regenerative therapies such as honey, bees wax and eggshells. Pullulan is a water-soluble material containing monomers with three glucose sugars. Although it is derived from yeast or fungi, it is non-toxic and non-immunogenic, thereby a potential biomaterial for biomedical applications (Singh et al. 2016). Being a superabsorbent, it has been greatly explored for wound healing purpose. Cationized pullulan has shown potential ability to act as a carrier of pDNA for gene delivery applications in living cells (Jo et al. 2010). Similarly, dextrans are also widely used for wound repair and regeneration. Dextran based artificial skin grafts have shown great success in healing full-thickness burn wounds (Shen et al. 2015). The field of materials research is gaining popularity owing to the diversity of natural resources. Discovery of new bio-derived materials and studying their biological properties provide a wide scope in the field of tissue engineering and regenerative medicine.

6.5

Microbial Derived Biopolymers

Microbial derived biomaterials have received unprecedented attention in the field of biomedical engineering owing to their water-soluble properties, negligible toxicity, controlled drug release and extended circulation time (Mokhtarzadeh et al. 2016). Microorganisms stand as convenient synthesis factory units for simple and low-cost fabrication of functionalized organic/inorganic biocompatible materials (RodríguezCarmona and Villaverde 2010). These biopolymers are produced by microorganisms during fermentation. Some of the leading bacterial-derived polymers and their prominent functions in tissue engineering are shown in Fig. 6.5. Dextran was the

Fig. 6.5 Representative model biopolymers derived from microorganisms, the chemical structure, and their main functions. Figure reproduced with permission from (Mokhtarzadeh et al. 2016) 2016 Elsevier B.V

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first biopolymer discovered by Louis Pasteur as a by-product of wine fermentation during the mid-nineteenth century and Van Tieghem discovered Leuconostoc mesenteroides to be the microbial cell factory (Rehm 2010). In 1886, bacterial cellulose came into focus following intracellular polymers such as cyanophycin from cyanobacteria. After these initial discoveries, other medically and industrially relevant polymers including polyhydroxybutyrate, alginate, xanthan, etc. were unearthed gradually (Rehm 2010). This section expounds on the numerous microbial derived biopolymers and biomaterials in tissue engineering and their various types. General methods of biopolymer synthesis have been explained with the example of a few representative model biopolymers based on their extensive available information and application potential. Further, the relevant biocompatible properties, as well as applications of the bacterial-derived biomaterials pertaining to tissue engineering and regenerative medicine have been elucidated.

6.5.1

Types of Bacterial Polymers

The degradation of polymers may follow an active route of enzymatic catalysis or a passive route of decomposition by hydrolysis (Katti et al. 2002). As compared to enzymatically decomposed polymers, hydrolytically degradable polymers are considered more applicable for in vivo applications due to their minimal patient-topatient and site-to-site variation. The hydrolytically labile chemical bonds in polymers include amides, esters, ureas, orthoesters, carbonates or anhydrides (Nair and Laurencin 2007). Based on the chemical structure and composition, bacterial polymers can be clustered as (1) polyamides [such as poly(γ-glutamic acid) (γ-PGA), poly(ε-L-lysine) (ε-PL) and multi-L-arginyl-poly(L-aspartic acid)/ cyanophycin granule polypeptide (CGP)], (2) polyesters (such as polyhydroxyalkanoates), and (3) polysaccharides (including cellulose, dextran, alginate, hyaluronic acid, gellan gum, xanthan, curdlan, colonic acid, glycogen and K30 antigen) (Rehm 2010). Polyhydroxyalkanoates (PHA) can be further categorized into three groups based on the chain length of carbon units in the monomers: (1) short chain length (SCL) PHAs with 3–5 chain length, e.g. polyhydroxybutyrate (PHB), 3-hydroxybutyrateco-3-hydroxyvalerate (PHBV), 4-hydroxybutyrate (P4HB), (2) medium chain length (MCL) with 6–14 carbon units (e.g. poly(3-hydroxyoctanoate), P(3HO), poly (3-hydroxyhexanoate), P(3HHx), poly(3-hydroxydodecanoate), P(3HDD), and poly(3-hydroxydecanoate), P(3HD), and (3) long chain length PHAs (LCL), with greater than 14 carbon unit chain length (Lu et al. 2009; Nigmatullin et al. 2015). Microbes prefer to exist in matrix-enclosed microcolonies with protective niche and homeostatic environment termed as biofilms (Stoodley et al. 2002). They may exist in mixed bacterial species or in symbiosis (e.g. cyanobacteria) (Sarma et al. 2016). Biofilm is composed of extra polymeric substances and detritus, with the major component being the water. Exopolysaccharides act as the cement to this biofilm matrix (Sutherland 2001). These polysaccharides can be further divided into exopolysaccharides (extracellularly secreted to the growth media, for example,

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cellulose, dextran, alginate, hyaluronic acid, xanthan, etc.), capsular polysaccharides (remain attached to the cell after secretion, K30 antigen), intracellular polysaccharides (glycogen) or cell wall polysaccharides (Rehm 2010).

6.5.2

Biosynthesis and Purification of Bacterial-Derived Polymers

6.5.2.1 Polyamides Poly-γ-glutamic acid (γ-PGA) is an anionic polyamide synthesized by microbes as an extracellular/capsular viscous material. In 1937, it was first identified in the capsules of Bacillus anthracis (Nair and Laurencin 2007). It was demonstrated that the gram negative Bacillus species was capable of producing γ-PGA in the culture medium under denitrifying conditions (Cheng et al. 1989). PGA is composed of D- and L-glutamic acid units connected by amide linkages between γ-carboxylic acid and α-amino units (Shih and Van 2001). Since then, several bacterial species including Staphylococcus epidermidis, Bacillus halodurans, Bacillus megaterium, Bacillus amyloliquefaciens, etc., have been demonstrated to produce γ-PGA (Rodríguez-Carmona and Villaverde 2010). The downstream processing of γ-PGA is comparatively uncomplicated as it is secreted outside the cell by the Bacillus sp. It usually involves three basic processes including centrifugation/filtration (with 0.45 μm filter unit) to remove the biomass, precipitation with the help of methanol/ethanol/propanol/hydrochloric acid or metal ions. This is followed by dialysis of the product to eliminate low molecular weight impurities (Buescher and Margaritis 2007; Pérez-Camero et al. 1999).

6.5.2.2 Polyesters Polyhydroxyalkanoates (PHAs) belong to the family of biopolyesters consisting of hydroxyalkanoic acids as monomers. PHAs are secreted by microorganisms inside their cell cytoplasm under stress conditions (starvation of oxygen, phosphorous, nitrogen, sulphur, etc.) but in existence of carbon (carbohydrates, alkanes, fatty acids, organic acids, etc.) surplus (Sudesh et al. 2000). The PHAs are stored as energy reserve nutrients inside the cells which can be degraded by depolymerases to serve as a carbon source. Poly(3-hydroxybutyric acid), (PHB) was the first PHA identified in Bacillus megaterium in 1925 by Maurice Lemoigne and since then, it has been widely studied (Winnacker 2019). The PHAs are strongly associated with the biomass unlike the bacterial polysaccharides that are extracellular. The additional steps of release of the PHA granules preceding cell isolation make the product recovery quite challenging. PHA is recovered by employing different techniques such as enzymatic cell disruption, solvent extraction (using chloroform), mechanical/chemical methods, dissolved-air floatation, etc. (Winnacker 2019). The commercial debut of PHAs began in 1980s with microbial fermentation in similitude to industrial commercialization of antibiotics (Chanprateep 2010; Chen 2009).

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6.5.2.3 Polysaccharides The biosynthesis of the polymer precursors involves specific enzymes. These enzymes have also been the key targets for metabolic engineering to obtain tailored polysaccharides with unique properties (Ruffing and Chen 2006). Nucleoside diphosphate sugars (acids) (ADP-glucose, UDP-N-acetyl-glucosamine and GDP-mannuronic acid) are the precursors for direct biosynthesis of bacterial polysaccharides. Exopolysaccharides like dextran have low immunogenicity, and defined molecular mass fractions on acid hydrolysis, two very desirable traits for clinical applications. Hyaluronic acid is a straight chain polysaccharide with a repeating unit of disaccharide, glucuronic acid and N-acetyl-glucosamine. It is found as an important component of extracellular matrix and can be produced by Streptococcus species through fermentation. Aside seaweeds, bacteria such as Azotobacter vinelandii and Pseudomonas fluorescens are natural producers of commercial alginate from glucose. Bacterial cellulose (BC) is an unbranched linear polysaccharide with β-1.4glucoyranose monomeric units. BC was identified by Brown in 1880s, during vinegar fermentation (Cacicedo et al. 2016). The biomedical applications of BC were first reported in a series of patents between 1986 and 1990 (Stumpf et al. 2018). Although BC has the same chemical formula (C6H12O5)n, as natural cellulose from plants, it differs highly in its physico-chemical properties (Eslahi et al. 2020). But the cellulose product from the latter is preferred due to its purity and absence of unwanted components such as hemicellulose, pectin and lignin (Rahman and Netravali 2016). Extraction and purification processes of BC are cheaper, simpler and present less burden on environment, as compared to that obtained from the plant source (Cacicedo et al. 2016; Stumpf et al. 2018). The commercially high yielding BC, Acetobacter xylinum during carbohydrate metabolism, secretes cellulose into the culture medium with the help of the membrane protein complex, cellulose synthase (Yoshinaga et al. 1997). The presence of hydrogen bonds in BC allows facile interactions with additional polymers of interest and tuneable intricate shapes. The absence of inherent antibacterial property initiates scope to develop BC composites (Cacicedo et al. 2016; Stumpf et al. 2018). The BC can be tailored to function both as a scaffold and reinforcement agent by incorporating certain materials (nanoparticle, peptide, or a polymer) to fabricate composite BC. This can be done following either in situ or ex situ techniques. In the former, the supplementary material (additives, alternate carbon source, etc.) is added into the culture medium during the BC synthesis process while in the later (post synthesis modification), the agent can be incorporated into the fibrous BC matrix after synthesis (generally after purification of BC) by physical (absorption) or chemical (crosslinking/co-polymerization) engagement (Eslahi et al. 2020; Stumpf et al. 2018).

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Microbial Derived Biopolymers for Tissue Engineering

6.5.3.1 Poly-g-Glutamic Acid (g-PGA) Hydrogels find its application in tissue engineering aspects as well as several clinical and pharmaceuticals applications (Buescher and Margaritis 2007). γ-PGA has been used to prepare several three-dimensional hydrogels to be applied as scaffolds in tissue engineering. γ-PGA solution can be converted to hydrogel by simple technique of irradiation. Swelling and specific water content in the biodegradable hydrogel could be controlled with the irradiation time (Choi and Kunioka 1995; Kunioka 2004). The technique of suturing is common for tissue adhesion and wound closure (Shih and Van 2001). γ-PGA crosslinked to gelatin formed a hydrogel in presence of water-soluble carboiimide. The hydrogel exhibited superior performance in terms of adhesion and homeostatic capabilities as compared to the commercial fibrin glue with insignificant inflammatory response (Otani et al. 1996; Otani et al. 1998). PGA combined with chitosan or gelatin has proved to be a better hydrogel bioadhesive glue as compared to fibrin glue (Hsieh et al. 2005; Otani et al. 1999). Growth factors, proteins, etc., are important to support the competency and proliferation of tissue cells (Hsieh et al. 2006). Combination of chitosan with γ-PGA further enhanced the cytocompatibility, release kinetics of recombinant human bone morphogenetic protein 2 (rhBMP-2) and mechanical strength of the hydrogel scaffold (Hsieh et al. 2006). Sulphonated γ-PGA behaved as an anticoagulant similar to heparin (Matsusaki et al. 2002). At 72% of carboxyl groups sulphonation, γ-PGA supported interaction with fibroblast growth factor (FGF-2) and was protected from thermal and acidic inactivation. The seeded fibroblast cells grew well on the stable scaffold (Matsusaki et al. 2005). Protein adsorption properties, mechanical strength and hydrophilicity of the biopolymer could be attuned by addition of varying ratio of poly(acrylamide) to γ-PGA. The swelling property changed in response to external pH and temperature (Rodríguez et al. 2006). It has also been demonstrated that the swelling of the γ-PGA hydrogels could be altered by variation in pH and temperature. The pH decreased (99%) through hydrogen bonding. This makes BC biomaterials thermally stable even at 100  C, during sterilization by autoclaving. Sterilization without loss of biophysical properties is a very essential factor to be considered for biomaterials in biomedicine and clinical studies (Cacicedo et al. 2016). Highly hydrophilic and hierarchical structure of BC films endowed with high porosity promotes cell migration and facilitates cutaneous healing. Preservation of a humid environment, and re-epithelialization with scope of self-degeneration and healing are the important criteria for wound dressings (Portela et al. 2019). Inadequate healing of wounds might lead to rise of chronic wounds. Hence appropriate wound dressing is necessary to diminish pain and exudate aggregation, as well as enhance ECM production and proteolytic activities (Khalid et al. 2017). Bacterially produced BC has been shown to exhibit tuneable water release rate and water holding capacity. The hydrophilicity of BC has been attributed to the high surface area of the pellicles in the BC fibre ribbons that lock in the moisture (Portela et al. 2019). BC derived bandages have been shown to stimulate healing of chronic wounds and burns in their sterilized, protected and moist environment, better than the traditional dressings (Khalid et al. 2017). The hydrophilicity and porosity of BC can be tailored with addition of chitosan, aloe vera gel, alginate, etc. Chemical modification of BC for adsorption of proteins such as haemoglobin, lysozyme, etc. has been also studied for interaction with the tissue of interest. The BC scaffolds can also be tailored with nanoparticles, antibiotics or antimicrobial peptides to imbibe antibacterial properties. Several interesting composite BC biomaterials have been discussed in the review by Portela et al. pertaining to intended applications (Portela et al. 2019). BC sheets augmented with titanium dioxide (TiO2) nanoparticles exhibited better and accelerated healing of burn wounds in mice models as compared to pure BC bandages. Recovery at the burnt site was evident with infiltration of fibroblasts, epithelial cells and small blood vessels. The nanocomposite bandages had antibacterial activity against Staphylococcus aureus (S. aureus) and Escherichia coli (E. coli) owing to the oxidative stress imparted by metal nanoparticles (Khalid et al. 2017). Another prerequisite for tissue engineering biomaterials often overlooked is the in vivo degradability. Although BC has a wide spectrum of applications in tissue regeneration and wound patches, the human body cannot degrade it. Hence, modifications of such biopolymers to impart biodegradability are necessary to eliminate surgical procedures for removal of such scaffold.

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Conclusion and Future Directions

In the quest of addressing the demand for artificial tissues, biomedical engineers rely heavily on employing the bio-derived materials. The use of decellularized ECM derived biopolymers as a biocompatible, immune proof and highly biomimetic scaffolding biomaterial is ever expanding. Nevertheless, the individual components extracted from ECM such as collagen, gelatin, fibrin or HA have shown great potential in tissue engineering applications. These biomaterials hold suitable physico-chemical and biological properties that has led to their immense usage in both pure and composite forms. Nature inspired biomaterials obtained from insects’ cocoons or exoskeletons further provide an alternative to the synthetic and allogenic biomaterials. The materials obtained from insects, sea weeds, shells, corals and other biological sources can be made largely available and cost effective for an affordable healthcare market. Furthermore, the tiny bacterial cell factories can deliver desired biomaterials with functional properties. They can be manipulated by the modern synthetic biological and genetic engineering tools to modify the product with unique characteristics to suit the target tissues. The microbial derived polymers can be decomposed to low weight molecular products by the microbial flora over synthetic polymers. Practise of mixed culture fermentation devoid of sterilization steps, recombinant strains that accumulate high polymer granules or exploring renewable substrates as fermentation feed can work a long way to reduce the cost of biopolymer production. Reduced cost, environmental benign nature, along with simple process technologies can boost the commercial production and widen the application horizon in tissue engineering and allied regenerative fields for these biocompatible microbial derived polymers. It would not be an overstatement if we mention that such bio-derived biomaterials are foreseen to provide patients with the desired organ/ tissue transplants before long. We can conclude with an optimistic note that these promising materials will bridge the gap between demand of donor organ unavailability and patients’ requirement in the most effective and efficient manner.

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Trends in Functional Biomaterials in Tissue Engineering and Regenerative Medicine Deepika Arora, Prerna Pant, and Pradeep Kumar Sharma

Abstract

Material science, particularly biomaterials, has developed as one of the mainstays of tissue engineering. Owing to their multifunctional, dynamic, and interdisciplinary capabilities, functionalized materials including natural, synthetic, or composites are now being extensively applied to organ development and regenerative medicine. Biomaterials usually do not have all the desirable attributes to be used as such in tissue bioengineering. Advancement in the various processes of surface modification and functionalization has improved the surface properties of biomaterials tremendously rendering them as optimum candidates for scaffolding and mimicking extracellular matrix (ECM), which is a prerequisite of tissue bioengineering. By treating with various functionalization systems (such as physicochemical, mechanical, radiation, and biological), the surface chemistry, conformation, bioactivity, biodegradability, bioavailability, biocompatibility, mechanical strength, etc., of biomaterials can be transformed to facilitate unified adaptation towards physiological biomimicking. These functionalized materials concurrently perform the edified functions which have been incorporated to

D. Arora (*) Biosystems and Biomaterials Division, National Institute of Standards and Technology, Gaithersburg, MD, USA Skeletal Biology Section, National Institute of Dental and Craniofacial Research, National Institutes of Health, Department of Health and Human Services, Bethesda, MD, USA P. Pant Department of Biomedical Engineering, University at Buffalo, State University of New York, Buffalo, NY, USA P. K. Sharma Food Drug and Chemical Toxicology Group, CSIR-Indian Institute of Toxicology Research, Lucknow, Uttar Pradesh, India # The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 B. Bhaskar et al. (eds.), Biomaterials in Tissue Engineering and Regenerative Medicine, https://doi.org/10.1007/978-981-16-0002-9_7

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address needs for medical and organ regeneration applications. In this chapter, the most applied methods of surface/bulk functionalization, their characterization and role in tissue engineering and development of different human organs and tissues are described. Keywords

Tissue engineering · Regenerative medicine · Functionalized biomaterial · Surface functionalization

Abbreviations ADSCs ALD Bio-GelMA BMSC CaP CAPS CCS CGSM CPCs CS-PG D-GUN ECM ELPs EPD FGF FPC GAGs HA HAG HEMA HS-PG HVOF LbL LENS PAA PCL PCT PDA PDGF PECVD PEG PEM

Adipose derived stem cells Atomic layer deposition Biomimetic gelatin methacrylamide Bone marrow mesenchymal stem cells Calcium phosphate Controlled-atmosphere plasma spraying Carboxymethylated chitosan Cold-gas spraying method Calcium phosphate cements Chondroitin sulfate proteoglycan Detonation-gun spraying Extracellular matrix Elastin-like polypeptides Electrophoretic deposition Fibroblast growth factor Fetal pulmonary cells Glycosaminoglycans Hyaluronic acid Hydroxyethyl methacrylate-alginate-gelatin Hyaluronic acid-g-poly(2-hydroxyethyl methacrylate) Heparin sulfate proteoglycans High velocity oxy-fuel spraying Layer-by-layer Laser engineered net shaping Poly(acrylic acid) Poly(ε-caprolactone) Proximal tubule cells Polydopamine Platelet-derived growth factor Plasma-enhanced chemical vapor deposition Polyethylene glycol Polyelectrolyte multilayer

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PEO PET PGA PGS PHB PLA PLGA POC POSS-PCU PPS PU PVA PVA PVD SF SLPC TIPS VEGF

7.1

217

Polyethylene oxide Polyethylene terephthalate Polyglycolic acid Poly(glycerol sebacate) Poly[(R)-3-hydroxybutyrate] Poly(lactide) Polylactic-coglycolic acid Poly(1,8-octanediol citrate) Polyhedral-oligomeric silsesquioxane-poly(carbonate-urea) urethane Polypropylene sulfide Polyurethane Poly(vinyl alcohol) Polyvinyl alcohol Physical vapor deposition Silk fibroin Somatic lung progenitor cells Thermally induced phase separation Vascular endothelial growth factor

Functionalized Biomaterials

The designing of novel smart functionalized biomaterial constructs compatible to human physiology is essential for various biological and clinical applications. Tissue engineering has evolved gradually after the collective efforts from materials science, physics, basic and clinical biology and it usually refers to the fabrication and embedding of suitable bio-scaffolds that can be entrenched with cells or functionally active biomolecules. The ultimate goal of tissue engineering is to tailor biocompatible and functional constructs that reinstate, uphold, and recover injured tissues or organs (Keane and Badylak 2014). Availability of suitable biomaterials for frame support is the key prerequisite for successful construction of any biological graft or engineered tissue/organs. Both natural and synthetic polymers are used in tissue engineering, but the usage of these materials as an embedding scaffold is not straightforward, and it requires surface/bulk modifications in many aspects, depending upon the target involved (Falentin-Daudre 2014). Biocompatibility, mechanical strength of tissue, release or inter-communiqué of indispensable elements or signals, and adverse immune response are the major challenges in the successful deployment of reproduced graft as an implant in medical purposes (Falentin-Daudre 2014; Wu et al. 2015; Yoshida et al. 2006). Unfortunately, most materials do not consent drugs to bind or allow biological cells to cultivate properly on their surfaces. Therefore, they need to be improved or functionalized by coating or adding suitable functional groups so that they can be accepted in a physiologically relevant environment. By virtue of these surface modifications, biomaterials tend to

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Fig. 7.1 Schematic showing several advantages in cellular function and behavior such as cell adhesion, proliferation, stem cells differentiation, etc., and controlled drug release by using surfacemodified/functionalized biomaterials

behave more biocompatible and appropriate for tissue engineering as shown in Fig. 7.1. Moreover, the surface modification also helps in reducing unwanted outcome of biomaterials in the host as well as improving controlled delivery of drug molecules. Besides serving as scaffolds, surface-modified biomaterials have shown promising outcomes in homing stem cells outside the body and an increased differentiation potential in the target cells with enhanced functionality in vitro (Zhang and Kohn 2012). Stem cells-loaded constructs require a natural milieu that augments and controls their growth and differentiation for functional tissue regeneration and therefore, biomimicking of the tissue environment plays a critical role. Functionalization of biomaterial has geared up the mimicking of tissue environment and allows stem cells to grow and differentiate fully into the desired cell type to construct an autologous graft (Chen and Liu 2016). The modified architectural surfaces can provide advantageous sites for cell–extracellular matrix, cellular attachment, and extracellular matrix remodeling and cell–cell interaction. Moreover, the implanting scaffolds must possess biocompatibility. The behavior of stem cells to biomaterials majorly relies on surface property and physicochemical properties of the biomaterial (Aiyelabegan et al. 2016). Prior to use of biomaterials, their

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functionalization by attachment of suitable molecules to their surface via different modification techniques such as physical or chemical methods is common in stem cell regeneration-based applications. Blending of extracellular matrix (ECM) ingredients in biomaterial is the most common practice during the functionalization of scaffolds. It imparts functionalities via reconstitution of ECM in the constructs that mimics the tissue nativity and allows the cells to grow in a natural microenvironment. Naturally occurring collagen fibril, particularly type 1 collagen, has been extensively used for such functionalization of various biomaterials (Meyer 2019). Gelatin, a denatured hydrolyzed derivative of collagen is also being used in embedding scaffolds as it is biocompatible and tunable to derive more suitable forms via chemical modifications (Santoro et al. 2014). Other novel strategies of surface/bulk modifications are now being developed and adopted to ensure the reproducibility, durability, cost-effectiveness, environment safety of functionalized biomaterials to ensure optimum performance, and efficacious application in tissue engineering.

7.2

Surface Functionalization Methods

Functionalized biomaterial-scaffolds form the basis of tissue bioengineering and regenerative medicine, since it bears the important tasks of cell adhesion, proliferation, differentiation, both in physical and chemical terms, in vitro and in vivo. Modifications made to their surface properties through various modification methods (physical/chemical/biological), could bring about a significant influence on cellular response and physiological environment (Kyzioł et al. 2017). Surface interaction of biomaterials plays a critical role in allowing transplantable grafts and implants (e.g., artificial liver, bones, and dental implants) to be accustomed well within the host body without mounting an overt immune reaction, which is the prime cause of failure of such grafts in clinic. Therefore, surface properties of biomaterials need critical evaluation before application to construct a graft or implant. A human body response towards a graft or implant majorly relies on biomaterial compatibility and physicochemical characteristics of its surface. Depending on the site of graft implantation and its projected application, various features are to be considered for the smart biomaterial to provide a desired response such as (a) hemocompatibility, (b) osseointegration, (c) overcome immune response (such as pyrogenicity and allergenicity), (d) non-toxicity, (e) carcinogenicity, (f) genetic changes (mutagenicity), (g) blood clotting (thrombogenicity), etc. (Tang et al. 2008; Wang 2013). The optimal choice of functionalization method to be employed in any particular case is primarily dependent upon the type of scaffold material (i.e., metals, natural polymers, synthetic polymers) intended to be used (Haider et al. 2020; Pina et al. 2019). Broadly, depending on the intended use, functionalization of biomaterials is achieved by surface modifications where the aim is to amend the surface of biomaterials. The most common methods and subcategories of surface alteration are shown in Fig. 7.2.

Fig. 7.2 Broad and subcategories of functionalization techniques employed in tissue engineering

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Usually, functionalization and alterations to the surface of substrates are performed by (a) roughening or patterning of the surfaces according to need, (b) coating/layering of the bioactive molecules or biofilms onto the surface to increase biocompatibility, and/or (c) incorporation of bioactive molecules for systematic delivery (Kyzioł et al. 2017). Though, each of the method of surface modifications is well-developed, yet their application to intended modification depends on the way by which a particular interaction/process can modify a biomaterial (Table 7.1). Owing to the different types of interactions such as a weak hydrogen bond, electrostatic interactions, denaturation of adsorbed molecules, etc., that take place during surface modification, the success of functional biomaterial largely depends on them since cell viability and ECM biomimicking are robustly determined by surface modifications (Rana et al. 2016). With the increase in demand for functionalized materials in clinical and translational settings, the exploration of surface techniques and biomedical substrates is continuously emerging. A schematic of some commonly used surface functionalization methods is shown in Fig. 7.3.

7.2.1

Surface Roughening and Patterning

Surface roughening is an effective and often simpler way of altering the surface topology without fetching any chemical changes. This method can cause a noteworthy upsurge in the surface area of the material with secure portions for cell movement, improves the cells attachment in scaffolds (Fig. 7.2). Different mechanical techniques, that include grinding, polishing, blasting, machining, oxygen and plasma deposition, are usually used for coarsening of surfaces, developing adhesion, and generating hydrophilicity in biomaterials. In particular, roughness or patterning of the scaffolds provides tremendous impacts on cell attachment, growth, and maturation (Narendrakumar et al. 2015). Mechanical methods are routinely used to modify surface properties of metallic biomaterials such as titanium and its alloys that are valuable to many medical and dental applications (Bruck 1978). Preconditioning with plasma etching is an important strategy before applying the coating of other desired methods, as etching causes substantial changes on the upper layer of the material (chain scission), thus advantageously supports the adjunction of biomolecules or biofilms on the surfaces (Kyzioł et al. 2017). In case of surface patterning, the structure featured type of amendments was employed (in its micro and nanoscale) on the material surfaces (Fig. 7.2). Lithography is a widely used technique of surface patterning that can control shape and size of a scaffold (Rashidi et al. 2014). Photolithography is the widely used form of lithography, where the photoirradiation is employed to create patterns (usually of size 5–100 mm) for stem cells research (Curtis and Wilkinson 1997).

Drug delivery, immunomodulatory techniques

Bone tissue engineering, drug delivery

Soft tissue, bone tissue engineering, and drug delivery

Bone tissue engineering

Bone tissue engineering

Physical adsorption of active biomolecules

Langmuir–Blodgett method

Physical vapor deposition

Electrophoretic deposition

Biomedical application Bone, cardiac, neuronal, and skin tissue engineering

Surface films and coatings

Functionalization methods Surface roughening and patterning

Titanium dioxide coated polymethyl methacrylate (PMMA) films, polyethylene oxide-polypropylene oxide triblock copolymers modified poly(lactic acid) Biomedical magnesium alloys, titanium alloys, titanium masked surfaces Hydroxyapatite nanoparticles-decorated aerographite scaffolds

Type I collagen immobilized polycaprolactone fibrous scaffolds

Bioactive glass (zirconium titanate composite thin films)

Modified biomaterial scaffolds Bioceramic or biopolymer scaffolds

Table 7.1 Biomedical applications of the surface modifications methods

To induce bioactivity on nanoparticles/surfaces to enhance cellular attachment and facilitate cellular functions

Bone healing through metal matrix composites

Functions Enhances attachment, proliferation, and differentiation, boosts vascular invasion, controlled deposition of calcium phosphate-based minerals Improves biocompatibility, bioactivity, improved material-host interface, and stability to implanted grafts Modulates cell behavior, improved proliferation, cell adhesion, growth, and differentiation Induction of macro-porosity in materials, in vivo behavior of polymers and their interaction with proteins

Taale et al. (2019)

Hacking et al. (2007), Narayanan et al. (2015)

Piwoński et al. (2013), Schöne et al. (2017)

Mattanavee et al. (2009)

Hashmi (2014), Mozafari et al. (2016)

References Gerberich and Bhatia (2013), Mitra et al. (2013)

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Smart grafts for tissue engineering

Bone tissue engineering, immunomodulation, gene/ cellular delivery systems, wound healing

Surface modification by addition of signaling biomolecules

Delivery systems for tissue engineering, scaffolds engineering

Cartilage tissue engineering

Tissue engineering of respiratory tract, cartilage and soft tissue, drug delivery, and wound healing Bone and blood vessel tissue engineering and biomedical devices

Radiation methods

Chemical vapor deposition (plasma polymerization) Sol-gel technique

Chemical methods (alkali acid hydrolysis; adsorption via covalent bonding)

Spraying techniques

Cationized gelatin modified poly(lactic acid) nanofiber scaffolds Manganese incorporated bioactive glass, hydroxyapatite, and βtricalcium phosphate treated biphasic calcium phosphate scaffolds Poly(lactide-co-glycolide) (PLGA) microspheres coupled with peptides, poly (ethylene glycol)methacrylate Apatite-coated polymeric scaffold, growth factor induced ECM polymer scaffolds (e.g., fibroblast growth factor (bFGF collagen composites)

Functionalize tantalum surfaces, poly-ɛ-caprolactonedecellularized bone matrix/ bio-Oss hybrid material

Sprayed alginate hydrogel, bilayered fibrin/polyurethane scaffold

Control release of factors, cell–matrix interaction, modulate cellular signaling

Surface modifications, sterilization, cell adhesion, tenability, and cellular responsiveness

Improves release of ions, cells responsiveness, and enhanced mechanical properties

Cell adhesion and differentiation

Enhances cell adhesion, protects from nonspecific binding, and allows postmodifications reactions

Provide adhesive properties for cells and enhance mechanical performances

Bishop et al. (2014), Dang et al. (2018), Davis et al. (2011) Fujisato et al. (1996)

Benson and Materials (2002), Mittal et al. (2010), Zhang et al. (2018b)

Barrioni et al. (2017), Houmard et al. (2012)

Al Kayal et al. (2020), Goor et al. (2017), Hasan et al. (2018), Mas-Moruno et al. (2015), Nyberg et al. 2017; Richbourg et al. (2019), Smith et al. (2007) Chen and Su (2011)

Al Kayal et al. (2020), Thiebes et al. (2015), Tritz et al. (2010)

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Fig. 7.3 A Schematic of some common surface functionalization methods

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Natural

Natural

Natural

Gelatin

Alginate

Chitosan

Delivery of angiopoietin-1 analogue Myocardial infarction

Chitosan–collagen hydrogel Chitosan–hyaluronan/ silk fibroin patch

Ischemic heart

Myocardial infarction

Alginate–chitosan Hydrogel

Injectable hydrogel

Myocardial infarction

To develop scaffolds for cardiac grafts

Injectable alginatehydrogel

Hydrolyzed denatured collagen with bioactive protein

Purpose of study To study the defect of right ventricular wall

Advantage Improved in vitro growth of endothelial and bone marrow cells and enhanced angiogenesis Better degradation kinetics when used with synthetic polymers Improved cardiac function, scar thickness, attenuated ventricular dilatation Promoted tissue repair by inhibiting cell death and increased angiogenesis and also effective in preventing LV remodeling Enhanced engraftment of stem cell and their survival Improved function and survival of endothelial cells Improved cardiac function, reduced LV

(continued)

Chi et al. (2013)

Miklas et al. (2013)

Liu et al. (2012)

Deng et al. (2015)

Landa et al. (2008), Ruvinov et al. (2011), Leor et al. (2009)

Gupta et al. (2007), Weinberg and Bell (1986)

References Miyagi et al. (2011), Segers and Lee (2011), Gupta et al. (2007)

Tissue/ organ type Cardiac Modification or alteration Growth factor immobilization (VEGF)

Table 7.2 Type of biomaterial and their common modifications/manipulation in tissue engineering

Type of biomaterial Natural

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Biomaterial used Collagen

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Tissue/ organ type

Type of biomaterial

Natural

Synthetic

Biomaterial used

Fibrin glue

Poly(lactic-co-glycolic acid

Table 7.2 (continued)

Remodeling of native microenvironment of myocardium

To study cardioprotection postMI

Binding of insulin-like growth factor (IGF)-1 to PLGA nanoparticles

To study preservation of cardiac function post-MI

Fibrin glue scaffolds

Poly(L-lactic acid), poly(ε-caprolactone), and collagen nanostructured scaffold

Fibrin-based clot formation for repairing cardiac wall damage in acute MI

Purpose of study

Fibrinogen and thrombin

Modification or alteration

Improved cardiac function and preservation of infarct wall thickness Cardiomyocytes embedded in scaffolds showed comparable growth and cellular organization to native myocardium Nanoparticles of IGF-1-complexed PLGA prolonged retention of IGF-1 in tissue, prevented cardiomyocyte death, and augmented LV function

dilatation, and enhanced wall thickness Sealing of the ruptured myocardium

Advantage

Chang et al. (2013a)

Mukherjee et al. (2011)

Terashima et al. (2008), Iemura et al. (2001), Okonogi et al. (2013), Kin et al. (2012), Wu et al. (2012) Christman et al. (2004)

References

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Synthetic

Synthetic

Synthetic

Carbon nanotubes

Polyurethane

Polyethylene terephthalate

To grow cardiomyocytes,To construct heart valves

Cardiovascular products such as vascular grafts and pediatric shunts

Polyurethane film

Polytetrafluoroethylene

To construct vascular grafts in different configurations

To support in vitro culture of neonatal rat cardiac cells

Scaffold made up of carbon nanofiber/ gelatin hydrogel

PET grafts coated with collagen or albumin

To study the growth of cardiomyocytes in vitro

Chitosan/carbon scaffold Improved cardiomyocytes survival and muscle functions by increasing the expression of myosin heavy chain, troponin T, and connexin-43 proteins Prevented pathological deterioration (e.g., ventricular dilation) Cardiomyocytes were able to grow and form a multilayered contractile tissue construct Grafts showed reduced thrombogenicity, decreased restenosis and hemostasis. Grafts were also less prone to calcification and biochemically inert, resistant to allergic and inflammatory response Promoted endothelialization with less calcification

Trends in Functional Biomaterials in Tissue Engineering and Regenerative. . . (continued)

Kudo et al. (2002), Nagano et al. (2007)

Aumsuwan et al. (2011), Verbelen et al. (2010), Miyazaki et al. (2007), Miyazaki et al. (2011)

McDevitt et al. (2003), Kütting et al. (2011), Silvetti et al. (2012)

Zhou et al. (2014)

Martins et al. (2014)

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Liver

Tissue/ organ type

Natural

Natural

Gelatin

Type of biomaterial Synthetic

Collagen

Biomaterial used Metals

Table 7.2 (continued)

Gelatin/polyurethane hydrogels

Mixture of heparin and gelatin

Gelatin hydrogels

Implantable hydrogel

Hydrogel of collagen/ chitosan Collagen type I-hyaluronan hybrid hydrogel Bioprinting of 3D liver using cell-based bioink comprised of hepatocytes and stellate cells Optimization of multilayered hepatocytes laminated into gelatin hydrogels Vasculature coating material to construct endothelialized vascular tree in decellularized livers To construct bioengineered liver

Hepatocytes culture

Purpose of study Construction of stents and heart valves

Collagen and glycosaminoglycans

Modification or alteration Titanium and stainless steel

Controlled pore sizes and better interconnectivity

Enhanced attachment of endothelial cells and better vascular patency

Increased longevity of hepatocytes for more than 2 months

Advantage Better strength and biocompatibility of stents and valves Imparted high mechanical integrity to hepatocytes and promoted molecular signaling Excellent biocompatibility Better liver microenvironment simulation

Xu et al. (2008)

Hussein et al. (2016)

Wang et al. (2006)

Mazzocchi et al. (2018)

Wang et al. (2003)

Zhang et al. (2017)

References Koh et al. (2011), O'Brien et al. (2010)

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Natural

Natural

Natural

Natural

Hyaluronic acid (HA)

Fibrin

Alginate

Chitosan

Alginate-based scaffolds Hydrogel consisting of glycyrrhizin (GL), alginate (Alg), and calcium (Ca) Microfibers of chitosan

To evaluate the formation of liver cells spheroids

In preparing the tissue seed by co-encapsulated hepatic cellular aggregates and endothelial cords In vitro hepatocytes culture Liver tissue engineering

Fibrin hydrogels

Fibrin hydrogels with PLGA

To optimize suitability for differentiation and stimulation of hepatocytes Endothelialized liver tissue

Used to provide better attachment and migration of liver cells

Fibrin-based hydrogel

HA hydrogels with cell adhesive proteins and peptides

Spheroids demonstrated improved liver functions

Maintained hepatocyte phenotype Better proliferation and liver-specific functions of cells

Implantable bioartificial liver with functional vascular network Ectopic implantation for tissue expansion

Better proliferation and maintenance of hepatoblast and hepatic progenitor cells Better hepatocytes differentiation and viability

Trends in Functional Biomaterials in Tissue Engineering and Regenerative. . . (continued)

Lee et al. (2010)

Jammalamadaka and Tappa (2018) Tong et al. (2018)

Stevens et al. (2017)

Wang and Liu (2018)

Van Vlierberghe et al. (2011), Christoffersson et al. (2018), Turner et al. (2007) Bruns et al. (2005)

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Tissue/ organ type

Type of biomaterial

Natural

Natural

Natural

Biomaterial used

Polyhydroxyalkanoates

Cellulose

Agarose

Table 7.2 (continued)

Agarose–chitosan scaffold

Nanofibrillar cellulose hydrogel Nanocrystals of alginate and cellulose hydrogel

Scaffold of poly (3-hydroxybutyrate-co3-hydroxyvalerate-co3-hydroxyhexanoate) (PHBVHHx)

Chitosan nanofibers with fibronectin coating on the surface

Modification or alteration Hybrid scaffolds of chitosan and gelatin

Liver tissue construction using human umbilical cord multipotent stromal cells (MSCs) and hepatocyte-like cells Liver tissue engineering To develop a hybrid bioink for bioprinting a 3D liver-mimetic honeycomb In vitro 3D liver model with primary hepatocytes

Purpose of study To organize hepatocytes microstructures Co-cultured liver models

Promoted liver cell culture Hybrid bioink possesses excellent shear-thinning property 3D model displayed liver-mimetic physicochemical properties, biocompatibility, and enhanced metabolic activity

Advantage Scaffolds were more suitable for hepatocyte culture. Hepatocytes formed colonies in a co-culture platform for a prolonged period of time Liver tissue performed very similar to the native organ till 28 days of culture

Tripathi and Melo (2015)

Bhattacharya et al. (2012) Wu et al. (2018)

Su et al. (2014)

Rajendran et al. (2017)

References Jiankang et al. (2009)

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Synthetic

Synthetic

Synthetic

Poly(ethylene glycol) (PEG)

Poly(vinyl alcohol) (PVA)

Poly(lactide-coglycolide) acid (PLGA)

Biodegradable PLGA hydrogels

PVA/gelatin hydrogels

Transparent PVA hydrogels

Used for developing 3D hepatocellular carcinoma (HCC) model A transplantable 3D liver structure for end stage liver disease

Co-culture (hepatocytes and patient derived iPSCs)-based perfusable 3D organoids To fabricate a 3D liver tissue representing a native hexagonally arrayed lobular structure Used for biomedical applications

Co-culture of hepatocytes and nonparenchymal cells

PVA hydrogel resembled microstructure similar to the porcine liver tissue Development of a long-term HCC model to study migration

3D culture displayed an advanced hepatic function for at least 5 months

Perfused liver-on-achip with enhanced viability of organoid

Encapsulation of hepatic cells

A hydrogel of variable chain length of PEG polymer conjugated with bioactive factors PEG hydrogel

PEG hydrogel

Formed a biocompatible matrix that allowed survival of encapsulated primary hepatocytes Better growth and function of hepatocytes in culture

Used for liver-on-achip model

PEG hydrogels

Trends in Functional Biomaterials in Tissue Engineering and Regenerative. . . (continued)

Kim et al. (1998)

Moscato et al. (2015)

Jiang et al. (2011)

Ng et al. (2017)

Schepers et al. (2016)

Bhatia et al. (2014)

Lee et al. (2015)

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Tissue/ organ type

Type of biomaterial

Synthetic

Biomaterial used

Poly(lactide) (PLA)

Table 7.2 (continued)

PLA hydrogels

To reconstruct 3D stacked hepatocyte

Membrane of degradable and microporous poly(d, l-lactide-co-glycolide) Biodegradable copolymers of L-lactide: Glycolide

Co-culture of rat hepatocytes with stellate cells

Fabrication of absorbable vascular anastomosis device (AVAD)

Transdifferentiation of stem cells into mature hepatocytes

Purpose of study

Collagen-coated PLGA

Modification or alteration survived well on the 3D polymer scaffolds under both static and flowing conditions Biomimicking the microenvironment to allow stem cells differentiation into mature liver cells Better liver-specific function as compared to the monolayer culture AVAD showed compatibility in absorption and intact anastomosis in minipig showing success in the liver transplantation A long-term culture of rapidly self-organizing three-dimensional liver cell spheroids

Advantage

Riccalton-Banks et al. (2003)

Park et al. (2019)

Kasuya et al. (2012)

Li et al. (2010)

References

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Synthetic

Synthetic

Poly(e-caprolactone) (PCL)

Poly(acrylic acid) (PAA)

To determine the competency of mouse hepatic cells in a culture To provide a framework for capillary-like network

Electrospun nanofiber of PCL/chitosan

Polymeric mixture of PAA and polyethyleneimine conjugated with elastin-like polypeptides (ELPs)

Polycaprolactone (PCL) framework with collagen bioink To study different liver functions such as differentiation, morphology, aggregation, etc

To study the differentiation potential of MSCs into hepatic cell type Liver tissue engineering

PLLA and gelatin based electrospun nanofiber scaffolds Hybrid PCL-ECM scaffolds

To assess adhesion and proliferation of hepatocytes

Electrospun nanofibers of poly(L-lactic acid) (PLLA) coated with type I collagen

Better simulation of microenvironment for a heterotypic co-cultured 3D liver Rat hepatocytes cultured with ELP-polyelectrolyte conjugates profoundly exhibited high liverspecific functions

PCL-ECM scaffolds recapitulated a niche microenvironment for hepatocytes Better matrix for the construction of liver models

Discretely aligned nanofibers (disAFs) represented a suitable method of large-scale hepatic cultures Controlled migration of hepatic stellate cells

(continued)

Janorkar et al. (2008)

Lee et al. (2016)

Semnani et al. (2017)

Grant et al. (2017)

Zhang et al. (2018a)

Feng et al. (2010)

7 Trends in Functional Biomaterials in Tissue Engineering and Regenerative. . . 233

Tissue/ organ type Lung

Type of biomaterial Natural

Natural

Natural

Biomaterial used Albumin

Fibrin

Collagen

Table 7.2 (continued)

As a hemostatic material in minimizing the damage to pulmonary arteries during thoracic surgery To develop a matrix material suitable to simulate mechanical properties of a single alveolar wall

Fibrinogen/thrombinbased collagen fleece (TachoComb).

Collagen–elastin fiber hydrogel

To construct lung alveolar-like structures

Used as an implanting gel to allow angiogenesis on rat lung surface

Purpose of study Recellularization of decellularized lung scaffold

Collagenglycosaminoglycan scaffold

Fibrin gel

Modification or alteration Albumin with collagen/gelatin scaffolds

A better Young’s modulus was noted in collagen-elastin hydrogel as compared to collagen alone

Advantage Grafting of albumin on decellularized lung scaffolds allows cell engraftment and cell– tissue interaction Implanted gel resulted in neo-angiogenesis on lung surface by controlling adjacent physical and chemical signals Dissociated lung cells demonstrated the capability to form lung histotypic structures TachoComb based hemostasis of pulmonary artery injury was safe and reliable

Dunphy et al. (2014)

Ikeda et al. (2012)

Chen et al. (2005)

Mammoto and Mammoto (2014)

References Aiyelabegan et al. (2016)

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Bone

Synthetic

Synthetic

Poly-lactic-co-glycolic acid (PLGA)

Polyethylene glycol (PEG)

Synthetic

Synthetic

Polyglycolic acid (PGA)

Poly(lactide-coglycolide) (PLGA)

Natural

Gelatin

Polyethylene glycol (PEG)-substituted polylysine/PEBP-bPBYP-g-PEG Copolymer PLGAPCL

Crosslinking of poly (ethylene glycol) to decellularized lung scaffolds from rats

Polyglycolic acid (PGA) with somatic lung progenitor cells (SLPC) Scaffolds of porous foam of PLGA

Gel foam sponge

Bone regeneration

To construct 3D pulmonary tissue by incorporating fetal pulmonary cells (FPC) to porous PLGA scaffold To increase biomechanical property and antienzymatic stability of lung decellularized scaffolds Nano-drug delivery systems to target lung metastasis

Implantable inoculum of gel foam and fetal rat lung cells for lung regeneration To promote alveolar tissue growth

Osteointegration and bone formation were comparable to the preformed,

Better and controlled drug (paclitaxel) release

Crosslinking of poly (ethylene glycol) had no toxicity with decellularized lung scaffold

PLGA foams scaffolds facilitated ingrowth of FPC

Better in vitro alveolar tissue growth

Regeneration of alveolar-like structures was achieved

(continued)

Ulery et al. (2011), Shim et al. (2015)

Zhang et al. (2015)

Xie et al. (2020)

Mondrinos et al. (2006)

Cortiella et al. (2006)

Andrade et al. (2007)

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Tissue/ organ type

Synthetic

Synthetic

Poly-L-lactic acid (PLLA)

Type of biomaterial

Poly(epsiloncaprolactone)

Biomaterial used

Table 7.2 (continued)

Bone replacement material

Composite of made up of poly(L-lactide-coglycolide)/ hydroxylapatite and beta-tricalcium phosphate developed by laser sintering Calcium phosphate cements (CPCs) with ultrafine fibers of poly (epsilon-caprolactone) Composite scaffold containing a recombinant bone morphogenetic protein 2 (rhBMP2)

Induction of angiogenesis for bone repair

Microsphere scaffolds of poly(lactide-coglycolide) (PLGA)

Bone formation

Used as bone filler material

Purpose of study

Modification or alteration

Composite material facilitated interconnective channels and cement resorption for bone growth PLLA scaffolds offered enhanced carrying capacity of rhBMP2 for inducing bone formation

non-resorbable membrane of titanium mesh An obvious vascular growth was noted in PLGA scaffolds harboring VEGF releasing ADSCs and endothelial cells Possibility of fabrication of porous scaffolds for bone constructs

Advantage

Chang et al. (2007)

Zuo et al. (2010)

Simpson et al. (2008)

Jabbarzadeh et al. (2008)

References

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Synthetic

Natural

Poly(ethylene glycol) (PEG)

Collagen

Ethylene glycol methacrylate phosphate incorporated PEG hydrogels Collagen scaffold containing carbonatesubstituted hydroxyapatite (HA) crystals formed by rapid prototyping Collagen scaffold with osteoinductive HA particles

Nanofibrous membrane of electrospun poly-Llactic acid (PLLA) equipped with collagenous guided bone regeneration (GBR) membrane Copolymerized with tyrosine-derived polycarbonates

To incorporate microchannels to allow the flow of nutrient rich media throughout the scaffold Bone tissue regeneration

To study the differentiation potential of mesenchymal stem cells (MSCs) into osteogenic lineage To study the mineralization potency and MSCs viability

Regeneration of dense bone

Gleeson et al. (2010)

Capable of promoting osteogenesis and repair of critical-sized bone defects

(continued)

Sachlos et al. (2006)

Nuttelman et al. (2006)

Briggs et al. (2009)

Cai et al. (2010), Shim et al. (2010)

Able to transport mass nutrient thoroughly within the scaffold

Promoted spreading and adhesion of hMSC

Allowed abundant bone formation

Electrospun nanofibrous membrane improved regeneration of cortical bone

7 Trends in Functional Biomaterials in Tissue Engineering and Regenerative. . . 237

Dental implants

Tissue/ organ type

Metallic

Natural

Gelatin

Titanium

Type of biomaterial Natural

Biomaterial used Chitosan

Table 7.2 (continued)

Scaffolds of biomimetic gelatin methacrylamide (bio-GelMA) hydrogel fabricated by thermally induced phase separation (TIPS) Plasma-sprayed hydroxyapatite (HA) coating on titanium surfaces Surface coating of different composites such as carbonate apatite (CO3–Ap),

Modification or alteration Nanofibers of electrospin chitosan containing hydroxyapatite and crosslinked with genipin Copper (II)–chitosan containing strontium– hydroxyapatite Gelatin scaffolds crosslinked with transglutaminase

To obtain biocompatibility and mechanical strength of implants To improve the bioactivity of implants

To study the deposition of novel calcium phosphate To develop a matrix for cell substrate and growth factor release system To mimic bone ECM and physical architecture

Purpose of study Bone tissue engineering

Low cost alternative to coat titanium surface for constructing dental implants Hydroformed HAp has greater osteoconductivity than HAp

Advantage Composite scaffold facilitated proliferation, differentiation, and maturation of osteoblast-like cells Better osteogenesis with angiogenesis and antibacterial activity Developed scaffold was a suitable candidate for bone constructs Bio-GelMA facilitated better osteogenic differentiation of adipose derived stem cells (ADSCs)

Kuroda and Okido (2012)

Vindigni et al. (2009), Hung et al. (2013)

Fang et al. (2016)

Echave et al. (2019)

Gritsch et al. (2019)

References Frohbergh et al. (2012)

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Skin

Collagen

Natural

To impart antibacterial activity in the dental implant

Silver nanoparticle loaded composite of chitosan/hyaluronic acid coated on titanium surface via layer-bylayer method Hydrogels of collagen with carboxymethylated chitosan (CCS) prepared by enzymechemical double crosslinking To use the hydrogel as skin scaffold

Surface fabrication of titanium implants for better adhesion

To create bioactive surface of the implant by increasing the biomineralization

Covalent modification of type I collagen by coating of polydopamine (PDA)

HAp/collagen, or HAp/gelatin by using pyroprocessing and hydroprocessing Coating a film of multilayered casein/ chitosan by layer-bylayer technique

Hydrogel promoted skin regeneration at the wound site

Multilayer film stimulated osteogenic differentiation, cell attachment, and proliferation of human mesenchymal stem cells (HMSCs) Surface modification by PDA coating enhanced differentiation and adhesion of MC3T3E1 cells Effective antifouling activity

Trends in Functional Biomaterials in Tissue Engineering and Regenerative. . . (continued)

Cao et al. (2020)

Zhong et al. (2016)

Yu et al. (2014)

Li et al. (2016)

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Kidney/ bladder

Tissue/ organ type

Synthetic

Natural

Chitosan

Type of biomaterial Natural

Poly [(R)-3hydroxybutyrate] (PHB)

Biomaterial used Chitosan

Table 7.2 (continued)

Cell surface heparin sulfate proteoglycan and chitosan

Modification or alteration Nanoparticles (ZnO, Fe3O4, and au) loaded electrospun hybrid poly (lactic acid)/ chitosan biomaterials Surface modification of PHB with PEG or EDA using radiofrequency glow discharge method To study the transcellular pathways such as transport of water and ions without losing the function of rat renal proximal tubule cells (PCT)

Graft for bladder reconstruction

Purpose of study To prepare scaffold for skin tissue engineering

Scaffolds repressed the growth of calcium oxalate and enhanced uroepithelial cell viability Scaffold played a significant role in PTC proliferation and differentiation

Advantage Hybrid nanofibers mimicked ECM substantially close to native state

Chang et al. (2013b)

Karahaliloğlu et al. (2016)

References Radwan-Pragłowska et al. (2020)

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Surface Films and Coatings

These surface modification techniques involve “surface activation via attachment of polymer chains or synthesis of functionalities on surfaces” without altering the compositional characteristics of the material. It comprises the coating of an additional functional group to the surface (coating) or any other polymer-based films. There are broadly three methods of functionalization, namely physical chemical, and via radiation (Thakur et al. 2017). Physical modifications basically involve physical connections and depend upon the type of interactions (like electrostatic interactions, hydrogen bonding amid hydrophilic/hydrophobic residues) and binding affinity between the shielding layer and the adsorbed functional groups (Zhou and Pang 2018). Coatings of proteins/peptides like collagen, laminin, integrin, fibronectin, chitosan, and gelatin (ECMs) are commonly done by physical modification. The thickness of the resulting layer is largely dependent on the interaction among various factors like viscous force, surface tension, gravity (Richbourg et al. 2019). Chemical modification methods usually utilize deposition/coating of specific molecules via chemical conjugations methods such as alkali hydrolysis, covalent adsorption, covalent interactions, ionic interactions, acid etching, etc. Some of the common chemical, physical, and radiation methods are described below.

7.2.2.1 Physical Methods 7.2.2.1.1 Physical Adsorption of Active Biomolecules Physical adsorption modification technique is known for its simplistic procedure, which involves incubation of biomaterial substrate into a biomolecule’s solution (Rana et al. 2016). Upon incubation, these biomolecules attach to the surface of the substrate through various electrostatic interactions (such as hydrogen bond, weak van der Waals forces, or hydrophobic interaction). Preconditioning with plasma etching process is found to be useful with this type of method. Due to the increment in hydrophilicity through etching the adhesion strength of the material surfaces gets improved. The limitations like inability to control the orientation of adsorbed molecules, weak binding with surfaces are noticed with this method. However, being a simple and gentle procedure, this technique is reported to be suitable to deal with fragile assemblies and biomolecules. The application of this type of functionalized methods includes proteins/biomolecule, ECM (vitronectin, laminin, integrin, matrigel, growth factors like VEGF) deposition on to the surface of the scaffold materials for improved proliferation, cell adhesion, growth, and differentiation (Joddar and Ito 2011; Tallawi et al. 2015). 7.2.2.1.2 Langmuir–Blodgett Method Langmuir–Blodgett is one of the promising methods of functionalization and is useful for the creation and deposition of single or multiple organic thin monolayer (also known as Langmuir film) of amphiphilic substances on the exterior of solid material (Hussain et al. 2018; Roberts 2013). By virtue of the amphiphilic nature of Langmuir films, the arrangement of molecules takes place at the air–water interface

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such that it offers several advantages such as: (a) defined thickness of monolayer, (b) homogeneous deposition to larger surface areas, and (c) multilayered structure of different compositions. Langmuir monolayer methods are also considered as powerful measures for studying interfacial properties, such as structural and physicochemical attributes of the polymers under experimental conditions (Przykaza et al. 2019). This technique also finds its application in tissue engineering. In a recent study, the authors reported that a surface-modified Ti-matrix (developed by the deposition of monolayer of dihexadecyl phosphate onto a titanium surface) demonstrated a substantial improvement in the proliferation of osteoblasts and therefore, supporting the use of this method for the modification of osteogenic biomaterials (de Souza et al. 2014; Rana et al. 2016). 7.2.2.1.3 Physical Vapor Deposition Physical vapor deposition (PVD) is another coating process which includes a number of methods of thin-film deposition via condensation of a vaporized solid onto a substrate. PVD is an eminent vacuum coating procedure that supports the deposition as mono- or multilayered pattern, multi-graduated coating, and also allows configuration or structural changes in alloys (Baptista et al. 2018; Inspektor and Salvador 2014). Importantly, atomic deposition process can be used in any medium including vacuum, gaseous, plasma, or electrolytic environment (Porteiro et al. 2018). Evaporation

Evaporation is also known as vacuum deposition, is a very simple process which involves the thermal evaporation of atoms or molecules. These particles travel in the deposition chamber without colliding with the gaseous molecules present in it and allow condensation over the surface of material (Mattox 2010). Common thermal mechanisms that are frequently used are resistive heating and electron beam heating. In resistive heating mode, materials are allowed to evaporate via heating with a filament keeping on a boat crucible (Rockett 2008). On the contrary, e-beam heating is used for refractory substances. An e-beam gun having accelerating electrons with a high voltage (10–20 kV) is used either electrostatically or magnetically collimated and strike upon the surface of material undergoing evaporation (Rempe 2019; Yu and Lee 2014). Deposition by Sputtering

Sputtering is the process by which highly energetic ions are used for bombardment upon required surface using scattering of the solid surface atoms in a backward direction (Kammara et al. 2016). It is a kind of etching procedure which is appropriate for micromachining, surface cleaning, depth profiling, etc. (Wasa et al. 2012). The deposition rate is usually smaller as compared to the process of evaporation; however, it has several advantages over other PVD methods (Ye 2015). It can allow bombardment of the target surface with energy electrons.

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Plasma immersion ion implantation and deposition (PIII&D)

It is a very flexible technique that holds the advantage of simultaneous conduction of ion implantation as well as their deposition. This method is an amalgamation of plasma and ion beam technology (Lu et al. 2012). Taking into consideration the industrial standpoint, this is a very useful method to be applied to biomaterials since it can be used for irregular-shaped materials and can control coating composition (Paterlini et al. 2017). It can selectively enhance surface properties (up to several hundred nanometers) while keeping the bulk part unmodified. Thus, it finds various uses in the biomedical industry (Chu et al. 2002; Yoshida et al. 2013). 7.2.2.1.4 Electrophoretic Deposition Electrophoretic deposition (EPD) is used to form thin/thick films or coatings under the electric field. In this method, an electrophoresis mechanism (that allows charged particles to move in an electric field) is followed to deposit the suspended particles on a substrate in an orderly fashion (Neirinck et al. 2013). Thin-film deposition methods are simple and provide films of high purity and good structural properties. However, due to some limitations like high processing costs, requirement of stringent instrumentation and gaseous waste treatment, researchers are trying to explore alternative techniques (Darband et al. 2017; Furuya et al. 1972). In recent times, the curiosity in EPD technique has been widely increased, as this can fabricate surface uniform deposition to recapitulate better microstructural homogeneity. This method can also be used to fabricate bulk materials, coatings, nanomaterials, and 3D complexes and porous structures (Boccaccini et al. 2010) (Corni et al. 2009; Lovsky et al. 2010). 7.2.2.1.5 Spraying Techniques Thermal spraying techniques collectively refer to the process of spraying of metal/ non-metal material in a molten or semi-molten state over the substrate (0.5–2 mm thick coatings). In this method, a feedstock of particles of molten, semi-molten metals, or ceramics is pushed towards the target surface to deposit a coating. Different methods such as flame spraying, atmospheric plasma spraying, arc spraying, detonation-gun spraying (D-GUN), high velocity oxy-fuel spraying (HVOF), vacuum plasma spraying, controlled-atmosphere plasma spraying (CAPS), cold-gas spraying method (CGSM) are frequently involved in these techniques (Oyinbo and Jen 2019; Pawlowski 2008; Singh et al. 2012; Sorrentino and Tiginyanu 2011). Among these, plasma spraying is common and comprises spraying of atmospheric or vacuum plasma. In this method, plasma-forming gases like Ar, He, H2, N2, etc., are used for igniting the high-frequency electrical arc between the cathode and anode, which generates a high energy and causes melting and spraying of particles onto the substrate leading to the formation of splat, splat layering, and coating before they actually flatten out and solidify over the material (Fauchais 2004). This method is deemed favorable for deposition of bioceramic coatings (Sorrentino et al. 2011). Owing to their flexibility and versatility, these spray methods are found to be useful for broader range of applications (repair to

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replacement) in biomedical and tissue engineering (Cizek and Matejicek 2018; Heimann 2018).

7.2.2.2 Chemical Methods 7.2.2.2.1 Adsorption Via Covalent Bonding In this functionalization technique, biomolecules are chemically adsorbed (covalently bind) to the biomaterial surface via relevant functional groups. Unlike physical adsorption, chemical adsorption is known for more efficient coating, higher surface stability, being retained over a longer period (Tallawi et al. 2015). Furthermore, with this method, higher biocompatibility, immobilization of biomolecules onto polymeric biomaterials, integration of adhesive peptides on substrate surface are achievable, which makes this technique even more important for tissue engineering (Bagno et al. 2007). Recently, Pandey et al. have shown that hybrid nanoscale modified surfaces (using silanization adsorption) displayed venerable properties for supporting cell adhesion and growth (Hasan et al. 2018). In another study, this technique has been used as an antifouling coating so that after implantation, the material can be prevented from nonspecific protein and cell adhesion (Goor et al. 2017). 7.2.2.2.2 Alkali Acid Hydrolysis The main aim of alkali acid hydrolysis-based functionalization method is to improve the water solubility of the substrate. In this method, between the polymer chains, due to diffusion of protons, cleavage of ester bonds takes place and results in formation of different functional groups such as hydroxyl ( OH) and carboxylic ( COOH) on the surface upon hydrolysis (Guo et al. 2015). Leonor et al. have developed a bonelike apatite on a biodegradable polymer of starch by using this method. Due to rise in the functional group (e.g., hydroxyls and carboxylic acid) on surface, a heterogeneous apatite growth along with the calcium binding on SEVA-C specimens was recorded which demonstrated the potential use of this method for surface functionalization of polymers used for bone tissue engineering applications (Leonor et al. 2007). In another study, attempt has been made to improve the efficiency of this surface alkali acid hydrolysis method. This technique has been used with some modifications (using citric acid as the wash solution) to improve the hydrophilicity of polymer surface (Qin et al. 2019). The diminution in the surface-water contact angle and increment in the surface roughness were observed on the polymer surface which advocate this method for functionalization of polymeric biomaterials for different applications in tissue engineering. 7.2.2.2.3 Chemical Vapor Deposition Chemical vapor deposition (CVD) method involves the deposition of thin films that utilizes chemical reactions involved in vapor phase precursors. In this technique, substrate is exposed to one or more volatile precursors that react on the surface of substrate to yield required deposition. In this approach, heat is used as source of energy, which initiates and controls the process. A high temperature is required for

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deposition, which is useful for limiting the types of substrates and materials used for coating, particularly those that are highly thermally sensitive (Choy 2003; Kaivosoja et al. 2013). Notably, heat is not the sole source of energy, with plasmas and photons being used as well (Griesser 2016). Nowadays, many types of CVD methods are in use: 1. APCVD: atmospheric pressure chemical vapor deposition (leads to the formation of a uniform coating), 2. LPCVD: low-pressure chemical vapor deposition (causes an increase in hardness and resistance to corrosion), 3. LECVD: laser-enhanced chemical vapor deposition (causes enhanced resistance to wear and corrosion), 4. PECVD: plasma-enhanced chemical vapor deposition (responsible for an increase in wear as well as corrosion resistance), 5. PACVD: plasma-assisted chemical vapor deposition (enhances chemical stability, improves biocompatibility, and imparts corrosion resistance) (Thakur et al. 2017). Among these, plasma-based technologies have gained tremendous popularity as it can be useful for the inception of immobilized proteins or biomolecules onto the biomaterial surface. PACVD (also referred sometimes as PECVD) is the common plasma-based method where the UV radiation is been employed to generate radicals and react with surfaces with a low thermal resistance. Plasma-Enhanced Chemical Vapor Deposition

Plasma-enhanced chemical vapor deposition (PECVD) uses plasma as a source of energy in order to trigger ions and radicals present in the chemical reactions, resulting in the layer formation on materials surface. The most prominent advantage of this technique is that it uses low temperature that facilitates the deposition of layers which cannot be deposited via high temperature. Also, its deposition rate is comparatively higher and controlled easily. Plasma Polymerization

Plasma polymerization or glow discharge polymerization is a process derived from PECVD that uses organic or organometallic precursors to construct thin films that are plasma-polymerized (Santhosh et al. 2018). It includes fragmentation followed by the deposition of organic/organometallic precursors while allowing the functional groups to remain from the monomers (Chen and Su 2011; Laurano et al. 2019; Montazer and Harifi 2018). Plasma has been categorized into different categories based upon the resulting consequences of the interaction with the material such as plasma polymerization, plasma treatment, and plasma etching. Plasma treatment is different from plasma polymerization in that it uses gases like N2, O2, NH3, CH4, Ar for embedding chemical modification onto a surface or for the creation of radicals required for crosslinking followed by surface grafting (Laurano et al. 2019; Santhosh et al. 2018).

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Atomic Layer Deposition

Atomic layer deposition (ALD), also one of the subclasses of chemical vapor deposition, practices the successive use of a gas phase chemical process (into the reaction chamber) for construction of thin films of a variety of materials (Johnson et al. 2014). Various advantages like unvarying deposition of conformal films and controllable thickness (3D surfaces) have generated a substantial interest in this method. The application range of this method is extensive, it includes nanopatterning, 3D nanoporous structures, fuel cells, desalinations, catalysis, and various biomedical applications (Oviroh et al. 2019).

7.2.2.2.4 Sol-Gel Technique The sol-gel process which is usually a wet-chemical technique involves chemical reactions such as hydrolysis and condensation of metal or silicon alkoxides to develop highly pure inorganic oxides or hybrid materials. In this method, the solution evolves gradually towards the formation of a gel-like network containing both a liquid phase and a solid phase (Danks et al. 2016; Laurano et al. 2019). It is advantageous because it can provide homogeneous hybrid materials at a low temperature, and this can incorporate a number of compounds with a high level of purity (Suslick 2001). Single/multiple-component ceramic materials, thin solid films, porous materials, glass fibers, and catalytic materials can be prepared by this technique and these are frequently used in bio-tissue engineering (Esposito 2019). Manuel et al. have employed this sol-gel synthesis technique (using calcium nitrate tetrahydrate and triethyl phosphite precursors) for the fabrication of highly porous scaffolds of calcium phosphate (CaP). They report that this coating not only improves the mechanical strength, but also modifies the topography and composition of the scaffold surface which supports bone formation (Houmard et al. 2012).

7.2.2.2.5 Layer-by-Layer (LbL) Deposition Thin film fabrication is usually performed by layer-by-layer deposition method which is a prominent method of thin-film fabrication. Classically, in this assembly process a stepwise adsorption of complementary molecules takes place on a substrate surface under the influence of electrostatic and/or non-electrostatic interactions (Gentile et al. 2015). Numerous studies have advocated that LBL technique is a competent, facile, flexible, and versatile approach to coat biomaterials with controlled structures, properties, and functions (Decher 1997). Different nanoformulations such as core–shell nanoparticles or nanocapsules have been attained by this method that are used for controlled drug-delivery systems (both single and multidrug targeting) (De Villiers et al. 2011). It has been reported that the formation of particles by layer-by-layer method facilitates the production of multifunctional, stimuli-responsive carrier systems. Moreover, improved crosslinking in collagen/hyaluronic acid (Col/HA) polyelectrolyte multilayer (PEM) film with LBL method has enhanced proliferation and improved spreading and differentiation of mesenchymal stem cells (Rana et al. 2016).

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7.2.2.3 Radiation Methods Radiation methods are one of the most important functionalization approaches for modifying polymer surface properties. This approach has been quite significant in drug-delivery systems where small biomolecule or stimuli are incorporated and release a required amount of drug under physiological conditions (Zhang et al. 2018b). Previously, ultraviolet or selective synchrotron radiations have been exploited to improve surface characteristics of polymers such as polyurethane and polystyrene-film-based construction systems. Competent surface functionalization was observed in both types of films suggesting the potential of this technique to functionalize polymer surfaces (Weibel 2010). Another study of surface modification by grafting acrylic acid onto poly(ethylene terephthalate) film using gammarays and embedded silver nanoparticles has advocated that the hybrid (PET-g-PAA/ Ag) film has quite strong antibacterial features (Ping et al. 2011). Functionalization through X-ray synchrotron radiations provides a high spatial resolution probe which plays an essential role to study the heterogeneity from micron-scale to meso- and nanoscale in hybrid biomaterials (Mastrogiacomo et al. 2019). Various studies have reported that this technique is useful for constructing ceramic scaffolds and other bone regeneration applications (Polo-Corrales et al. 2014).

7.2.3

Surface Modification by Addition of Signaling Biomolecules

Construction of biocompatible and suitable biomaterials is the aim of surface modification techniques. Clinical applications of various tissue-engineered grafts remain compromised due to several critical factors including functionality and hostgraft rejection. Though these attributes pose a great challenge in the synthesis of novel biomaterials yet various techniques of surface modification have been employed to achieve functionality of various grafts such as vascular grafts, artificial liver, dental implants, etc. Indeed, a considerable advancement has been reported across the globe related to the synthesis of these grafts for clinical applications. In addition to the classical methods of surface modification (e.g., physical and chemical), biological modification of biomaterials has gained significant momentum in the recent past (Su et al. 2018). Application of signaling biomolecules to the surface of biomaterials has enriched the process of surface functionalization owing to their biocompatibility, bioactivity, and relative abundance in nature. Bioactive biological materials possess natural complexity and multifunctional domains that allow their integration with any biomaterial for a suitable biological response during the construction of biocompatible constructs. Hence functionalization of biomaterials using bioinspired surface technologies has opened a new avenue to revolutionize the area of regenerative medicine and body implants. In addition to bioactive chemicals, various naturally occurring bioactive biomolecules such as peptides/proteins, growth factors, and glycosaminoglycans (GAGs), etc., can be easily integrated with biomaterial for desired application (Su et al. 2018; Tallawi et al. 2015). Integration of these bioactive molecules has conferred manifold advantages to biomaterials in their functionality and

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biocompatibility. For example, peptides derived from fibronectin containing ArgGly-As (RGD) sequence have been extensively studied and used for better cell adhesion to scaffolds (Hersel et al. 2003). Remodeling of ECM is critical for tissue engineering and biosurface mimicking is essential to provide a homing atmosphere to the seeded cells. Different proteins/peptide sequences of collagen, fibronectin, laminin, gelatin, etc. have been successfully used in mimicking ECM for growing cells (Bierbaum et al. 2012; Tallawi et al. 2015). Tissue microenvironment is quite complex in terms of its biological composition, chemical nature, and physiological functions besides mechanical strength. Certain soluble factors such as growth factors, hormones, cytokines, etc., play an essential role in defining a tissue microenvironment. Various isolated and purified growth factors [e.g., vascular endothelial growth factor (VEGF), fibroblast growth factor (FGF), platelet-derived growth factor (PDGF)] have been shown to regulate and initiate angiogenesis, cell growth, migration, regeneration, and differentiation of different cell types including homing of stem cells integrated onto a biomaterial. Osteointegration, osteogenesis of bone and dental implants, antifouling surfaces, antithrombogenicity, bactericidal biofilm coatings, anti-corrosive effects, and drug-delivery systems are some other important examples of bioinspired surface modifications (Su et al. 2018; Tallawi et al. 2015).

7.3

Functionalized Scaffolds Towards Organ Development

Tissue engineering has paved a way to develop promising strategies for constructing functional tissues/organs and grafts. With the advent of various natural and synthetic biomaterials, tissue engineering has improved a lot and successfully generated relevant grafts and tissues for both basic and clinical applications. Biomaterials are indispensable for tissue engineering and suitability of biomaterial to the desired application actually depends on the type of biomaterial used. Various biomaterials have been successfully used in developing different types of biological products such as vascular grafts, cartilages, bones, liver, lungs, etc. (Lam and Wu 2012; Mammoto and Mammoto 2014; Mazzocchi et al. 2018; Rho et al. 2006). Designing and construction of tissue-engineered organs and grafts is quite challenging and newly emerging methods of development are being reported across the globe. The functionality of tissue-engineered grafts/tissues/organs for the desired application remains a major challenge. Often, the biomaterials provide good scaffold support and allow tissue engineers to construct relevant prototypes. However, sometimes the biomaterial used does not support viability and functionality of cells in vitro. Therefore, various modification methods have been applied to tissue engineering for developing stable, functional, and long-lasting products for appropriate uses (De Mel et al. 2012; Ren et al. 2015). As we have already discussed regarding the different methods of modification in biomaterials, here we will emphasize on the modification/manipulation in biomaterials for their better application in tissueengineered products and their functionality (as shown in Fig. 7.4). The ultimate purpose of surface or bulk modification is to derive a compatible, non-toxic, and easy-to-use biomaterial for a functional tissue product. Functionalized biomaterials

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Fig. 7.4 Applications of biomimetic organoids developed using functionalized biomaterials

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provide excellent ECM support to cells and mimic the tissue microenvironment for better adaptability for the grafted cells (De Mel et al. 2012; Ren et al. 2015). Various studies have shown that surface modification of polymeric biomaterials results in improved characteristics of scaffolds and also a better mimicking of the tissue native microenvironment. By altering the physicochemical attributes of biomaterials via different surface modifications, tissue engineering can possibly be applied for the repair of any tissue types including cardiac, liver, skin, bone, and lungs repair (Table 7.2).

7.3.1

Cardiac Tissue

Tuning of physicochemical properties by using different physicochemical modifications, such as electrochemical polishing, surface patterning and roughening, plasma coating, chemical etching, and passive or covalent layering/coating of the surface, has gained considerable interest in designing permanent biocompatible cardiovascular grafts (De Mel et al. 2012; Moorthi et al. 2017). Such modifications have shown that topography changes result in better physical properties that affect cell behavior, thrombogenicity, and protein adsorption in vascular grafts. For example, some plasma proteins like fibrinogen bind to surface receptors like plasma receptors more efficiently than flat surfaces. Adsorption of some other proteins such as albumin, fibronectin, vitronectin has been shown to be affected by modulating topological changes via physicochemical modifications (Watson 2009). These physicochemical modifications facilitate protein adsorption, cellular interaction, and behavior at blood-graft surface. Over the past decade, advancement in nanotechnology has also contributed to the modification in surface properties of biomaterials and nanoscale modifications such as the development of nanocomposites has emerged as a favorable solution to the development of permanent vascular grafts (Ho et al. 2017; Lalegül-Ülker et al. 2018; Shokraei et al. 2019). Highly viscoelastic, antithrombogenic polyhedral-oligomeric silsesquioxane-poly(carbonate-urea) urethane (POSS-PCU) nanocomposites are being used for designing vascular bypass grafts, stents, heart valves, etc. (Moorthi et al. 2017; Watson 2009). Surface of biomaterials is prone to contamination and therefore, to achieve non-fouling, non-adhesive surfaces, the passivation of materials has been done using polymers like polyethylene glycol (PEG), dextran containing hydrogels, and polyethylene oxide (PEO). Such modifications have clearly demonstrated their versatility, better suitability, and application in cardiovascular grafts construction. For example, grafts of ePTFE coated with polypropylene sulfide (PPS)-PEG showed better cell adhesion and reduced thrombus formation when a comparison of heparinized vs. non-heparinized grafts was done. Similarly, the elastomer poly(1,8-octanediol citrate) (POC) coated luminal surface of ePTFE did not compromised graft compatibility and also minimized thrombosis as compared to control grafts (De Mel et al. 2012). Though such manipulations in surface modifications have enabled tissue engineers to construct stable, compatible, non-fouling permanent vascular grafts, however, emphasis has also been given to

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the strategies where biofunctionalization of biomaterial is done in order to mimic signaling cascades and tissue microenvironment to promote tissue regeneration or replacement with biological functional grafts. Hence, creating biomimetic surface using bioactive molecules is the thrust area of regenerative medicine. In the quest for generating biomimetic grafts, stem cell-based endothelialization of cardiovascular grafts is now being considered for extensive research. In addition, the development of injectable hydrogels encompassing cells and scaffolds is one of the approaches to repair cardiac injury in situ (Cui et al. 2016; Lam and Wu 2012). However, successful translation of tissue engineering to regenerative medicine strictly depends on the availability of suitable biomaterials that can be modified via surface/bulk modification to allow cells to grow in a native environment without rejection, toxicity, and sustained viability. Some of the studies have focused on using cardiac patches of functional scaffolds comprising of cells and functionalized biomaterials [such as bone marrow mesenchymal stem cells/silk fibroin/hyaluronic acid (BMSC/SF/HA) or chitosan– hyaluronan/silk fibroin (chitosan–HA/SF)] to improve the cardiac function after myocardial infarction (Chi et al. 2012; Chi et al. 2013). Use of surface-modified polymers (e.g., polyurethanes and polyesters) and biomolecules (e.g., collagen, laminin, heparin, etc.) has been shown to reduce thrombogenicity via better topological, mechanical, and cellular responses that improve cardiac repair (Lam and Wu 2012; Ren et al. 2015). Such advantageous properties of biofunctionalized polymers are actually attributed to some important peptide sequences that are similar to fibronectin (such as REDV, RGD, GRGDSP, etc.), laminin (IKLLI, LRE, PDSGR, YIGSR, etc.), and collagen type I-derived sequences (such as DGEA, Tenascin-C-derived peptides D5 and D50) (Fittkau et al. 2005; K-I et al. 1989; Wei et al. 2011; Zhang et al. 2010; Zheng et al. 2012). These peptide sequences are well known to form a network with cell surface receptors to facilitate cellular functions like cell adhesion and differentiation. Some recent studies have clearly demonstrated that polymers like PU/PEI are effective in preventing platelet adhesion when their surface is modified with low-molecular weight biomolecules like heparin, hyaluronic acid (HA).

7.3.2

Liver

Liver is a vital organ that maintains body homeostasis by performing multitude of function including bile synthesis and secretion, metabolism of macromolecules and drug metabolism, etc. Critical or severe injuries to liver usually require transplantation of healthy liver tissues to patient. However, due to the scarcity of healthy donors, most of the patients die during the course of treatment. Untiring efforts of various tissue engineers across the globe have paved the way to develop/construct artificial liver tissue/grafts that in near future could be used clinically for liver injury treatment. Besides primary hepatocytes, the use of an appropriate biomaterial has remained the key to mimic various biochemical and structural cues, including ECM, that are required for developing bioartificial liver by facilitating the growth of

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hepatocytes and cell–cell interaction (Bhatia et al. 2014; Ye et al. 2019). However, due to the complex functionality and architecture of liver, it has always remained quite challenging for tissue engineers to develop functional and long-lived artificial liver tissue grafts. With the advent of knowledge in histology and functional biology, researchers have employed complex tissue details of liver to simulate functionality and architectural microenvironment with the help of suitable biomaterials (Chen and Liu 2016). For a heterotypic tissue like liver, it is essential to place all the cell types (hepatocyte, stellate cells, Kupffer cells) within a compatible ECM that may be derived from surface-modified biomaterials (Hosseini et al. 2019; Jain et al. 2014). Both natural and synthetic biomaterials are widely used to create artificial liver tissue (da Silva et al. 2020). However, developing a fully functional liver for transplantation in the lab is still under intense investigation. One of the challenges in developing a functional and viable graft is to bio-mimic the ECM that governs various cellular and physiological cues such as cell adhesion, cell–cell interaction, cell–matrix interaction, and signaling myriads that are important for liver function. The native ECM of liver is comprised of several proteins like collagen, hyaluronans, laminin, fibronectin, and elastin. Collagen type 1 protein predominates as a porous structural scaffold over which matrix deposition takes place, whereas hyaluronans, laminin, fibronectin, vimentin, elastin, and sulfated chondroitin sulfate proteoglycan (CS-PG) and heparin sulfate proteoglycans (HS-PG) are rich in periportal region. The chemistry resulting due to the gradient of these matrix proteins interplays a synergistic role with soluble factors and regulates cell behavior (Arriazu et al. 2014; Baiocchini et al. 2016; da Silva et al. 2020; Kim et al. 2016). Therefore, optimum distribution of ECM components, oxygen gradient, and nutrient transport are some of the prerequisites to achieve the goal of a functional liver tissue. Though naturally obtained biomaterials provide better cell adhesion, biocompatibility, and degradability they lack mechanical strength to provide structural scaffolding. Whereas synthetic biomaterials hold the advantage of better mechanical strength but possess poor biocompatibility and degradability (da Silva et al. 2020). Therefore, sometimes hybrid scaffolds are more suitable for tissue engineering of liver tissue. These provide all the essential components and microenvironment that supports the growth of hepatocytes in vitro (Guan et al. 2017; Kazemnejad 2009). For example, hepatocytes encapsulated in the complex of thiolated heparin and PEG were able to sustain their viability and function for up to 20 days (Kim et al. 2010). Synthetic scaffolds are relatively poor in cell binding sites and therefore, modification of their surface or bulk as a whole is usually required for their optimum use in tissue engineering. Simplest strategy to modify synthetic polymers is to add or integrate cell adhesion or binding ligands/motifs from natural structural polymers like collagen, fibronectin, laminin (Tallawi et al. 2015). In addition to the coating using natural polymers, modification of physicochemical properties by surface modification that enhances the adsorption of growth factors has also improved the recapitulation of appropriate ECM for hepatocytes growth (Nikolova and Chavali 2019). Hence, addition of ECM to 3D scaffolds has been viewed as a successful strategy to liver tissue engineering. For example, heparinized collagen scaffolds have been shown to afford better liver function of differentiated mesenchymal stem cells, as

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they tend to store more glycogen and express late liver-specific markers as compared to the collagen-seeded MSCs (Aleahmad et al. 2017). Hydrogels in both scaffold or injectable forms have been considered as an attractive opportunity to be used in liver tissue engineering. Hydrogels are highly hydrated, easy-to-add soluble growth factors that can mimic natural ECM due to their comparable softness as well as cell-friendly behavior (Mantha et al. 2019; Zhu and Marchant 2011). In one interesting study, alginate hydrogels were used to produce microcavitary platforms by adding cell-laden gelatin microspheres that allow cells to form spheres within the space created by dissolving the gelatin by enzymatic degradation. Moreover, these microspheres were further retrieved from the alginate scaffold by treating with citrate (Hwang et al. 2010; Lau et al. 2012). A combination of gelatin hydrogels and sugar (e.g., galactose)-containing poly (vinyl alcohol) has been shown to serve as a better substrate for liver cell proliferation and improved liver functions like albumin secretion (da Silva et al. 2020). Similarly, surface modification of chitosan nanofibers using galactose ligands has resulted in a scaffold that slows down the degradation while enhancing the functionality of primary hepatocytes (Lalegül-Ülker et al. 2018). Until now, different biomaterials including natural as well as synthetic have been evaluated and used in the construction of liver tissue and biofunctionalization of scaffolds by surface modification using various approaches has improved the outcomes of tissue-engineered liver. Some of these examples are listed in Table 7.1.

7.3.3

Lung

Artificial lung biofabrication is also another important aspect of regenerative medicine to provide considerable solution to the crisis of functional lungs for transplantation in patients. Tissue engineering has made advancement in lung biofabrication by integrating stem cell biology and scaffold chemistry (Prakash et al. 2015). Since a scaffold is an essential component of tissue-engineered grafts, so lung tissue generation also requires it. Therefore, packing of stem cell progenitors with proper scaffolding is being explored extensively with the aim to construct bioengineered lung tissue for clinical applications. Importantly, the emphasis has also been given to the efforts where autologous grafts are being generated using patient’s progenitor cells to avoid immunological graft rejection and allow a better transplantation success rate (Rouchi and Mahdavi-Mazdeh 2015). In addition to bioengineered lungs, researchers are also developing various constructs/scaffolds that find their application in drug delivery, as pulmonary surfactants, lung vasculature repair, etc., in various lung airway diseases (Farré et al. 2018; Prakash et al. 2015). Though the construction of a whole functioning lung has not been achieved so far, considerable success has been gained in the application of biopolymers that allow considerable growth of artificially cultured lungs cells. In this endeavor, both natural and synthetic materials have been exploited for the construction of lung and associated airways tissues (Farré et al. 2018).

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Some of the common biopolymers and their modifications for improving the function, biocompatibility, and longevity of lung tissue/airway tree have been provided in Table 7.1. 3D collagen scaffolds have been used in reconstructing an artificial alveolus by seeding rat alveolar type II epithelial cells on them (Zhang et al. 2011). Synthetic polymers such as polyglycolic acid (Cortiella et al. 2006), polylactic-co-glycolic acid (PLGA) (Mondrinos et al. 2006), and pluronic F-127 (Cortiella et al. 2006) are also considered significant with respect to the architectural framework of lung scaffolds. Either their scaffolds or injectable hydrogels have shown to mimic alveolar structure when seeded with lung progenitor cells. In another study, the longevity of human lung epithelial cells has been shown to enhance by several weeks by using a macroporous matrix of hydroxyethyl methacrylate–alginate–gelatin mixture (Singh et al. 2013). Employing scaffold modification techniques such as electro-spinning has also shown that the electrospun scaffolds prepared from nondegradable polyethylene terephthalate recapitulated the topography of air smooth muscle cells and demonstrated ability of bronchoconstriction (Bridge et al. 2015; Jun et al. 2018). Use of decellularized scaffolds is another promising strategy for constructing whole organs, including lungs (Gilpin and Yang 2017; Tapias and Ott 2014). Here, the ultimate approach is to use decellularized scaffold obtained from the natural tissue to construct a functional, biomimetic, and customized ECM to generate whole lung tissue by replacing with the correct combination of cells such as lung alveolar epithelial cells and endothelial cells. In other strategies, seeding of lung progenitor cells, or stem cells onto a 3D scaffold, along with some soluble factors, has shown cellular differentiation and development of anatomically similar lung tissue (Nonaka et al. 2016; Scheers et al. 2018). In spite of some disadvantages associated with the natural or synthetic polymers, their functionalization/modification is sometimes required to gain optimum functionality, mechanical strength, cell–matrix interaction, cell–cell interaction, adhesion, and cell proliferation. Some of the modified scaffolds such as polymeric elastin containing polyaniline or a mixture of polyglycolic acid/polylactic acid (Fakoya et al. 2018), gel foam sponges (Andrade et al. 2007), and commercial available benzyl esters of hyaluronic acid and crosslinked Hylan (Vindigni et al. 2009) have been used in lung tissue engineering. Other polymers such as polyethylene terephthalate (PET) or its nanocomposites and polyurethane (PU) fibers have also found appropriate in remodeling of tracheal scaffolds (Chiang et al. 2016). Some of the highly tunable synthetic hydrogels like polyvinyl alcohol (PVA), poly(ethylene glycol), and elastomeric poly(glycerol sebacate) (PGS) have also been used in reproducing complex lung architecture. Collagen–glycosaminoglycan complex has been investigated for the production of alveolar-like structure (Chen et al. 2005). These studies clearly demonstrate that simulating the complex ECM of a heterotypic organ like lung requires use of either multiple biomaterials or modified scaffolds of a polymer. Different compositions of scaffolds such as films (e.g., gelatin-modified poly(ε-caprolactone) (Kosmala et al. 2017), electrospun nanofibers (e.g., poly(ε-caprolactone) (PCL)/depolymerized chitosan) (Mahoney et al. 2016), copolymer [e.g., hyaluronic acid-g-poly(2-hydroxyethyl methacrylate) (HEMA)]

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(Radhakumary et al. 2011), and cryogels [e.g., 3D macroporous hydroxyethyl methacrylate–alginate–gelatin (HAG)] (Singh et al. 2013), etc., have been found to support better tissue engineering of lung, tracheal, or alveolar-like structures by mimicking lung ECM (Fakoya et al. 2018).

7.3.4

Bone

Orthopedic injuries often require transplantation of bone grafts to repair bones in patients. Tissue engineering has made significant contribution in developing various bone grafts for clinical applications. However, constructing such grafts is not really straightforward as certain important considerations such as biocompatibility, oesteoconductivity, cell adhesion, viability, vascularization, and cost-effective biomaterials are critical in developing functional bone grafts (Amini et al. 2012). There are two important aspects to repair bone injury, these are: (1) to replace with artificial but biocompatible grafts and (2) to allow patient’s own response to regenerate the bone by using a suitable formulation of progenitor cells with compatible scaffolding. Though allogenic- or auto-grafts offer a number of advantages for bone transplantation, they suffer from severe limitations such as poor donor availability, high cost of transplantation, immunoreactivity/host rejection with allogenic grafts, etc., that have impeded the success rate of transplantation therapy (Greenwald et al. 2001). Therefore, the promising opportunities offered by tissue engineering have opened substantial avenues to overcome these shortcomings in the near future. Bone tissue engineering focuses on delivering functional artificial grafts while exhibiting biocompatibility, minimum or no immune reactivity, and a better mimicking of bone tissue environment by employing suitable metallic biomaterials. Most often, these materials are required to be modified since the surface of these materials helps in biomimicking ECM, supports biochemical signaling, homing of cells, and deposition of calcium (Mitragotri and Lahann 2009; Ponche et al. 2010; Qiu et al. 2014; Suzuki et al. 2006). Surface modifications have enabled researchers to tune surface properties appropriately while designing materials used for making bone grafts (Hu et al. 2019). Various metallic biomaterials, ceramics, and polymers have been used for bone replacement as they provide excellent mechanical strength and post-surface modifications, they have also been found suitable to mimic bone environment without any inflammatory response, which is usually generated from the wearing of biomaterials due to friction within the joints. Various surface fabrication/modification processes to coat external surface of biomaterials such as rapid prototyping, ion beam-assisted deposition, and plasma coating have been utilized for creating biocompatible biomaterials for bone grafts that can be used as a long-term transplant opportunity. Moreover, surface modifications in bone grafts and dental implants have been viewed as substantial strategy to deliver functional and biocompatible grafts for commercial use (Qiu et al. 2014). Various examples of surface modification and derived biomaterials that have been used in bone tissue engineering are shown in Table 7.1.

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Titanium is a widely used metallic biomaterial that finds great application in implants and grafts that does not only offer mechanical strength, but also affords biocompatibility upon certain surface modifications. For example, titanium coated with tantalum via a rapid prototyping method such as laser engineered net shaping (LENS) has shown better osteointegration (Balla et al. 2010). Fabrication of highly adhesive Ti–6Al–4 V substrate has been achieved by coating hydroxyapatite (HA) using ion beam-assisted deposition (Cui et al. 1997). Plasma spray is one of the frequently used techniques for the surface modification of commercial bone implants. For example, polymeric composites externally coated with bioactive HA have been explored in constructing bone implants using the plasma spray technique (Auclair-Daigle et al. 2005). Bone tissue engineering has also provided importance to some other pertinent components of bone grafts such as construction of osteoinductive biomaterials, hybrid biomaterials that can offer biocompatibility, surface adhesion, porosity, as well as biodegradability. Such biomaterials are generated with the aim to elevate the tissue regenerative capacity of patients by adding suitable cell species to replace or cover the damaged or sheared bone using the patient’s own bone formation capacity (Amini et al. 2012).

7.3.5

Dental Implants

Dental implants have shown a remarkable success in the clinic and frequently being used for tooth replacement and other dental repair options. With the advent of materials science and tissue engineering, the successful construction of various bioactive implants has been put forward for clinical applications. Titanium is one of the most suitable biocompatible material that is widely studied and used in dental implants. In addition, the surface modification techniques such as layer-by-layer deposition, hydroprocessing, covalent immobilization, and plasma spray, etc., have enabled researchers to develop biocompatible, bioactive, and osteointegrative/ osteoinductive surfaces to enhance adhesion of cells, differentiation of mesenchymal stem cells to osteogenic lineage, and antibiotic activity within the implants (Jemat et al. 2015; Smeets et al. 2016; Yeo 2020). Dental implants are a classical example of the amalgamation of materials science and tissue engineering that provide reliable and better treatment options to patients in the clinic. Most importantly, the survival of dental implants is quite longer than any other implants (e.g., bones) used clinically (Smeets et al. 2016; Yu et al. 2014; Zhong et al. 2016). Some of the typical surface alterations and their application in dental implants are shown in Table 7.1. Surface-modified titanium and its alloys has shown remarkable success as a transplantable, biocompatible, bioactive, and osteointegrative implant with better cell adhesion, longevity, and biomineralization (Kuroda and Okido 2012; Li et al. 2016; Yu et al. 2014). Surface modification of biomaterials in other soft tissue engineering areas such as skin, kidney, and bladder has also emerged over the past few years. By using various modification techniques, the possibility of the construction of suitable biomaterials fabricated with functional surface allowing cell adhesion, biocompatibility, cell differentiation, functionality, ECM simulation,

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antibacterial activity, etc., has been shown. This allows for the creation of better and functional implants/grafts for clinical uses (Cao et al. 2020; Chang et al. 2013b; Karahaliloğlu et al. 2016; Radwan-Pragłowska et al. 2020).

7.4

Conclusion and Future Perspectives

Over the last few decades, surface modification of biomaterials has gained much importance across the globe as far as development of functional tissues and grafts is concerned. Surface modifications of biomaterials have enabled researchers to tune or produce functionalized surface for designing and constructing novel, functional, stable, and viable tissue-engineered product for various biomedical applications. Ranging from their applications from remodeling of ECM to vascular grafts, dental implants, artificial liver and lungs, bones, cartilages, etc., biomaterials with appropriate surface modifications have enabled homing of different cell varieties including stem cells for constructing novel and better functional grafts. Advancements in biomaterial functionalization has promisingly opened some new avenues in tissue engineering to develop functional, viable and transplantable grafts for repair and regeneration. Various evolving methods of surface modifications such as surface roughening, patterning, covalent modification, layering, etc., have enabled researchers to develop various strategies to pack cells into a desired architecture with different scaffolding materials and varying a degree of their functionalization. Functionalization/modification of biomaterial surface has augmented various biological processes such as osteointegration, vascularization, cell adhesion, differentiation, blood compatibility, ECM remodeling, antifouling surfaces, etc., besides excellent mechanical strength and anti-corrosive properties of implants. Integration of tissue engineering and regenerative medicine with material science, nanotechnology, and basic biology has really been instrumental so far and still evolving to fulfill the unmet need of artificial but functional grafts with biocompatibility and less immune reactivity to offer successful transplantation in near future. Acknowledgement Deepika Arora is a Fulbright-Nehru Postdoctoral Fellow. Pradeep K Sharma is a recipient of ICMR-DHR International Fellowship.

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Trends in Bioactive Biomaterials in Tissue Engineering and Regenerative Medicine G. P. Rajalekshmy and M. R. Rekha

Abstract

Tissue engineering and regenerative medicine holds great potential in repairing damaged cells and tissues and thereby restoring the lost function of any organ. However it is now well established that other than a compatible scaffold appropriate micro-environmental cues also play a significant role for an engineered tissue to deliver its functional characteristics. Recent studies have revealed the significance of bioactive biomaterials and biomolecule delivery along with tissue engineered constructs. Biomolecules mainly consist of growth factors, hormones, drugs, proteins, cells, etc., which can be either loaded onto the scaffolds or delivery vehicles. These bioactive molecules have defined roles which include the regulation of cellular differentiation, proliferation, and migration. In this chapter, we describe the various modes of biomolecule delivery and its relevance in tissue engineering/regenerative medicine. Keywords

Bioactive · Scaffolds · Regenerative · Biomolecule · Drug delivery

Abbreviations ASC BMP-2 CGN

Adipose derived stem cells Bone morphogenetic protein-2 Cartilage oligomeric protein

G. P. Rajalekshmy · M. R. Rekha (*) Division of Biosurface Technology, Biomedical Technology Wing, Sree Chitra Tirunal Institute for Medical Sciences & Technology, Thiruvananthapuram, Kerala, India e-mail: [email protected] # The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 B. Bhaskar et al. (eds.), Biomaterials in Tissue Engineering and Regenerative Medicine, https://doi.org/10.1007/978-981-16-0002-9_8

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COMP DCF DFO ELP ePTFE ECM FGF GDNF GAG GNPs GMCSF HFP HAp HIF 1-α IFN γ LbL LCST MMP MeTro NGF PCL PU PDLLA PEGMA PEG PGA PHB/CTS PLA PLGA PMAM SCF SDF1 hSMSC TiO2 TGF β3 VEGF

8.1

Citalopram-loaded gelatin nanocarriers Deferoxamine Diclofenac Elastin like polypeptides Expanded polytetrafluoroethylene Extra cellular matrix Fibroblast growth factor Glial cell line-derived neurotrophic factor Glycosaminoglycan Gold nanoparticles Granulocyte-macrophage colony-stimulating factor 1,1,1,3,3,3-hexafluoro-2-propanol Hydroxyapatite Hypoxia inducible factor 1-α Interferon γ Layer-by-layer Lower critical solution temperature Matrix metalloproteases Methacryloyl-substituted tropoelastin Nerve growth factor Poly (ε-caprolactone) Poly urethane Poly(D, L-lactic acid) Poly(ethylene glycol) methacrylate Poly(ethylene glycol) Poly(glycolic acid) Poly(hydroxybutyrate)/chitosan Poly(lactic acid) Poly(lactic acid-co-glycolic acid) Polyamidoamine Stem cell factor Stromal derived factor 1 Synovium-derived mesenchymal stem cells Titanium oxide Transforming growth factor β3 Vascular endothelial growth factor

Tissue Engineering

Tissue and organ failures are serious medical conditions for which organ transplantation, surgical repair, and drug therapy are the recommended treatment options. Tissue engineering is an emerging multidisciplinary field, which aims to regenerate

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damaged tissues instead of replacing them. This is a combination of the principles and technologies from the material, engineering, and biosciences to develop functional substitutes for damaged tissues and organs. In this technique cells are grown on highly porous scaffold biomaterials which act as a template and guide the growth of the new tissue as well (Obregón et al. 2017). To elaborate, tissue engineering relies on various components including biocompatible scaffolds with appropriate cells grown in it, loaded with required biomolecules which include drugs, growth factors, genes, etc., in a combination. The scaffolds are designed to influence the physical, chemical, and biological environment of a cell population. There are mainly two approaches to develop engineered tissue. In the first approach, the scaffold itself can be used as a cell support device upon which cells are seeded in vitro; cells are then encouraged to lay down matrix to produce the foundation of tissue for transplantation. In the second one, scaffold is used as a growth factor or drug delivery device to guide the growth of new tissue (Howard et al. 2008). The biomolecule loaded scaffold upon implantation recruits cells from the body to the scaffold, promotes cell growth, and ultimately forms tissue. This combination of cells, signals, and scaffold is often referred to as a tissue engineering triad. As this chapter deals with bioactive scaffolds that promote tissue engineering we will be focusing on the biomaterials that are designed to elicit cell favorable cues that promote growth and differentiation, biomolecule loaded scaffolds for delivering appropriate growth factors, proteins/gene, etc.

8.2

Bioactive Scaffolds

The development of bioactive materials is the most important advancement in the field of biomaterials. Bioactivity can be incorporated into the material through biological recognition by incorporation of bioactive molecules, adhesion sites, and cleavage sites for enzymes. Materials can be modified to transform under external stimuli such as light, temperature, pH, or chemical composition. There are three strategies used to develop biomimetic materials (Glaser and Viney 2013) • Incorporation and release of bioactive components. • Surface modification of the biomaterial with specific binding motifs. • Nanoscale patterning of the material. The first generation biomaterials based therapies were limited to its composition and mechanical properties of the tissue to be replaced. But these approaches do not support the tissue microenvironment. The new biomaterials are developed based on the principles of biomimicry. The bioactive biomaterials will mimic the natural extracellular matrix composition and architecture and provide necessary bioactive cues for the overall control of cellular functions (Fratzl 2007). Biomimetic biodegradability is generally required for a tissue engineering scaffold. The degradation rate should match with the neo tissue formation. Ideally linear

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poly esters such as poly (lactic acid) (PLA), poly (glycolic acid) (PGA), and their copolymers poly (lactic acid-co-glycolic acid) (PLGA) are commonly used. They are biocompatible and possess tunable biodegradability. Poly (ethylene glycol) is widely used as a scaffold for soft tissue engineering application, since it possesses similar mechanical nature of the soft tissues. But they do not have appropriate biodegradability. So the enzymatic biodegradability is used to synthesize biomimetic material. PEG copolymers are tethered with oligopeptide sequence specific for the cleavage by the enzymes matrix protease. Such hydrogels can be specifically degraded by cells secreting matrix metalloproteinases (MMPs) such as collagenase. By tailoring gel degradation properties, gels with optimal properties can be fabricated to support initially compressive load while simultaneously supporting the formation of neotissue (Bryant et al. 2004). Tissues differ in their elastic properties, so that materials with biomimetic mechanical properties are needed. Engineering such tissues has been a continuous effort especially for cardiac muscles, blood vessels, and heart valves. Poly (ε-caprolactone) (PCL) and polyurethanes (PU) are typically used for such applications. PCL is semi crystalline polymer having low glass transition temperature and highly elastic at room temperature. The major disadvantage of PU is the involvement of toxic precursors during synthesis (Ghaee et al. 2019). Elastin is another ECM protein having good mechanical properties such as extensibility and elastic recoil. It is synthesized as soluble precursor tropoelastin and converted to insoluble elastin. Elastin plays significant role in functionality of many tissues such as blood vessels, heart valves, and skin. Apart from providing mechanical strength, they have role in signaling process also. Different elastin based scaffolds are synthesized, namely electrospun scaffolds, elastin based hydrogels, elastic sealants, and synthetic elastin like materials (Daamen et al. 2007). To synthesize electrospinning elastin scaffolds, tropoelastin is dissolved in low boiling point solvents 1,1,1,3,3,3-hexafluoro-2-propanol (HFP). After fibered position, chemical crosslinking can be done using the fumes from a 25% aqueous solution of glutaraldehyde (Rnjak-Kovacina et al. 2011). By modulating electrospinning parameters, mechanical properties, pore size, and fiber size can be controlled. To develop as a wound healing scaffold, dermal fibroblast can be grown within the scaffold, which subsequently secretes collagen fibers and fibronectin (Nivison-Smith and Weiss 2011). Elastin based synthetic hydrogels are developed which can withstand great amount of water incorporation while maintaining the integrity. This can be used as wound healing scaffold since it provides rapid restoration of damaged tissues. Lin et. al. synthesized electrodeposited hydrogels made up of silk-tropoelastin alloys. Silktropoelastin alloys are a combination of recombinant human tropoelastin and regenerated Bombyx mori silk fibroin. They possess tunable physical and biological properties. Electro deposition of alloys helps in the assembling of gels with spatial and temporal controllability. Silk and tropoelastin carry opposite charges and the overall electric charges modulated by enzymatic coupling between silk and tropoelastin. Electrodeposited gels prepared from silk-tropoelastin protein alloys provide novel versatile tool for coating material in nanoparticle based drug delivery

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systems. They possess enhanced tissue adhesion and desired cellular attachment and growth, with programmable gel matrix degradation (Lin et al. 2015). Surgical glues are one of the emerging biomedical tools which replace suturing and stapling. They reconnect ruptured tissues having low adhesion, inappropriate mechanical strength, and poor performance in biological environments. Annabi et al developed highly elastic methacryloyl-substituted tropoelastin (MeTro) surgical sealant. The recombinant tropoelastin was introduced by photocrosslinking, which provides tunable adhesion properties. In vivo experiment carried out in incisional model of artery sealing in rats showed MeTro sealant treated groups had normal breathing and lung function. The material also showed tunable degradation in vitro by the action of matrix metalloproteinase-2 (MMP-2). Lower concentration of the hydrogel had faster degradation by in vivo experiments (Annabi et al. 2017). Synthetic elastin like polypeptides (ELPs) have found promising applications in tissue engineering. The key interesting features that make them promising are its non-immunogenic and biodegradable nature. Since ELP’s are synthesized, either chemically or genetically, their amino acid sequence, molecular weight, etc., can be designed, also specific recognition and functional motifs can be introduced. Elastin like polymers (ELPs) are obtained by synthetic strategies that require chemical methods. The viscoelastic properties of ELPs can be modified by crosslinking approaches. They can be synthesized in different forms such as gels, films, foams, and fibers. Cell recognition sites like RGD or REDV can be incorporated into ELPs to regulate cell responses by mimicking native protein substrate. Different bioactivities and specific functionalities can be incorporated at precise locations to modulate biocompatibility and mechanical properties (Nettles et al. 2010). Yeboah et al. described the development of recombinant fusion protein comprised of SDF1 and an elastin like peptide that confers the ability to self-assemble into nanoparticles. Topical application of SDF1 promotes neovascularization and rapid re-epithelialization. SDF1-ELPs can act as drug depots that supply biomolecules over extended period of time. The particles are also stable in the presence of elastase, so can be used for the treatment of chronic wounds. Apart from wound healing applications, SDF1-ELP fusion protein nanoparticles can be used for other applications such as myocardial infarction, where SDF recruit stem cells to promote local tissue regeneration (Yeboah et al. 2016).

8.3

Incorporation of Bioactive Components

8.3.1

Bioactivity by Incorporation of Adhesion Sites

Extensive research has been performed to incorporate adhesion promoting sites within the biomaterials. Cell adhesion is a dynamic process occurring through the interaction of cell surface molecules with appropriate ligand. They involve in signal transduction for gene expression, cytoskeletal dynamics, and growth regulations. The surface properties of the material determine the cell adhesive interactions.

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Biomaterial surface can be modified by blending, surface deposition, or electrostatic attachment. During surface deposition process, adhesion molecules can be bound to the biomaterials by weak forces such as van der Waals, hydrogen bonding, or electrostatic interaction. Adhesion molecules can be blended with biomaterials to create composites. This results in uniform distribution of molecules in the biomaterial matrix. Blend composites are produced as thin films or 3D polymeric constructs. Blending can also increase the encapsulation efficiency of cells. Electrostatic attachment can be of two types, such as layer-by-layer (LbL) assembly and electrochemical polymerization. Using layer-by-layer technique, alternative layers of polycationic and polyanionic materials are deposited and self-assembled by electrostatic interaction to produce nanoscale coating (Rao and Winter 2009). Lotfi et al studied the influence of different poly electrolyte multilayer films on gingival fibroblast cell responses. Multilayered films are made by using layer-by-layer technique based on alternating oppositely charged polyelectrolytes on glass probes. These can be used for orthopedic surgery, ophthalmology, urology, aesthetic surgery, and other domains (Lotfi et al. 2013). During electrochemical polymerization, electrically conducting charged crystalline or semi crystalline polymers chains are doped with opposite charge. This technique is employed for electrical prostheses. Natural extracellular matrix provides platform for cellular adhesion and activation of signaling pathways. However ECM based biomaterials showed limitations such as poor mechanical strength and immunogenic reactions. But reduced immune reaction is observed with synthetic polymers; however, they lack biological activity for cell adhesion and function. Integrin receptors are responsible for direct signaling through interaction with binding proteins. The active conformation of integrin receptors increases the affinity for ECM ligands. Mechanical properties and chemical environment regulate integrin activation, adhesion, and signaling. To improve cell material interaction, synthetic bioadhesive motifs are developed for material surface modification. Short linear sequences of amino acid derived from proteins are incorporated into the biomaterial. Most commonly used sequence includes fibronectin (e.g., RGD, KQAGDV, REDV, and PHSRN), laminin (e.g., IKLLI, LRE, LRGDN, PDGSR, IKVAV, LGTIPG, and YIGSR), collagen (e.g., DGEA, GFOGER), and elastin (e.g., VAPG). RGD (Arginine–Glycine–Aspartate) sequence is a ubiquitous receptor adhesion motif found in most ECM proteins. Enhanced cell spreading and attachment was observed with RED modified biomaterials (Ghaee et al. 2019). There are many researches carried out to incorporate adhesive polypeptide chains to the biomaterial surface. RGD sequence from fibronectin can be immobilized by amino terminal primary amine via glycyl spacer. To induce adhesion, spreading, and cytoskeletal organization, approximately 105 copies of RGD sequences per cell are required. But in the case of bioactive pentapeptide YIGSR from laminin, immobilized single orientation could produce different biological activities. The cell adhesion strength depends on the surface concentration of adhesion ligands and cell migration rate depends on strength of cell adhesion. For tissue engineering applications, degradable biomaterials are used, so the surface modifications are not effective due to rapid disintegration. Hence new materials were developed so as to

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display ligand continuously on the surface as it degraded and remodeled. A copolymer of poly (lactic acid-co-lysine) was synthesized by Quirk et al for suture and surgical staple applications. The ε-amino groups on the lysyl comonomer provide sites for ligand grafting and the peptides on the surface move through the material during degradation of poly(lactic acid) (Quirk et al. 2001). In another study, poly (lactic acid) was surface modified with immobilized RGD peptide linked with polyL-lysine. The material showed good spreading of cells (Ardjomandi et al. 2012). Ospina et al synthesized collagen-functionalized poly (lactic acid) by grafting and prepared electrospun scaffolds to grow cells in vitro. The bioactivity was compared with scaffolds made up of PLA-collagen blend and observed that, there was fourfold increase in cell adhesion than PLA-collagen blend (Ospina-Orejarena et al. 2016).

8.3.2

Nanopatterning

Cellular recognition cues can be modified by changing nano topographical characteristics of the presented ligand. Topography can be altered by surface roughening technique. Mechanical and chemical etching of titanium surface can improve the cell adhesion properties of the material. Nanopatterning allows the controlled application of binding motifs similar to that of physiological spacing. Block copolymer nanolithography is an effective technique for spatially controlling cell receptor– ligand interaction. This technique utilizes spherical micelles filled with nanometersized particles that form distinct hexagonal patterns when exposed to reactive gas plasma. Functionalizing the particles with bioadhesive peptides renders a geometrically patterned surface for cell adhesion (Maheshwari et al. 2000). Zamuner et al. designed scaffolds for bone tissue regeneration to promote cell adhesion and proliferation. Selective functionalization of glass and titanium surfaces with adhesive peptides mapped on sequence of human vitronectin helped to selectively increase osteoblast attachment and adhesion. However, the peptides got completely cleaved after 5 h. To overcome the enzymatic degradation, they synthesized a retro-inverted peptide, which was completely stable within the media. A grafted retro-inverted peptide induced cell focal adhesions and filopodia and increased the osteoblast adhesion and gene expression (Zamuner et al. 2017).

8.3.3

Bioactivity by Incorporation of Growth Factors

Growth factors are powerful regulators of cellular behaviors. The biological activity of growth factor depends upon its presentation to the cells in space and over time. The extracellular matrix plays significant role in storage and release of growth factors in spatiotemporal manner. The biomaterials which mimic the function of the natural ECM are necessary for growth factor delivery. Immobilization of growth factors to the biomaterials is important for artificial organs, tissue engineering scaffolds, and regenerative medicine applications. Immobilization is carried out to

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enrich the local concentration via direct contact with cell membranes, so that significant effect will be obtained at lower concentration. Of the growth factors, VEGF has a significant role in tissue regeneration and is widely explored for suitable delivery so as to sustain its biological activity. Free VEGF gets cleared and degraded fast resulting in poor therapeutic efficacy. Maia et al. demonstrated prolonged biological activity of the VEGF functionalized with dextran compared to that of free molecule. VEGF-functionalized dextran was cross-linked with adipic acid dihydrazide to form a degradable gel and incorporated into endothelial cell containing fibrin gel to modulate the activity. This construct might be effective for enhancing pro angiogenic activity of VEGF in ischemic tissues and improve the biological activity of vascular cells (Maia et al. 2012). In another study by Nur et al., tried to immobilize Fibroblast growth factor (FGF)-2 on amine modified nanofibril-deposited surfaces. Because of the pivotal role of FGFs in cellular proliferation and developmental pathways, a number of studies have attempted to modulate the micro environment to stabilize growth factor for tissue engineering applications. The activation of the FGF receptor requires the formation of ternary complex of growth factors (FGFs), FGFRs, and heparan sulfate proteoglycans. Heparan sulfate proteoglycans help to stabilize FGF by inhibiting proteolytic degradation (Nur-E-Kamal et al. 2008). Ham et al. developed N terminal modification for immobilization of interferon γ (IFN-γ). It was then coupled with azide group and tethered to dibenzocyclooctyne modified chitosan and hyaluronan via “click” chemistry. Tethered IFN-γ could able to induce neuronal differentiation of neuronal stem cells (Ham et al. 2017). Biotin– streptavidin binding was used for immobilization of granulocyte macrophage colony-stimulating factor (GMCSF), SCF, thrombopoietin, and interleukin-3. They showed sustained activity over a week in bioreactor and showed long-term growth response (Worrallo et al. 2017). One of the key challenges in tissue engineering is overcoming cell death within the scaffolds due to limited diffusion of oxygen and nutrients. An immobilized gradient of growth factors from periphery to the center of a porous scaffold guides the cells to overcome the necrotic core. Quantitative investigation of immobilized GFs can be analyzed by gradient formation techniques. Odedra et al. demonstrated VEGF gradient on porous collagen scaffolds. The first step was the activation of VEGF-165 with 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide/sulfo N-hydroxysuccinimide and then applied to the center of the scaffold to create the gradient. The gradient was formed in radial direction across the scaffold. More endothelial cells are guided into the scaffold, demonstrating high cell density at the center. So, such a VEGF gradient promoted cell migration, not the cell proliferation. This scaffold shows promising characteristics for in vivo tissue engineering studies (Odedra et al. 2011). Dual growth factor delivery using biocompatible scaffolds is another strategy used in tissue engineering. In nerve tissue engineering, delivery of nerve growth factor (NGF) and glial cell line-derived neurotrophic factor (GDNF) can be delivered simultaneously to promote peripheral nerve repair. NGF and GDNF were encapsulated into poly(D,L-lactic acid) (PDLLA) and poly

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(lactic-co-glycolic acid) (PLGA) nanofibers via emulsion electrospinning. Sustained release of both growth factors was also achieved. Another strategy to enhance the bioactivity of immobilized growth factors is the co-immobilization with adhesion factors. The interaction of integrin and growth factor receptors is very important in the designing of growth factor immobilization. Ettelt et al. developed biotin–streptavidin coupled system to modify TiO2 with immobilized BMP-2 to promote bone and soft tissue growth. Co-immobilization yielded increased osteoblast activation compared to either protein alone. So combined tethering of different ECM proteins will facilitate the cell specific reaction of implants (Ettelt et al. 2018).

8.3.4

Bioactivity by Physiochemical Interactions

Biological activities can be brought onto biomaterial surface by physico-chemical interactions. Some of the important biological activities such as anticoagulant activity can be introduced into the biomaterial by charge modification. By providing appropriate degree of sulphonation polyurethane exhibits heparin like activity, by binding to antithrombin III and thrombin, to catalyze complex formation between these two proteins. Heparin can also be used for other biological activities such as binding to growth factors or interfering with a growth factor’s binding to its receptor.

8.3.5

Bioactivity by Material Transformation

Biomaterials can express biological activity via ability to transform the material properties in situ, i.e. at the site of implantation. This property is more relevant in sealants and cell-tissue barriers. A water soluble macromer was developed based on poly(ethylene glycol) central block with oligo(lactic acid) flanking block and terminal acrylates. Central block provides water solubility, flanking oligo esters determine the degradability, and terminal group regulates the polymerizability. These soluble materials rapidly transformed into elastic hydrogel on exposure to light in the presence of photoinitiators. This material can be used as a barrier and depot for local drug delivery to prevent formation of scar tissue adhesion. Another mode of material transformation includes gel swelling or collapse in response to temperature changes. This property used with polymers displaying lower critical solution behavior or upper critical solution behavior. Such materials are used as drug delivery depots. Biomaterial monolayers for therapeutics have been developed. They bind to the tissue surface based on electrostatic interactions. Amine reactive poly(ethylene glycol) was chemically grafted with lysine residues, which inhibits cell adhesive interactions after surgical tissue damage (Hubbell 1999). The development of smart biomaterial requires strict control over the material surface properties. The response of the cultured cells depends on the surface characteristics of the material. Cells are arranged in distinct pattern during

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development. It would be beneficial if patterned environment is provided in vitro for regulating cell behavior.

8.4

Bioactive Inorganic Biomaterials for Tissue Engineering

Inorganic biomaterials play attractive role in regenerative medicine due to their tunable properties. The properties can be biophysical or biochemical cues that could able to direct tissue regeneration. They regulate cellular responses including cell–cell and cell–matrix interactions. Ions released from the mineral based biomaterials play significant role in tissue specific functions. The released bioactive ions can induce phenotypic changes in cells and also modulate the immune micro environment for tissue healing and regeneration. Biomolecules can be easily sequestered and released from mineral based biomaterials. Mineral based micro or nanoparticles can be easily ingested by the cells and provide cues by the release of ionic dissolution products. These particles can be easily combined with various polymeric scaffolds for controlled delivery. The biophysical properties such as shape, size, surface to volume ratio, topography, stiffness, and charge of the biomaterials can be modulated for regenerative medicine applications. The stiff biomaterials shown to facilitate osteogenic differentiation, while soft biomaterials can induce chondrogenic differentiation (Engler et al. 2006). Inorganic biomaterials including monolithic and polymer composites can respond to cellular signals and interact with endogenous immune system and stem cells to stimulate in situ tissue regeneration. Da Tren Chou et al. developed 3D printed iron-manganese biodegradable scaffold for bone regeneration. Iron based alloys have high strength and slow corrosion rate. The scaffold exhibited similar tensile and mechanical properties of natural bone and had a porosity of 36.3%. The cell viability studies carried out in pre-osteoblast cells showed good cytocompatibility and cell infiltration into the pores was also observed. The primary studies revealed that Fe-Mn alloy is a promising material for craniofacial biomaterial applications (Chou et al. 2013). In another study by Schussler et al. developed scaffold that could promote normal fracture healing by endochondral ossification. They have used vanadyl acetyl acetonate combined with fibrous porous scaffold consisting of polycaprolactone. The differentiation of human mesenchymal stem cells was evaluated using three different induction media such as osteogenic, chondrogenic, and osteo/chondrogenic media, which mimic endochondral ossification. The controlled release of vanadium was observed for 28 days. Almost 1000-fold increase in VEGF gene expression was noticed, which indicates promotion of angiogenesis by vanadium ions (Schussler et al. 2017).

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Injectable Biomaterials

Injectable biomaterials are extesnsively applied in the field of tissue engineering. They possess specific characteristic features such as minimally invasive nature and in situ formation of scaffold. The areas where injectable biomaterials applied include bone, cartilage, cardiovascular tissue, filling of bone defects, and skin tissues (Kona et al. 2011). The gelation kinetics of the injectable biomaterials is very critical for tissue engineering application. Thermal gelation is very faster than pH or ionic gelation process. In situ polymerization can occur by the action of some crosslinking chemicals. Thermo gelling polymers are widely used in drug delivery and tissue engineering applications. They can undergo entropically driven phase separation above their lower critical solution temperature (LCST). They show rapid gelation following injection, as their LCST is below body temperature and forms well hydrated networks. The presence of hydrophobic groups such as methyl, ethyl, and propyl are characteristics of temperature sensitive polymers. The pH responding polymer network possesses acidic or basic group that accepts or releases protons in response to pH changes. Electrically responding hydrogels are made up of polyelectrolytes. They are able to swell and shrink according to the electric field. The rate of polymerization should be sufficiently quick within adequate period of time (Kretlow et al. 2009).

8.6

Bioactive Scaffolds: Tissue Engineering Applications

8.6.1

Neural Tissue Engineering

Enormous progress has been made in the field of neural tissue engineering for regulating both central nervous system and peripheral nervous system. Polymers are largely used due to their high versatility than metals and ceramics. The physical, chemical, and biological properties of polymers can be modulated depending on the application. They can be used as drug delivery vehicle, hydrogels, nerve conduits, and scaffolds. Polymeric structures support growing neuritis and regulate biological cues for axonal growth (Boni et al. 2018). Natural polymers are highly beneficial for neural tissue engineering due to their biocompatibility and degradation kinetics combined with chemically tunable properties. The components of ECM like collagen, natural polysaccharides such as alginate and chitosan, and other polymers derived from insects like silk are widely used for synthesis of neural scaffolds. Collagen scaffolds can repair small nerve gap (5 mm) and physiologically similar to graft repair. Collagen conduits can also act as internal neural fillers, thus increasing the quality of peripheral nerve regeneration over longer gaps. Collagen hydrogels can be used for the regeneration of sciatic nerve gap of 15 mm. In a study conducted by Hiroshi et al described collagen-PGA tube as promising biomaterial for nerve conduits. In this study, they developed a hybrid artificial nerve conduits by filling PGA with dedifferentiated fat cells and

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applied to a rat nerve defect model and the facial nerve regenerative ability of the conduits (Fujimaki et al. 2019). Gelatin is a denatured protein product derived from collagen and is less immunogenic in nature. Gelatin as such and also in combination with other polymers, is widely explored for various biomedical applications including neural tissue engineering. By electrospinning method, the optimization and manipulation of mechanical, biological, and kinetic properties of the nerve conduits are possible. Gelatin is often cross-linked with genipin which enhances both biocompatibility and stability of the cross-linked product. They also provide biological cues for the differentiation of seeded cells. The in vitro biological assay of the conditioned electrospun scaffolds using rat allogeneic mesenchymal stromal cells confirmed its biocompatibility and differentiating potential using rat brain extracellular matrix (Baiguera et al. 2014). Gelatin nanoparticles have used to enhance the biocompatibility of neural tissue scaffolds. Naseri et al developed poly lactic acid/cellulose acetate scaffold loaded with gelatin nanoparticles. Here poly lactic acid was fabricated as electrospun core and cellulose acetate as the fibril shell. The scaffold was then coated with citalopramloaded gelatin nanocarriers (CGNs). The cytocompatibility was evaluated using Schwann cells. Then the scaffold was developed into a nerve guidance conduit and surgically implanted into sciatic nerve defect in Wistar rats and full functional recovery of the injured nerve was observed (Naseri-Nosar et al. 2017). Hyaluronic acid hydrogel enhances the survival rates and proliferation of neural precursors and also supports neurite outgrowth, differentiation, and proliferation. They are promising materials for peripheral nerve regeneration therapies. Tarus et al. developed a hyaluronic acid based extracellular matrix with tunable stiffness and density tethered with cell adhesive RGD peptide, which mimics the mechanical properties of the brain matrix. Cells seeded on the surface of the hydrogel experienced an optimum neuronal outgrowth as a function of ligand density neurites. They progressed within the gels in vertical direction depending on the structure of hydrogel (Tarus et al. 2016). Alginate gels are also promising materials for nerve regeneration. The leading application of the alginate gel is in the treatment of spinal cord injury, where it has been observed that, it is successful in regenerating small nerve gaps ranging from 2 to 4 mm. Wen Hen et al. developed long-term three dimensional culture system using integrin ligand modified alginate hydrogel that encapsulates neural progenitor cells. The porosity of the hydrogel was adjusted by optimizing alginate concentration (Wen et al. 2019a). Chitosan hydrogels have been successfully used in neural tissue engineering. They promote cell adhesion, cell survival, and neurite outgrowth. Aligned nanofibrous scaffolds in nanoscale dimensions are suitable for the alignment of nerve tissues. The aligned forms developed by Afarin et al. by combining poly (hydroxybutyrate)/chitosan (PHB/CTS) nanocomposites showed suitable hydrophilicity, mechanical properties, and morphology for nerve tissue engineering applications. Amino ethyl methacrylate derivatized, photocrosslinked chitosan hydrogels were synthesized. They facilitated enhanced neurite differentiation from cortical neurons and also help in enhanced extension of dorsal root ganglia. They

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also facilitated the differentiation into tubulin positive neurons and astrocytes (Karimi et al. 2018).

8.6.2

Vascular Tissue Engineering

Synthetic polymers are widely used for vascular engineering. Various tissue engineering strategies are raised to rectify the flaws of synthetic polymers and provide bioactive biomaterials for in situ arterial regeneration. The material should have similar mechanical properties of the native vasculature and should promote cell growth, facilitate matrix reproduction, and inhibit thrombogenicity. Currently expanded polytetrafluoroethylene (ePTFE) and Dacron (polyethylene terephthalate) are the widely used synthetic materials for graft synthesis. Dacron is a very resistant and biostable thermoplastic polyester. They are used for large diameter vascular prostheses, especially for arterial sutures and construction of valve rings. They prevent hydrolysis of the graft due to highly crystalline and hydrophobic nature. Polyurethane grafts have been used for the last 40 years, which are ideal for bypass procedures. But they showed degradative behavior causing aneurysm. To enhance patency of the grafts, several methods are developed to enhance the patency rates, such as linking heparin to graft surface helps to reduce the thrombogenic activity. In another method, coating of carbon on the luminal surface helps to reduce the electronegativity, thus reduces thrombus formation. Coating with fibrin glue can also able to improve endothelialization (Adipurnama et al. 2017). The major disadvantage of synthetic grafts is the rejection within few months by the immune system if the diameter is smaller than 6 mm. To overcome such situations, tissue engineering modalities have been increasingly adopted. Materials with adequate mechanical strength to withstand the long-term hemodynamic stresses are developed. They can able to withstand infection and provide satisfactory graft healing (Carrabba and Madeddu 2018). In a study reported by Zheng et al., they fabricated an electrospun PCL vascular graft functionalized with RGD sequence bearing molecule (Nap-FFGRGD) along with the hydrophobic moiety naphthalene. The hydrophobic moiety enables selfassembling in such a way that the RGD sequences forms as a coating on the surface. The healing nature was evaluated in rabbit carotid arteries for 2 and 4 weeks. They noticed that RGD modified PCL grafts were remained patent compared to non-modified grafts. This suggested that RGD modification significantly improved the hemocompatibility of the PCL graft. There was also threefold increase in endothelial coverage of the PCL-RGD grafts than unmodified graft (Zheng et al. 2012). Silk fibroin has got lot of attraction in the field of vascular tissue engineering. Silk based grafts are used as flow diverting devices and stents. In a study conducted by Uden et al. developed three layered silk fibroin/poly urethane vascular graft by electrospinning method, can be applied as long-term hemodialysis vascular access. Polyurethane provides mechanical properties for the graft. This proposed approach may also represent a step forward to in situ engineered hybrid vascular access with

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potentialities for vein-graft anastomosis stability, early cannulation, and biointegration (Van Uden et al. 2019). Another sulfated silk fibroin was developed by Liu et al. by reaction with chlorosulphonic acid in pyridine and scaffold developed by electrospinning method. They observed that, sulfated silk fibroin has anticoagulant activity and both endothelial cells and smooth muscle cells strongly attached and proliferated on their surfaces (Liu et al. 2015).

8.6.3

Cardiac Tissue Engineering

The main aim of the cardiac tissue engineering is the development of cardiac graft, heart tissue substitutes that can be efficiently implanted in the organism to regenerate the tissue, to provide fully functional heart. The regeneration of myocardium has been increasingly explored, without causing any side effects such as immunogenicity. The major hurdles of the cardiac tissue engineering involve scaffold material selection, scaffold material production, cell selection, and in vitro cell culture. Synthetic polymers involved in myocardial tissue engineering include polyglycolic acid (PGA), polylactic-l-acid (PLLA), polylactic glycolic acid (PLGA), and polyurethane. Polymers with adjustable degradation rate, good porosity, biocompatibility, and elastomeric properties are selected which favor the tissue contraction inherent to the cardiac function. Natural polymers which resemble extracellular matrix are selected that can hold the cells together similar to native tissue. Collagen types I and III and fibrin are extensively investigated for cardiac tissue engineering because of their natural interaction with the cells. Pattern of the biomaterial also significantly influences the cell growth. McDevitt et al. demonstrated that cardiomyocyte cultured on polyurethane films with printed pattern allows the two dimensional alignment of cells and presented with contractile responses (McDevitt et al. 2003). Engineered heart tissue can be constructed using collagen type 1 and extra cellular matrix proteins. Zimmerman et al. developed myocardial infarction rat model using this technique. The developed heart tissue showed differentiated myocytes, formed patent vasculature that was anastomosed with recipient blood vessels. After implantation, the construct showed the formation of capillary networks. Electrical and mechanical stimulations played critical role in cell maturation and tissue coordination (Zimmermann and Eschenhagen 2007). Growth factors can also be incorporated into the bioreactor system depending on its design. For generation of large constructs, cell sheet technology and decellularized matrices can be employed. The decellularized whole heart can be seeded with different cardiovascular lineages and placed in perfusion bioreactors with appropriate mechanical and electric stimulations, provides a great promise for the development of new tissues. The successful development of engineered tissue construct is the optimum interaction between different cell populations. This creates a signaling network which is essential for communication of different cell compartments and mechanotransduction.

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Biomaterial Based Stem Cell Therapy in Regenerative Medicine

Stem cells have wide range of application in regenerative therapy, due to their ability to regenerate damaged or diseased tissues. The major limitations include poor cell persistence and engraftment upon cell transplantation. Biomaterials provide artificial extra cellular environment, which helps to modulate the stem cell behavior. Naturally derived biomaterials are widely accepted for regenerative medicine. Kim et al. developed TGF-β1 incorporated chitosan collagen hydrogel that induced chondrogenic differentiation of human synovium-derived stem cells. TGF-β1 delivered from the hydrogel in a controlled manner. The burst release was reduced by the conjugation of TGF-β1 to MeGC hydrogel. Collagen impregnation and TGF-β1 delivery significantly enhanced cellular aggregation, followed by deposition of ECM by the encapsulated synovium-derived mesenchymal stem cells (hSMSCs) (Kim et al. 2015). Gelatin is one of the promising biomaterials for stem cell delivery, since it supports cell growth and differentiation. Dong et al. designed a hydrogel based on thiolated gelatin containing multifunctional PEG for improving stem cell delivery. This injectable hydrogel was designed with highly tunable properties such as mechanical properties, biodegradability and cellular responses can be finely controlled by changing the hydrogel formulation and cell seeding density. Spontaneous gelation was occurred within 2 min. Murine adipose-derived stem cells (ASCs) were encapsulated into the hydrogel. In vivo studies showed that in situ formed hydrogels could able to improve angiogenesis and accelerate wound closure. The ability of the hydrogel to regulate stem cell behavior in 3D culture can be used for regenerative therapeutic applications (Dong et al. 2017). Bioceramic materials can also be used for stem cell delivery. Among bioceramics, zirconia-based ceramic has attracted great attention due to higher mechanical strength, fracture toughness, and biocompatibility. Zirconia implants can support stem cell differentiation. In a study conducted by Kitagawa et al. designed a culture system for hMSC differentiation. The microwells were composed of zirconia as the ceramic substratum facilitated the adhesion of hMSC to substratum. hMSC clusters can be differentiated homogenously into hyaline chondrocytes and express specific gene like Col II, aggrecan (ACAN), and cartilage oligomeric protein (COMP). By the expression of non-hyaline chondrocyte genes CD105, Col X, and Col I, the pellets became heterogeneous in distribution. This novel microwell substratum technology directed the differentiation of hyaline chondrocyte cells from hMSCs and this provides a valuable system for experimental and potential clinical studies (Kitagawa et al. 2013). Porous materials can also be used besides synthetic and natural polymers. Some of the inorganic porous 3D structures used for cell culture scaffolds include graphene foam, multiwall carbon nanotubes, glass microtubes, and graphene oxides (Table 8.1). They possess high mechanical stability and high porosity with tight interconnectivity, which make them ideal for highly interactive 3D culture. The high porosity allows deeper and more uniform nutrient transport and the cells can freely migrate along the structure without any significant resistance. Spencer et al.

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Table 8.1 Scaffolds for 3D stem cell culture Scaffold material Type 1 collagen hydrogel

Fibrin hydrogel

Gelatin hydrogel

PLA 3D printed scaffold

Stem cells Human neural progenitor derived astrocytes Human endometrial stem cells (hEnSCs) Mouse ESCs

Human adiposederived stem cells (hASCs) hMSCs

3D aligned poly(L-lactic acid) (PLLA) and polyacrylonitrile (PAN) nanofibers by electrospinning Methacrylate-modified hyaluronic acid hydrogel

Human iPSCderived NPCs

Polyurethane hydrogel

Adult mouse NSCs

Application Promotes axon growth on dorsal root ganglion (Führmann et al. 2010) hEnSCs are more efficiently differentiated into Schwann cells (Bayat et al. 2016) 3D culture enhances the differentiation of ESCs into functional thyroid tissue (Antonica et al. 2017) Allows higher levels of cell proliferation within bioprinted strands (Narayanan et al. 2016) Support neuronal activity and induce cell growth along the lengths of the nanofibers (Wu et al. 2018) Layered hydrogel influences migration and differentiation (Zhang et al. 2016) Thermo-responsive biodegradable polyurethane bioink for 3D printing (Hsieh et al. 2015)

developed three-dimensional graphene foams to promote osteogenic differentiation of human mesenchymal stem cells. Multilayered graphene foams were fabricated by growing graphene in 3D Ni scaffolds. Nickel was then subsequently removed by FeCl3 etching and then investigated for promoting hMSC attachment, maintenance of cell viability, and spontaneous morphological changes. The results showed that 3D graphene foams induced the spontaneous osteogenic differentiation of hMSCs without the need for extrinsic biochemical manipulation. The fate of hMSC was influenced by cytoskeletal tension and cell shape. Rho/ROCK cascade was observed to control osteo-/ adipogenic lineage commitments in which extreme intracellular tension was shown to bias the cell towards osteogenesis. Such novel graphene-based strategies can be used for the development of osteoconductive tissue engineered constructs. The fabrication of 3D GFs is very cost effective and highly scalable approaches for tissue engineering construct development (Crowder et al. 2013). The major disadvantages of such inorganic porous scaffolds are their degradation profile. They are much harder to degrade than bulk materials. This limits their clinical applications. Another disadvantage is the opaque nature which limits the light transmission, making in situ imaging very difficult.

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Scaffolds for Biomolecule Delivery

Scaffolds are three-dimensional biomaterials used as implants or injectables to aid in tissue regeneration and may be loaded with cells or therapeutic molecule and are used to support damaged tissues or organs. Scaffolds are successfully utilized in various fields of tissue engineering such as bone formation, periodontal regeneration, cartilage repair and development, tendon repair, valve replacement, and ligament replacement. Scaffolds for delivering therapeutic molecules to the targeted site so as to promote cell/tissue growth are being explored widely as they can be also used as controlled delivery vehicle over a long period of time.

8.8.1

Properties

Tissue engineering scaffolds help in the cell colonization and transmission of physical and chemical cues for tissue growth. The synthetic scaffolds are meant for local delivery of proteins and growth factors for tissue repair and regeneration. The scaffold for drug delivery application should possess the following characteristics: • There should be homogenous distribution of drug throughout the scaffold. • Able to release the drug at a predetermined rate. • At physiological temperature the drug binding affinity should be low to allow stable drug release. The physical dimension, chemical structure, and biological activity of the scaffold should be stable over a period of time.

8.9

Biomolecule Delivery Systems

There are numerous biomolecule delivery strategies employed in tissue engineering (Fig. 8.1), which can be used depending on the tissue of interest. The scaffolds serve as synthetic extracellular matrix for cellular organization in three-dimensional architecture. Depending on the site of application and tissue of interest, the required form and the properties of the scaffolds may vary (Drury and Mooney 2003). Biomolecule loaded delivery systems can be either incorporated within the scaffolds or can be delivered independently via injections.

8.9.1

Hydrogel-Based Systems

Hydrogels possess various biological applications due to its high-water content and biocompatibility. Biomolecules are released from these highly hydrophilic scaffolds through mechanical stimulation, hydrolytic degradation, or upon swelling by environmental stimulation. The release behavior can be regulated from few days to

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Fig. 8.1 Different systems for biomolecule delivery (a) microspheres, (b) nanoparticles, (c) emulsions, (d) membranes, (e) liposomes, (f) microchips, (g) hydrogels, (h) dendrimers, (i) injectables, (j) elastomers, (k) micelles

several months by controlling physical and chemical properties. Hydrogels can be administered in minimally invasive manner, so used as potential carriers of cells and proteins. Hydrogels are capable of protecting drugs and proteins from the harsh environment at the site of release. They have potential role in wound healing applications. The incorporation of viable cells into the scaffolds is potentially a challenge. Since the hydrogels are made up of highly hydrated polymers, they are widely employed for such application. Hydrogels have different functions in the field of tissue engineering. They can be applied as space filling agents, as biomolecule delivery vehicle, and as threedimensional cellular organization and signal presentation for tissue formation. Hydrogels can maintain a desired volume and structural integrity for the required time. They have found great utility in preventing post-operative adhesions and in plastic and reconstructive surgery. A dextranomer/HA copolymer was used by Seibold et al., as a bulking agent for vesico-ureteral reflux (Seibold et al. 2011). Hydrogels can be used as barrier to avoid restenosis or thrombosis. Hydrogel based on poly(ethylene glycol-co-lactic acid) diacrylate developed by bulk polymerization was able to prevent fibrin deposition and fibroblast attachment at the tissue surface (Hill-West et al. 1995). Hydrogels composed of chitosan and chitin derivatives are also proposed to use as biological adhesives (Zhao et al. 2001).

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Nanoparticle Based Systems

Nanotechnology offers site specific and target oriented delivery of medicines for treating chronic human diseases. Their distinctive site dependent properties like nano scale dimension enable them to make beneficial for wide range of applications. They can be used for controlled drug delivery, imaging of specific sites, probing DNA structure, biomolecule sensing, gene delivery, photo thermal ablation and more over many therapies utilize nanoparticles (Kingsley et al. 2013). The surface conjugation of gold nanoparticles, antimicrobial effect of silver metallic properties of metal oxides, fluorescence properties of quantum dots, and electro mechanical properties of carbon nanotubes are applied in gene delivery, cell mechano transduction, construction of 3D tissue complex structure, and controlling cell patterning. The nano size and large surface to volume ratio are widely explored in tissue engineering. Peptides and small proteins can easily diffuse across the membrane and facilitate uptake by the cells. Nanoparticles can be customized into sizes and surface characteristics in order to suit for any purpose. They also mimic the natural nanometer size scale of extra cellular matrix components of tissues. Gold nanoparticles (GNPs) and titanium oxide (TiO2) nanoparticles are used to enhance cell proliferation for bone and cardiac tissue regeneration. GNPs promote osteogenic induction in bone tissue engineering (Giljohann et al. 2010). TiO2 embedded nanocomposite polymers exhibited superior mechanical properties and showed higher tensile strength in reinforcing the scar after myocardial infarction (Kumar 2018). As drug release systems, nanostructures stay in blood circulatory system for a prolonged period and enable the release of drugs as per the specified dose. They cause very little plasma fluctuations and reduce adverse effects. Being nanosized, they can penetrate tissue structures and facilitate the easy uptake of the drugs, thus ensure action at the targeted site. Nanostructures can deliver drugs in ways such as passive or active methods (Fig. 8.2). The drugs can be incorporated into the inner cavity via hydrophobic effect for passive release. The prepared drug complex circulates in the bloodstream and driven to the target site by affinity or binding influenced by pH, temperature, molecular site or shape (Hasan et al. 2018). Drugs intended for release can be directly conjugated to the nanostructure for easy delivery. Here timing of release is crucial, as the drug dissociates from the carrier very quickly, there is chance of reduction of its biological activity.

8.9.3

Liposomes

Liposomes are self-assembled vesicles that are able to encapsulate aqueous solutions and hydrophobic compounds. Liposomes are widely investigated as drug carrier for various types of drugs. Liposomes have been used in different areas such as vaccines, imaging, therapeutics, and tissue engineering applications (Monteiro et al. 2014). The successful treatment of liposomes also depends on the route of administration. Stimuli responsive liposomes are used to release the therapeutics at the site of

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Fig. 8.2 Different release mechanisms for biomolecule delivery

action. The advantage of liposome loaded scaffolds is that the drug delivery from liposomes can be prolonged unlike that of parenteral delivery systems.

8.9.4

Micelles

Polymeric micelles are self-assembled core–shells nanostructures formed in aqueous solution of amphiphilic block copolymers. When concentration of block copolymers increases over a certain concentration, micelle formation occurs, known as the critical micelle concentration. Micelles can overcome the limitations of oral drug delivery by acting as a carrier and able to enhance drug absorption. They provide protection of the drugs from extreme harsh environment, help in the release of drugs in controlled rate, prolonged residence time by mucoadhesion, inhibition of efflux pump to improve drug accumulation. Micelles can deliver poorly water soluble drugs with well-maintained bioavailability (Xu et al. 2013). Zhang et al demonstrated ultra-long block copolymer fibrous micelles which can modulate cell orientation on surface. The degree of cell alignment increases with density of micelles. For high density micelles, nuclear alignment also observed.

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Cells showed competitive response to the micelle network with multi directionality alignment. This can be used to mimic the native fibrous networks surrounding cells (Zhang et al. 2018). Santo et al developed dextran-micelles internalized by rat bone marrow mesenchymal stem cells which demonstrated a pH responsive release profile and enhancement of 2D and 3D in vitro osteogenic differentiation. This also promoted a significant enhancement of bone formation in rat ulna defect in a dose dependent manner (Santo et al. 2015).

8.9.5

Microparticles

Micro encapsulation is one of the intelligent approaches with strong therapeutic impact due to its specific therapeutic properties and target specificities. They can encapsulate both water soluble and sparingly soluble agents to elicit efficacy with great potential. Microparticles provide initial architecture as well as micro environment required for supporting cell growth (Chau et al. 2008). Hydroxyapatite/collagen/phosphatidylserine scaffolds embedded with steroidal saponin loaded collagen microparticles were prepared by Yang et al. using porogen leaching protocol. The scaffold consisted of dense and loose layers with inter connected pores. Microparticles entrapped within the scaffolds by gradient distribution. Loose layer showed greater drug release compared to dense layer. Cell proliferation was also more in loose layer than dense layer. Such spatial and temporal control over drug release provides opportunities for tissue regeneration with optimum dose at the site and reduces undesirable drug release (Yang and Fang 2015).

8.9.6

Dendrimers and Elastomers

Dendrimers are unimolecular architects that can incorporate a variety of biological or chemical substances in a 3D architecture to actively support the scaffold microenvironment during cell growth. For drug delivery applications, they are employed in two ways such as formulation and nanoconstructs. In formulation approach, drugs are physically entrapped by non-covalent interactions, whereas in nanoconstructs, drugs covalently coupled. Polyamidoamine (PAMAM) dendrimers have shown to increase transdermal permeation and specific drug targeting. Dendrimers are highly branched, monodisperse and radially symmetric macromolecules of nanosize with a 3D structure. Due to these unique structure dendrimers are widely used in drug delivery applications (Chauhan 2018). Elastomers are polymers with viscoelasticity and have very weak intramolecular forces, with low Young’s modulus and high failure strain. Biodegradable elastomers have potential applications in soft tissue engineering, where the mechanical properties of the polymer scaffold match with the tissue to be grown. This polymer scaffolds can withstand repeated dynamic load, provide suitable surface for cell attachment and growth. They can ultimately degrade at a rate that allows the load to be transferred to the new tissue (Ye et al. 2018). Elastomers are popularly used in

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vascular tissue engineering applications, as they offer the ability to design implants that match the compliance of native tissue. By mimicking natural tissue environment, elastic materials able to integrate within the body to promote repair and avoid adverse physiological response. Elastomers can also be used in osmotic pressure drive drug delivery, which is effective in providing a constant release of drugs for prolonged time. More elastic materials such as poly (glycerol sebacate), collagen, and elastin closely match with the arterial compliance. Less elastic materials such as polyurethane, PCL, and silk also have utility when used together with more elastic materials (Hiob et al. 2016).

8.9.7

Microchips

Microchip offers both rate and time release of molecules. The device consists of substrate incorporated multiple reservoirs, capped with conductive membrane and wired with integrated circuit controlled by microprocessor. Reservoirs are fixed into substrate by chemical etching or ion beam etching techniques. Hundreds of reservoirs can be fabricated into a single microchip using microfabrication. The biomolecule to be released is injected into the reservoir. Reservoir can enclose multiple drugs in different doses (Eltorai et al. 2016). Farra et al. studied the in vivo pharmacokinetics of human parathyroid hormone released from microchip devices in human patients with osteoporosis. The release from the device was activated 8 weeks after implantation to allow formation of tissue capsule. It has been given for 4 months and wirelessly programmed to release doses from the device once daily for up to 20 days (Farra et al. 2012).

8.10

Scaffold Based Biomolecule Delivery

8.10.1 Delivery of Therapeutic Drugs The aim of drug delivery is for administering a pharmaceutical compound to achieve a therapeutic effect in human or animals. For this purpose, several drug delivery systems have been formulated and investigated. Delivering drug at a controlled rate, slow delivery, or targeted delivery is being investigated. Hydroxyapatite (HAp) scaffolds with high porosity, controlled pore size, and adequate hardness help in slow, controlled, and sustained release of drugs at the affected site. The major limitation of HAp in the fabrication of scaffolds for biomedical application is the limited control over pore size, shape, and distribution. Incorporation of selective biomolecules such as chitosan with HAp can induce bactericidal activity. Adding such new functionality is very useful for the protection of regenerated new tissues from infection. Zhang et al studied the drug loading and release profile of gentamicin sulfate in chitosan-Hap microtubes. The material showed high loading capacity of 976.6 mg/g with sustained release profile and

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improvement in mechanical properties (Zhang et al. 2017). In another study by Kim et al, coated HAp with Poly (ε-caprolactone) for delivery of the antibiotic vancomycin. The bare HAp particles showed initial burst release of 70–80%, whereas coated particles with only 44% initial release of vancomycin. The release rate was sustained over a prolonged period, which was depend on the degree of coating dissolution (Kim et al. 2004). To address the bone resconstruction Chen et al. (2017) developed a new drug delivery system in which the hydrophillic drug of interest, desferrioxamine, was loaded in liposomes which was then embedded in gelatin hydrogel (Gelma). The idea was to provide a prolonged sustained release of drug over a period of time thereby ensure the therapeutic effect which can lead to bone reconstruction. Additive manufacturing is another promising technology in regenerative medicine. Here scaffolds are made in a layer-by-layer manner, which enables the direct construction of complex structures with very high precision (Hammond 2012). The antibiotic gentamicin was layered with polyacrylic acid and a degradable component poly (β-amino ester) in the form of a thin film. The film showed a rapid release of gentamicin from the surface followed by sustained release over multiple weeks. The burst release behavior was due to diffusion and slow release by hydrolytic degradation (Moskowitz et al. 2010). Combining multiple drugs by co-encapsulating into a single carrier is a promising strategy in tissue engineering. Drug can be physically loaded into the core of the material or it can be covalently linked to the polymer backbone. By coating drugs by different processing materials, drugs can be quickly separated and released independently. Kang et al. developed thermo-responsive nanospheres for dual delivery of drugs. The nanospheres were prepared by conjugating chitosan oligosaccharide with pluronic F127 for simultaneous but independent delivery of kartogenin (KGN) and diclofenac (DCF). One of the drugs (KGN) was conjugated on to the outer layer of nanopshere and DCF was loaded in the inner core. On exposure to cold temperature the nanospheres deliver DCF immediately and KGN in a sustained manner which minimised LPS induced inflammation in chondrocytes and induced differentiation of mesenchymal stem cells into chondrocytes. The nanospheres were capable of suppressing the progression of osteoarthritis in rat model (Kang et al. 2016). In another study by Wang et al., described the development of novel drug release controlling system for multiple drug delivery. They have developed chitosan nanoparticles/PCL composite electrospun nanofibers with core–sheath structures, in which two model agents’ rhodamine B and naproxen were successfully loaded in the core and sheath regions, respectively. They demonstrated good controlled release and temporality, which provides a new way to release multiple drugs, especially in suturing and wound dressings (Wang et al. 2010).

8.10.2 Delivery of Therapeutic Cells Cell encapsulation techniques are used for the purpose of utilizing the cells to secrete the molecule of interest for sustained release. By this technique, a substance can be

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released over a long period of time in a manner responsive to the need of the body. Hydrogel scaffolds are often utilized to stabilize and deliver bioactive molecules and encapsulate secretory cells. Highly hydrated three-dimensional networks of polymers provide a place for cells to adhere, proliferate, and differentiate. They provide chemical cues for the cells through growth factors or by mechanical signals. Currently, hydrogel scaffolds are widely employed in various tissues such as cartilage, bones, muscle, fat, liver, and neurons. A novel stem cell bandage for cellular delivery was developed by Asawa et al. They cultured mesenchymal stromal cells on the surface of PEG-DMA hydrogels with RGD adhesive peptides and applied to the wound surface. The hydrogel allowed an initial cellular adhesion for multiple days with a decrease by day 15. This bandage type approach has advantage over direct stem cell injection or encapsulation, because it prevents diffusion of the cells away from the area of interest and allows access to the site (Asawa et al. 2018).

8.10.3 Scaffold Based Peptide Delivery The tissue engineering process involves complex cascades of peptides such as growth factors, cytokines, and other molecules. Growth factors are endogenous polypeptides, which act through surface receptors to regulate cellular activities. The outcome of growth factor therapeutics depends on the delivery mode due to their rapid clearance in vivo. Sophisticated material systems that regulate the release of growth factors represent new therapeutic modalities for treating wide variety of diseases. Growth factors have distinct therapeutic applications such as bone regeneration, neovascularization, cell proliferation and differentiation, etc., growth factors can be chemically immobilized or physically encapsulated into polymer matrices (Lee et al. 2011). The chemical modification of polymers and physical encapsulation of growth factors are critical to increase therapeutic efficacy. Poly PLA nanofibers were coated with hydroxyapatite nanoparticles to stimulate bone mineralization. Bone morphogenetic protein-2 peptide loaded liposomes were grafted into the scaffolds by covalent bond to regulate release of peptide. This scaffold was observed with favorable biocompatibility and satisfactory ability for promoting osteogenic differentiation of MSC (Mohammadi et al. 2018). Simultaneous or sequential delivery of multiple growth factors has been also exploited to enhance therapeutic efficiency. Composite polymers are used to design spatiotemporal delivery of multiple growth factors. Degradable alginate hydrogelbased delivery systems provided simultaneous delivery of osteogenic growth factor and other morphogens. Alginate was gamma irradiated to vary degradation rate, then covalently modified with RGD peptides to control cell behavior. Bone morphogenetic protein-2 (BMP2) and transforming growth factor-beta3 (TGF-beta3) were incorporated into the hydrogels and showed significant bone formation as early as 6 weeks after implantation. The study shows that appropriate combination regulatory signals, both soluble and biomaterial mediated, in cell-based tissue engineering

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approaches can be more efficent and effective for regeneration (Simmons et al. 2004).

8.10.4 Scaffolds for Gene Delivery Gene delivery has been used in regenerative medicine to create or restore normal function at the cell and tissue level. Cells with excellent proliferation capabilities and differentiation potential have been taken as the candidate for gene delivery. Multipotent and pluripotent stem cells are widely studied in this aspect (Fang et al. 2015). Stem cell sources such as umbilical cord, amniotic fluid, and fetal tissues can be used for gene delivery, as they are capable of differentiating into three germ lineages. Gene delivery can be performed either in ex vivo or in vivo conditions. During ex vivo method, outside the body cells exposed to delivery agent. While in the case of in vivo technique, the implanted construct matures within the body, so that there are no distinguishable differences between the area of implant and the surrounding area. Directing extracellular matrix in appropriate way is the key aim of regenerative medicine. Monaghan et al proposed antifibrotic interfering RNA therapy for remodeling extracellular matrix after cutaneous injury. Exogenous micro RNA (miR 29B) from collagen scaffold efficiently modulates remodeling response and reduces aggressive deposition of collagen type I after injury. Primary fibroblast cultured scaffold doped with miR-29B showed reduced level of collagen type I and expression of collagen type III mRNA observed up to 2 weeks of culture. In vivo application of this scaffold functionalized with miR-29B showed wound healing within 2 weeks, with improved collagen type III/I ratio and significantly higher matrix metalloproteinase (Monaghan et al. 2014). Biomaterial scaffold for bone tissue engineering can regulate cell behavior and induce bone growth. The commonly used bone repair scaffolds include hydroxyapatite, electrospun 3D scaffolds, and hydrogels. Li and coworkers developed an injectable chitosan based thermo sensitive hydrogel scaffold incorporated with BMP-2 plasmid DNA. They found that these scaffolds could able to enhance new bone formation in calvarial defects of rats (Li et al. 2017).

8.11

Biomolecule Loaded Scaffolds in Tissue Engineering: Applications

8.11.1 Bone Tissue Engineering Bone regeneration is possible by the stimulation of various kinds of cytokines and osteogenic precursor cells such as osteoblasts and osteoclasts. They induce bone regeneration and matrix mineralization. But in the case of any bone defect, there will be lack of mechanical support and 3D environment for the attachment and differentiation of cells. In such conditions, composite scaffolds are needed which can release

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drug in situ for long term and fill the defective area to provide mechanical support to the entire system (Amini et al. 2012). Bioceramics are one of the most commonly used materials for skeletal system regeneration. They are commonly used in tissue engineering due to their similarities with the mineral components of hard tissue as bone. Bioceramics are biocompatible and can be bioactive or bioresorbable. They are presented with common characteristics such as hard refraction, polycrystallinity, high melting temperature, low electric conductivity, and corrosion resistance. The major applications include hard tissue replacement, periodontal, cranial, maxillofacial, dental, spinal, and otolaryngology surgery. Bone is the second most transplanted tissue and it is considered as composite of inorganic and organic phase comprising of apatite and collagen/glycoproteins, respectively. It is a highly organized anisotropic structure comprising of nano to the macro scale which bestows this tissue with its unique strength, load bearing capacities, and mechanical behavior (Diaz-Rodriguez et al. 2018). Owing to the in-depth knowledge on bone tissue, biomimetic approaches are gaining wide used in developing scaffolds for bone tissue engineering applications to improve its mechanical behavior and performance. The biomimetic design of these scaffolds is thus now based on the composition of native tissue structure and/or composition of the bone tissue to be replaced. Calcium phosphates are commonly used for bone regeneration due to their osteoinduction and mimesis of tissue composition. Using 3D printing polymer–ceramic composites are developed for osteochondral regeneration. Such composites have improved mechanical characteristics and interfacial integration. Bioactive glasses are another family of ceramics used for bone regeneration, due to their ability to promote hydroxyapatite formation and osteoconductive character. To improve their performance, therapeutic molecules are incorporated, which aims to promote biologic repair and provide support for treatment.

8.11.2 Skin Tissue Engineering Skin tissue engineering uses a combination of cells, engineering principles and materials. The aim of skin tissue engineering is to develop 3D scaffolds with appropriate biomaterial which can functiona as extracellularmatrix and promote cell adhesion, growth as well as differentiation resulting in a functional construct. In this aspect, biomaterial and cell selection are both equally important. The scaffolds can be degradable or non-degradable and capable of delivering the loaded therapeutic molecules as well. Nanofibrous scaffolds provide large area to volume ratio thus more surface area for cell attachment and proliferation. Ghaee et al. synthesized a biomimetic scaffold using PCL nanofibers embedded in a hydrogel matrix of PEGMA, which resembles the ECM structure. This nanocomposite showed suitable porosity and mechanical properties for skin scaffold. This could able to deliver about 80% loaded curcumin in a controlled manner. This scaffold also

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offered antioxidant properties and showed in vitro biological performance for skin regeneration (Ghaee et al. 2019). ECM can be synthesized by decellularization process, which removes cells from ECM thus make it free of potential antigens causing inflammatory response or immune mediated implant rejection. The optimal method for decellularization is highly dependent on type of tissue. After the removal of cellular components, a 3D fibrous and porous scaffold is maintained which is composed of collagen fibers. The main benefit of the scaffold is its porous structure and macrostructure like the vasculature. ECM has great potential as cell and drug carriers. ECM based drug delivery system has been well established in skin wound healing. Stem cell therapy is another novel technology for regenerating damaged tissues. There are numerous sorts of stem cells including epidermal stem cells, melanocyte stem cells, mesenchymal stem cells, and human newborn foreskin stem cells. Engineered skin alternatives deliver a conceivable resolution to the problems of donor implant scarcity prevent from liquid loss and infection. About 2–5 cm2 skin biopsies picked from one individual and expanded in vitro for developing epidermal skin grafts. Then epidermis detached from the dermis and keratinocytes are chemically discharged and cultivated on mitotically incapacitated mouse fibroblasts (Boyce and Lalley 2018). Some of the skin substitutes used for skin regeneration are included here (Table 8.2).

Table 8.2 Tissue engineered products developed as skin substitutes Product AlloDerm

Composition Decellularized human dermis

MatriDerm

Bovine collagen and elastin

Hyalomatrix

Derivatized hyaluronic acid

Apligraf®

AllohF in collagen gel plus stratified allohK Cultured auto hK multilayer sheet

EpiCel® StrataGraft® ReCell® Reconstructed skin

AllohF in collagen gel plus stratified allohK Uncultured suspension of auto hK, delivered as a spray Auto hF on acellular scaffold of dermal extracellular matrix, plus stratified auto hK full-thickness burns

Intended use Repair or replacement of damaged or inadequate integumental tissue (Jansen et al. 2013) Burns, reconstructive surgery (Halim et al. 2010) Partial- and full-thickness wounds (Gravante et al. 2007) Diabetic foot ulcers (Carlson et al. 2011) Full-thickness burns (Sood et al. 2010) Partial-thickness burns (Schurr et al. 2012) Partial-thickness burns (Wood et al. 2012) Full-thickness burns, venous and mixed ulcers (Boyce and Lalley 2018)

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8.11.3 Cartilage Tissue Engineering Biomaterial scaffolds play important role in cartilage tissue engineering. The geometric design of the scaffolds has significant role in cell migration. MSC homing can be enhanced by the use of radially oriented scaffold with ordered and aligned channels than non-oriented scaffolds. Geometric properties such as elasticity and hydrophilicity can also affect cell attachment and proliferation within the scaffolds. Elasticity and stiffness of scaffolds are particularly very much important for cartilage tissue engineering, because it is subjected to cyclic mechanical forces. The degradation rate is also very important, faster rate may cause scaffold to collapse before new tissue formation. Due to proper tensile strength, high elasticity and good biocompatibility polyurethane scaffold materials are used for cartilage tissue engineering. The biodegradation properties and mechanical strength can be adjusted by its soft segment composition. MSC seeded 3D printed PU scaffolds were found to aggregate inside the scaffolds before differentiation into chondrocytes (Wen et al. 2019b). PU could be also made into microspheres to carry drugs for controlled drug release. Wen et al had prepared SDF-1 loaded microspheres to promote MSC migration. The effective concentration of 100 ng/ml released after 24 hours could induce the migration of hMSCs. The scaffold showed a significant GAG (glycosaminoglycan) deposition within 7 days. They have implanted the scaffold in rabbit articular cartilage defect and promoted cartilage regeneration.

8.12

Future Perspectives

The purpose of tissue engineering approach was initially to address the critical gap between growing numbers of patients listed for organ transplantation. It also focuses on prevalent conditions in which the restoration of functional tissue would answer currently unmet medical needs. Recent progress in this field suggests that such engineered tissues expanded the clinical applicability and represents a viable therapeutic option for life extending benefits of tissue replacement or repair. But to date, only handful complex products like cell seeded scaffolds have gained regulatory approval and also, they have obtained limited marketed penetration. So, a combination of advances in both clinical development and commercialization is needed in coming future. Bioactive materials are designed to have the features of extracellular matrix which plays a key role in cellular adhesion, migration, and new tissue formation. Unlike tissue engineering constructs loaded with cells, the mode of tissue regeneration by bioactive scaffolds is by providing a near native environment that can promote cell adhesion and tissue growth. Added advantage of current strategy is to deliver appropriate biomolecules through the scaffolds in controlled manner to stimulate the tissue regeneration. If the bioactive scaffold loaded with appropriate biological cues can elicit normal tissue growth and regeneration, the clinical translation will be achievable at a faster pace and the logistic issues associated with cell loaded constructs can be minimized. The bioactive biomaterials utilized for

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regenerative purpose should have required function for specific time period and able to function in the physiological environment without eliciting adverse reactions. However, developing bioactive materials with such accuracy and efficacy is challenging. Thus, it is necessary to have deep understanding regarding the mechanism of regeneration for exploring new approaches for designing multifunctional biomaterials that can result in clinical use.

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Trends in Stimuli Responsive Biomaterials in Tissue Engineering Rajiv Borah, Jnanendra Upadhyay, and Birru Bhaskar

Abstract

Native tissues and organs coordinate and execute their activities via dynamic, interlinked clusters of biochemical and biophysical attributes, which are differed throughout biological processes spatiotemporally. Passive biomaterials, developed with tunable structural, mechanical and biochemical properties, cannot mimic the dynamic features of the cellular environment and therefore, often lack of efficiency in tissue regeneration to restore full functionality. With the perspective to address this notion, stimuli responsive biomaterials have evolved as effective tool that replicate essential static and dynamic features of native tissues due to their capacity to alter physicochemical characteristics in response to physical/chemical/biological stimuli compatible to tissues and organs, facilitating on demand cell microenvironmental manipulation. The current chapter focuses on trends of stimuli responsive biomaterials explored for tissue engineering (TE) applications. Special emphasize has been devoted to those stimuli responsive biomaterials (e.g. electroactive biomaterials), which are sensitive to the stimuli that match with the native biophysical cues of tissues and can regulate those biophysical cues to modulate the regeneration associated cellular processes for faster and efficient tissue regeneration. Each category of stimuli responsive biomaterials has been discussed with a brief introduction and the mechanism of

R. Borah (*) Life Sciences Division, Institute of Advanced Study in Science & Technology, Guwahati, Assam, India e-mail: [email protected] J. Upadhyay Department of Physics, Dakshin Kamrup College, Mirza, Kamrup, Assam, India B. Bhaskar Brien Holden Eye Research Centre, LV Prasad Eye Institute (LVPEI), Hyderabad, India # The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 B. Bhaskar et al. (eds.), Biomaterials in Tissue Engineering and Regenerative Medicine, https://doi.org/10.1007/978-981-16-0002-9_9

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functionality followed by its applications in various TE applications. Biomaterials that respond to chemical and biological stimuli, have also been briefly addressed in the light of TE potential. The chapter also highlights the advantages-limitations and future directions of stimuli responsive biomaterials at the end. Keywords

Tissue engineering · Stimuli responsive biomaterials · Electroactive biomaterials · Tissue regeneration

Abbreviations 0D 1D 2D 3D 5-FU ADSCs Alg BaTiO3 BT C Ch CNFs CNTs Co Col CPs CS DDF DLC DMAEMA DNA ECM ES Fe G Gel GelMA GO HA HEMA LCEs

Zero dimensional One dimensional Two dimensional Three dimensional 5-fluorouracil Adipose derived stem cells Alginate Barium Titanate Barium titanate Cellulose Chitosan Carbon nanofibers Carbon nanotubes Cobalt Collagen Conducting polymers Chondroitin sulfate Dermal fibroblasts Diamond-like carbon Dimethylaminoethyl methacrylate Deoxyribonucleic acid Extracellular matrix Electrical stimulation Iron Graphene Gelatin Gelatin methacryloyl Graphene oxide Hydroxyapatite 2-hydroxyethyl methacrylate Liquid crystalline elastomers

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LCST MAP MEH-PPV MS MWCNT NB NCD Ni P3HT PAN PAni PAs PCBM PCL PCLF PDMS PEDOT PEG PEGS PEO–PPO–PEO PGA PGS PHBV PHEMA PLA PLGA PLGA/HA PLLA PLLA-PEG-PLLA PNiPAAm PNVC POxs PP PPy PSS PT PTCDI-C8 PTFE PVA PVDF PZT RGCs rGO SF

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Lower critical solution temperature Microporous annealed particle Poly(2-methoxy-5-(2-ethylhexyloxy)-1,4-phenylenevinylene) Magnetic stimulation Multiwall carbon nanotube Nitrobenzene Nanocrystalline diamond Nickel Poly-3-hexyl-thiophene Polyacrylonitrile Polyaniline Peptide amphiphiles Phenyl-C61-butyric acid methyl ester Polycaprolactone Polycaprolactone fumarate Polydimethylsiloxane Poly(3,4-ethylenedioxythiophene) Poly(ethylene glycol) Poly(ethylene glycol)-co-poly(glycerol-sebacate) Poly(ethylene oxide)-poly(propylene oxide)-poly (ethylene oxide) Polyglycolide Poly(glycerol-sebacate) Poly(3-hydroxybutyric acid-co-3-hydroxy valericacid) Poly(2-hydroxyethyl methacrylate) Polylactide Poly(lactic-co-glycolic) Poly(lactic-co-glycolic acid)/hyaluronic acid Poly-L-lactic acid Poly (L-lactic acid)-poly(ethylene glycol)-poly(L-lactic acid) Poly(N-isopropylacrylamide) Poly(N-vinylcaprolactam) Poly(2-oxazoline)s Polypropylene Polypyrrole Poly(4-styrene sulfonate) Polythiophene N,N0 -dioctyl-3,4,9,10-perylenedicarboximide Polytetrafluoroethylene Poly(vinyl alcohol) Polyvinylidine fluoride Lead Zirconate titanate Retinal ganglion cells Reduced graphene oxide Silk fibroin

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Si SMPs TCP TE TrFE UCST UV

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Silicon Shape-memory polymers β-tricalcium phosphate Tissue engineering Trifluoro ethylene Upper critical solution temperature Ultraviolet

Introduction

Tissue engineering (TE) has evolved as a realistic alternative to donor-dependent organ transplantation or autografting and allografting to repair a damaged organ or tissue. It is meant to develop living, functional tissues that can be employed to substitute, or repair tissues impaired because of disease, aging, congenital defects or physical damage by integrating biomaterial scaffold, cells and bioactive compounds (Vacanti and Langer 1999). Therefore, the choice and design of biomaterial is essential for the regeneration of new cells in vitro and in vivo, while ensuring its biocompatibility, bioactivity, durability, degradability, porosity, and flexibility at the same time. In TE, the biomaterial scaffold should act as the artificial extracellular matrix (ECM) capable to mimic the native cellular microenvironment of the particular cell type, which is needed to be regenerated. Hence, the spatiotemporal modulation of the physical and chemical properties of biomaterial scaffold is necessary to support favorable tissue regeneration. The native ECM interacts with cells dynamically through close co-ordination with the biophysical and/or biochemical cues for normal tissue function including regeneration. A biomaterial scaffold is also required to function in a dynamic fashion for effective and efficient tissue regeneration, which paved the way for “smart” or “stimuli responsive” functional materials in TE applications. Throughout the designing and creation of new materials that are able to respond to particular stimuli, nature provides countless touchstones that are configurable, reliable, and replicable. In reality, many living system substances vary spontaneously as per the environmental circumstances and their processes and actions to maintain and regulate normal functions. It involves alteration in form, dimensions, appearance or rigidity and depends on complicated models for feedback. Over the last decade academic and industrial research has thus been inspired to create new functional materials that imitate the sensitivity of natural living systems. Subsequently, the understanding of endogenous physiological behavior of cells and tissues along with the existence of several important physical and chemical cues, inspired researchers to develop a new generation of biomaterials, termed as “Smart” or “Stimuli responsive” biomaterials. Prior to this new generation of biomaterials, most of the biomaterials were used in a passive way, just as support for the cells and tissues through their bioactivity and suitable physiochemical properties such as

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biodegradability, mechanical stability, and porosity. Therefore, there is a growing interest in stimuli responsive materials for TE and regenerative medicine with the capacity to communicate and interact with cells. Stimuli responsive materials, also termed as “smart” or “intelligent” materials, are those, which can sense and respond to external stimuli or any alternation in the external environment (Cardoso et al. 2017). In rebuttal to single or multiple external stimuli, this exceptional category of materials exhibits variations in one or more of their physicochemical properties, i.e., size, shape, solubility, permeability, hydrophilicity, surface charge, electrical, magnetic, mechanical, and optical, etc. These external stimuli can be classified as physical (temperature, electrical, magnetic, mechanical stress, light, ultrasound, etc.), chemical (pH, ionic strength, electrochemical, etc.) and biological (enzymes, glucose, antigen, growth factors, receptors etc.) stimuli (Fig. 9.1). Physical stimuli can induce modifications in the energy dynamics of the materials, whereas the chemical stimuli modulate molecular interaction within the material or between the material and the surrounding environment. Biological stimuli associate with the specific biological functions such as enzymatic reactions, receptor recognition, activating regeneration associated processes, etc. Additionally, there are dual and multi-stimulus-responsive materials that respond to more than one stimulus concurrently. In regard to TE applications, stimuli responsive materials hold potential to elicit beneficial effect at cellular level through changes in their physiochemical properties upon any change in external stimuli, which can activate regeneration associated processes by modulating various important biochemical or biophysical events at cellular and molecular level. Therefore, it is important that the stimulus dependent behavior of a potential stimuli responsive biomaterial should be able to induce the beneficial effect during in vitro cell culture or in vivo to enhance the tissue regeneration and function. Although, there are a range of stimuli responsive biomaterials with respect to their sensitivity towards specific stimulus type, the real time cellular response is significantly observable and therefore, well explored with the stimuli responsive biomaterials, which can respond to electrical stimulation (ES) and magnetic stimulation (MS). The concept of these biomaterials is based on the intrinsic biophysical cues already present in the tissue. In fact, there are two approaches of using stimuli responsive biomaterials for tissue repair purposes. In the first case, the stimulus is used during fabrication of the biomaterials and there is hardly any or rare evidence of utilizing that particular stimulus in real time during in vitro cell culture or in vivo. In the latter’s scenario, the stimulus, which matches with the intrinsic biophysical/biochemical cues of tissues, is utilized to mimic the dynamic cell microenvironment. The present chapter mainly focuses on the stimuli responsive biomaterials of the second category with their underlying mechanisms of stimuli dependent actions in the light of cellular processes, and hence, a detailed discussion on electroactive and magnetoresponsive biomaterials, followed by thermoresponsive and photoresponsive biomaterials, has been presented. The chapter also summarizes the TE applications of chemical and biological stimuli responsive biomaterials along with dual and multi-stimuli responsive biomaterials. Notwithstanding, most of the stimuli responsive biomaterials were explored largely in diagnostic applications and

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Fig. 9.1 Concept of stimuli responsive biomaterials for effective tissue regeneration and functional recovery in combination of physical/chemical/biological stimuli

on demand delivery of drugs, protein, gene, and cell (Cabane et al. 2012), which are not within the scope of the present chapter.

9.2

Stimuli Responsive Biomaterials in Tissue Engineering

9.2.1

Electroactive Biomaterials

Bioelectricity holds a pivotal role in our body’s normal operation including movement, thinking, sensation, visualization with eyes, blood transportation through our circulatory system and healing of an injury (Ghasemi-Mobarakeh et al. 2011). Cargo

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phenomena including movement of ions through the plasma membranes and electrons along biomolecules regulate all the biological processes in the body. Electrical potentials (60 mV to 100 mV) exist inside and outside cells. The changes in the transmembrane potential influence cellular functions as depicted in Fig. 9.2 (Qian et al. 2019). Biological tissues, particularly heart, neural, skin, bone, and muscles, are used to regulate their physiological behaviors and to propagate electrical potential by means of their electrical conductivity mechanisms such as accumulation and flow of charge (Balint et al. 2014). Electrical activities are associated in modulation of range of molecular events in these tissues, engaged in the development, adaptation, repair, and regeneration of tissues. There are growing evidences of significant positive contribution of ES in a range of important biological processes relevant to TE, viz. angiogenesis, cell division, cell signaling, nerve sprouting, prenatal development, and wound healing (Balint et al. 2013). This inspired the development of electroactive biomaterials, because of their excellent contact with bioelectric fields in cells and tissues, for a faster pace than traditional non-conductive biomaterials, for improving regenerations, differentiation or function of both in vitro and in vivo. Electroactive biomaterials enable cells to obtain direct electrical, electrochemical, and electromechanical stimulation. Possible clinical uses of ES include wound care, bone regeneration, nervous repair, and ulcer care of the diabetic and bedridden patients with pressure sores. Some of the electroactive biomaterials were clinically translated as non-biodegradable cardiac pacemakers, cochlear implants, electrodes for deep brain stimulation, etc. These smart biomaterials simultaneously can be stimulatory to the tissues as well as can trigger controlled/responsive release of therapeutics loaded into them. Such systems offer an effective delivery method for physicians and scientists in wound care, making it easier for patients to implement

Fig. 9.2 Scheme of cellular response elicited by electrical stimulation (ES) through electroactive biomaterials based scaffolds for improved tissue regeneration and function (Qian et al. 2019)

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new therapeutic approaches. Electroactive biomaterials can adapt their chemical, electrical, and physical properties to the specific needs of their application. The electroactive biomaterials family includes conducting polymers (CPs), piezoelectrics, electrets, and photovoltaics, which are discussed in the following subsections.

9.2.1.1 Conducting Polymers CPs are the latest class of organic polymers integrating the electrical, magnetic, and optical properties of metals and inorganic semiconductors with conventional polymers’ mechanical properties, processability, etc. (Shimano and MacDiamid 2001). This fourth generation polymers are completely distinct structurally from conventional polymers or mixture of insulating polymer with a conductive material such as a metal or carbon powder. Alternating single and double bonds along the strongly conjugated backbone of CPs enable electron mobility and charge movement within and between polymer chains, which results in strong electrical conductivity (Shirakawa et al. 1977). While the electrical conductivities of insulating polymers are much weaker (1020–106 S/cm), CPs possess much greater conductivities in the range of 1–103 S/cm (Le et al. 2017). Essentially, electrical conductivity in CPs is aided by two important features, which are its intrinsic conjugated alternation of single and double bonds and doping (Heeger 2001). Fundamentally, the electronic configuration CP’s backbone is unique from other insulating polymers due to the former’s conjugated alternating single-double carbon-carbon or carbon-nitrogen bonds (Skotheim et al. 1997). The CP backbone contains a strongly localized “sigma” (σ) bond and a weakly localized “pi” (π) bond with sp2 hybridized carbon atom. This sp2 hybridized carbon atom with a single s and two p orbitals, facilitates one non-boned electron (π electron) as shown in Fig. 9.3. Electron delocalization occurs due to the formation of π-band by the overlapping of the unpaired out-ofplane pz orbitals. Usually, two of the 2p orbitals (px and py) hybridize with 2s orbital to form three sp2 hybridized orbitals leaving one pz orbital unhybridized (Fig. 9.3). These sp2 hybridized orbitals are arranged at an angle of 120 among them in a same plane, while the unhybridized orbital remains perpendicular to the plane. The head-

Fig. 9.3 Formation of σ and π molecular orbitals from two sp2 hybridized carbon atoms in conducting polymers (CPs)

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on overlapping of the hybridized orbitals gives rise to strong σ (sigma) bonds contributing to the polymer chain configuration. On the other hand, the unhybridized pz orbitals of two carbon atoms undergo sideways overlapping and form π (pi) bonds. The electron cloud in the π bond are highly delocalized, which enables charge mobility along the polymer chain and between the neighboring chains. Therefore, the charge delocalization in π-band has vital role in defining the semiconducting or sometimes, metal like electrically conductive nature of CPs. Besides the single-double bond alteration in CPs, they are naturally non-conducting. Doping is the second essential requirement to impart electrical conductivity in CPs, which can be done by using anionic or cationic chemical species. However, the doping mechanisms in CPs are unique. Contrary to the substitutional doping in inorganic semiconductors, the process of doping in CPs is interstitial (Macdiarmid et al. 1985). Doping in CPs is nothing more than a charge transfer reaction, resulting in the partial reversible oxidation or less often reduction of the polymer. Doping can modify an insulating or semi-conducting polymer into a polymer with conductivity in the metallic regime. During doping, the loosely organized electrons hop along the polymer chain in the conjugated network. The peculiar conjugation of bonds in CP’s backbone allows the electrons to delocalize, culminating them being shared by several atoms. The delocalized electrons, therefore, serve as charge carriers, which render conductivity. Actually, such delocalization of charge modifies the band structure of CPs creating localized defects such as polarons, bipolarons, solitons, and defect bands. When electrons are extracted or added from a polymer chain, cations or anions are formed. These cations or anions under the influence of an electrical field can hop from one position to another leading to higher electrical conductivity. 9.2.1.1.1 Conducting Polymers in Tissue Engineering The increasing popularity of electrical and electromagnetic stimulation in medical field stems from the perception of the inherent bioelectric features of body tissues. Living tissues create electromotive forces, preserve the necessary potential difference, and turn the current on and off by regulating current flow and storing charge (Ghasemi-Mobarakeh et al. 2011). With this understanding, application of ES externally was well explored to various cellular activities including cell adhesion (Li et al. 2017), proliferation (Enayati et al. 2020), cell migration (Tai et al. 2018), and protein synthesis for tissue regeneration (Wake et al. 2011). The utilization of electrical signals to regulate the local cell microenvironment is, therefore, essential in activating specific cell behavior to particular phenotypes in order to achieve tissue functionality for longer run. CP-based biomaterials bring outstanding scaffolding features by assisting ES to cells, which are needed to promote regenerating mechanisms in the case of specific stimuli responsive cells (i.e., neurons, myotubes, cardiomyocytes) (Balint et al. 2014). CPs have many benefits in terms of excellent extent and period regulation of electrical stimuli, formidable electrical and optical properties, a high conductivity/weight ratio, and the ability to catch and controllably release biological molecules through reversible doping, to alter charges from a biochemical reaction and to easily alter their electrical, chemical, physical, and

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other properties necessary for intended application. In addition, CPs can be rendered biocompatibility, biodegradability, and porosity, which can be further altered and regulated even after synthesis by stimulation (e.g., electricity, light, pH) or various chemical based material modification techniques. Thus, several CPs such as polypyrrole (PPy), polyaniline (PAni), poly(3,4-ethylenedioxythiophene) (PEDOT), polythiophene (PT), and poly(2-methoxy-5-(2-ethylhexyloxy)-1,4phenylenevinylene) (MEH-PPV) were shown to effect positively various cellular activities including cell adhesion, proliferation and migration, DNA synthesis and protein secretion both in vitro and in vivo. Given the potential advantages, CPs were explored for various TE applications including neural, cardiac, bone, muscle, and wound healing, which are summarized in Table 9.1.

9.2.1.2 Piezoelectric Material Piezoelectricity refers to the phenomenon of surface charge accumulation on a material exhibiting a net dipole moment and no center of symmetry under a mechanical stress, which was first discovered by Pierre and Jacques Curie in 1880 (Jacob et al. 2018). Materials displaying such property are termed as piezoelectric materials. This unique category of materials can convert mechanical energy acting on it into electrical energy and vice versa. The generation of transient surface charges in presence of mechanical deformation (e.g. compression, tension) is known as direct piezoelectric effect and the deformation due to externally applied electrical signal (e.g. applied voltage, reversed polarity) is known as indirect or converse piezoelectric effect, as shown in Fig. 9.4 (Tandon et al. 2018). They can be categorized as piezoelectric polymers and piezoelectric ceramics, which may be natural materials or hydrogel systems. The dipoles are randomly oriented in a piezoelectric material and in order to fully utilize its piezoelectric feature, the dipoles should be rearranged so as to yield a net electric dipole moment through a dipole alignment process, called poling. This can be achieved by application of a strong electric field at a temperature above the glass transition temperature of the material followed by cooling under the same electric field. 9.2.1.2.1 Piezoelectric Materials in Tissue Engineering The piezoelectric property has gained significant attention in the evolving TE strategies to provide in vivo microenvironment for enhanced cell-biomaterial interaction and modulating the cellular response towards desired tissue or organ regeneration. Mechanical deformation induced transient electrical stimuli within piezoelectric biomaterial makes it one of the best approaches in delivering ES to cells without any external power source and devising any external electrical connections. Tissues like bone, cartilage, tendon, dentin, and keratin, have the piezoelectric property. Mostly, all these tissues composed with collagen, it is the fibril structure responsible for piezoelectric behavior (Halperin et al. 2004). The significance of piezoelectric behavior on cell behavior, tissue regeneration, and remodeling was explored, which has driven the research towards the development of novel piezoelectric biomaterials for TE.

In situ polymerization of PEDOT in chemically crosslinked Alg matrix followed by freeze drying

Cardiac TE

Solvent casting method

Poly (3,4-ethylenedioxythiophene)/ alginate (PEDOT/Alg)

Neural TE

Electrospinning

Cardiac TE

• Electrically conductive PAni/PGS films offered. • Good attachment, growth and proliferation of C2C12 myoblasts, while invoking no harmful effect on cells through its acidic leachants. • Macroporous PEDOT/Alg scaffolds supported good attachment and proliferation of adipose derived stem cells (ADSCs). • Under ES through these conductive scaffolds promoted cardiomyogenic differentiation of ADSCs.

Neural TE

Electrochemical polymerization

(continued)

Yang et al. (2020)

Qazi et al. (2014)

Borah et al. (2018)

Molino et al. (2018)

Yan et al. (2016)

Poly (3,4-ethylenedioxythiophene) (PEDOT) Poly[2-methoxy-5-(2-ethylhexyloxy)-1,4-phenylene vinylene]/polycaprolactone (MEH-PPV/PCL) nanofibers Polyaniline/poly(glycerolsebacate) (PAni/PGS)

Neural TE

Polymerization-enhanced ball milling method

Aligned polypyrrole/graphene (PPy/G) nanofibers

References Chen et al. (2019)

Fabrication technique Electrospinning and electrochemical deposition

CP-based biomaterial Graphene oxide/polypyrrole/ poly-L-lactic acid (GO/PPy/ PLLA)

Outcome • GO/PPy/PLLA conduit in conjunction with ES successfully repaired 10 mm rat sciatic nerve defect with improved re-innervated gastrocnemius muscle and nerve conduction velocity. • In addition, myelin sheath thickness and axon diameter in GO/PPy/PLLA conduit with ES was comparable to autograft. • Aligned PPy/G nanofibers with ES supported enhanced viability, neurite outgrowth and anti-aging ability of retinal ganglion cells (RGCs) suggesting possibilities for regeneration of optic nerve via ES on these electroconductive nanofibers. • ES through PEDOT significantly improved viability, morphology and neural differentiation of PC12 cells. • MEH-PPV/PCL nanofibers with ES offered significant enhancement in neurite formation and neurite outgrowth of PC12 cells.

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Table 9.1 TE applications of various CP-based biomaterials Application Neural TE

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Muscle TE

Muscle TE

Lyophilization and oxidative polymerization

Electrospinning

Directional lyophilization

Ink-jet printing

Polypyrrole/alginate/chitosan (PPy/Alg/Ch)

Polyaniline/polyacrylonitrile (PAni/PAN) nanofibers

Polypyrrole/collagen/ chondroitin sulfate (PPy/Col/ CS) Poly(3,4-ethylenedioxythiophene): polystyrenesulfonate/gelatin (PEDOT:PSS/ Gel)

Muscle TE

Bone TE

Bone TE

Freeze drying

Poly (3,4-ethylenedioxythiophene: poly(4-styrene sulfonate) (PEDOT:PSS)

Application Cardiac TE

Fabrication technique In situ polymerization of PPy over nanopatterned silk Fibroin films fabricated by capillary force lithography technique

CP-based biomaterial Silk fibroin/polypyrrole (SF/ PPy)

Table 9.1 (continued) Outcome • Nanopatterned SF/PPy scaffolds mimicking the native myocardial ECM topography, maintained viability of cardiomyocytes for 21 days leading to increased cellular organization and sarcomere development with upregulated expression and polarization of connexin 43, a critical regulator of cell-cell electrical coupling. • Osteogenic precursor cells differentiated into osteogenic phenotype on porous PEDOT:PSS scaffolds with elevated expression of bone regeneration associated genes. • The electrically conductive porous scaffolds also facilitated cell infiltration, increased ECM mineralization, and osteocalcin deposition. • PPy/Alg/CS scaffolds were cytocompatible as assessed with MG-63 cells and facilitated biomineralization. • PAni/PAN electrospun nanofibrous showed higher proliferation of primary myosatellite cells and myogenic differentiation as compared to PAN nanofibers. • Aligned and 3D PPy/Col/CS scaffolds provided guided myoblast growth and organization with enhanced myotube formation and maturation. • PEDOT:PSS/Gel scaffold demonstrated good metabolic activity, adhesion, differentiation, and alignment C2C12 myoblast cells than those on pure Gel scaffold. • The conductive scaffold along with ES promoted cell alignment and enhance myotubes differentiation. Fortunato et al. (2018)

Basurto et al. (2019)

Hosseinzadeh et al. (2016)

Sajesh et al. (2013)

Guex et al. (2017)

References Tsui et al. (2018)

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Photo-polymerization and sol-gel technique for PHEMA hydrogel formation followed by oxidative polymerization of PPy

Oxidative polymerization of PPy followed by sol-gel and solvent casting technique

Sol-gel technique

Poly(2-hydroxyethyl methacrylate)/polypyrrole (PHEMA/PPy) hydrogel

Polypyrrole/poly(L-lactic acid) (PPy/PLLA) conductive membranes

Chitosan/polyaniline/poly (ethylene glycol)-co-poly (glycerol-sebacate) (Ch/PAni/ PEGS) hydrogel

Wound healing

Wound healing

Wound healing

• The conductive hydrogel was found to be superior to the commercial Hydrosorb@ dressing in terms of anti-bacterial activity and protein absorption. • In vitro ES through the hydrogel promoted fibroblast migration, while faster healing was observed in rat diabetic wound model with in vivo ES. • ES through the PPy/PLLA conductive membranes to primary human fibroblasts demonstrated upregulation of various genes associated with cell adhesion, remodeling and spreading, cytoskeletal activity, extracellular matrix metabolism, while repressed production of inflammatory cytokines/ chemokines and improved growth factor secretion and signal transduction. • The electroactive and self-healable hydrogels showed good free radical scavenging capacity, biocompatibility, and anti-bacterial activity. • The hydrogel demonstrated promotion of tissue granulation thickness and collagen deposition in a full thickness skin defect model with enhanced healing efficacy and blood clotting capacity as compared to commercial dressing through elevation of various growth factor associated genes. Zhao et al. (2017)

Park et al. (2015)

Lu et al. (2019)

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Fig. 9.4 Mechanisms of direct and converse piezoelectric effect. [Redrawn from Tandon et al. 2018]

All the tissues in human body are subjected to mechanical stimuli, and the mechanical forces via gated channels responsible for the activation of signaling cascades augments for tissue repair and regeneration. The conversion of mechanical stimuli into biological signal, called as mechanotransduction, is exerted in physiological functions like muscle and bone homeostasis, regulation of blood flow, respiratory and kidney systems. The mechanical forces including compression, torsion, tension, and shear stress exerted on cells cause changes in voltage and ion concentrations, which result in change in gene expression. Several membrane associated molecules such as cell junction molecules, ion channels, G-Protein coupled receptors, and cytoskeleton proteins are involved under mechanical stimuli and initiate the biological response by activation of signaling cascades (Zaszczynska et al. 2020). In particular, ion channels contribute for piezoelectric response. The cationic channels including monovalent (Na+ and K+) and divalent (Ca+2 and Mg+2) channels are activated immediately after activation of piezo channels. The advances in the material science, physiology, and stimuli responses in tissue repair and regeneration has developed the strategies to develop synthetic, natural or composite biomaterials, which are appropriate to facilitate the physical niche to stimulate the cell proliferation and differentiation. The inherent piezoelectric property exerted in various tissues of human body augmented to develop a variety of composite biomaterials having piezoelectricity and tested for their suitability in various tissue engineering applications depicted in Table 9.2. The characteristic features of ideal biomaterials for specific tissue application also have been considered while developing the piezoelectric biomaterials. For example, the mechanical property varies depending on the tissue type. The mechanical strength of the composites modulated with enhanced piezoelectric property is the adopted strategy for the development of piezoelectric biomaterials for bone TE application. Likewise, the low mechanical

Two-photon lithography

Electrospinning

Electrospinning

Non-solvent induced phase separation method

Electrospinning

Poly(3-hydroxybutyric acid-co-3hydroxy valericacid) (PHBV)- BT composite scaffold

PVDF-trifluoro ethylene (TrFE) scaffolds

PVDF/graphene oxide (GO)

Gold nanoparticles/PVDF

Fabrication technique Scaffolds synthesized by electrospinning at different voltages (12–30 kV) Press sintering

Ormocomp-BT nanoparticles composite scaffold

Hydroxyapatite (HA)-barium titanate (BT) composite implant

Piezoelectric biomaterial Polyvinylidine fluoride (PVDF)

Table 9.2 TE applications of various piezoelectric biomaterials.

Neural TE

Neural TE

Neural TE

Cartilage TE

Bone TE

Bone TE

Application Bone TE

Outcome Higher alkaline phosphatase activity and mineralization was observed on PVDF25 kV scaffolds. Bone formation was noticed on the implant surface, exhibited direction dependent growth. Piezoelectric and topographic cues improved the bone regeneration, herein BT nanoparticles induced piezoelectric cues. Improved chondrocyte activity, gene expression of collagen-II higher. Piezoelectric cues supports cartilage regeneration. Human neural progenitor stem cells differentiated into β-III tubulin cells and enhanced neurite extension exhibited in micron-aligned- annealed scaffolds. GO addition improved piezoelectric and mechanical properties supported cell adhesion, proliferation and differentiation of PC12 cell. This scaffold served as nerve conduit channel and stimulated cell function. Addition of au nanoparticles improved piezoelectric properties in the composite fibrous scaffold and supported enhanced cell adhesion and growth.

Motamedi et al. (2017)

Abzan et al. (2019)

Lee and Arinzeh (2012)

References Damaraju et al. (2013) Jianqing et al. (1997) Marino et al. (2015) Jacob et al. (2019)

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strength is required for soft tissue, wherein the attainment of improved piezoelectric property in the composite material is the vital factor to be considered. The piezoelectric ceramics include barium titanate (BaTiO3), lead zirconate titanate (PZT), and lead metaniobate have already been studied for biomedical applications, while toxicity and brittle nature of these materials limited their application in biomedical field (Nguyen et al. 2014). Piezoelectric polymers have gained the attention over piezo ceramics owing to the biocompatibility, easy fabrication, tunable mechanical properties, etc. and therefore, found extensive applications in various TE areas such as bone, cartilage, neural, etc.

9.2.1.3 Electrets Unlike the transient surface charges in piezoelectric materials, electrets are dielectrics possessing quasi-permanent electric charges or molecular dipoles capable to generate electric fields within and outside. The concept of electrets was first proposed by Oliver Heaviside in 1885, while the first electret was first fabricated by Mototaro Eguchi in 1919 (Mascarenhas 1980). Electrets are considered as electrostatic equivalent of a permanent magnet owing to their ability to store charges for extended periods of time. Depending on the situation, however, the amount of charges decays over time. The electrets fabrication process is similar to the poling process of piezoelectric materials. For that, a dielectric material is electrically polarized by applying a high electric field and heating to softening temperature followed by cooling to room temperature. While maintaining the same field strength (Goswami and Sen 2018). The externally applied high electric field induces ordered charge accumulation inside the dielectric substrate as shown in Fig. 9.5. The induced charge accumulation process involves displacement of internal and external charges, which ultimately get trapped inside and prevents internal charge relaxation resulting in prolonged electrization. Figure 9.5 depicts the four ways of electric polarization of a dielectric material to form electrets according to Kohlrausch (Jefimenko and Walker 1980). He asserted that polarization due to alignment of molecular dipoles in the dielectric is more stable than the polarization due to internal charge migration to surface or various layers within the dielectric and atomic charge migration to the opposite ends of the molecules in the dielectric. Examples of electrets include organic materials such as ebonite, naphthalene, polymethyl-methacrylate, and many polymers, and inorganic materials such as sulfur, quartz, glasses, steatite, and some ceramics. 9.2.1.3.1 Electrets in Tissue Engineering The role of electret based materials in TE has gained considerable attention due to the ability of delivering ES to tissues without the need of external power source as in the case of piezoelectric materials. However, electrets have a static charge storage mechanism in contrast to dynamic charge generation in piezoelectric materials, which offers prolonged stability of the electret effect. The electret state has been used as a basis for understanding membranes, neural signals, biological memory in regeneration, electrically mediated tissue growth, and other phenomena in different biophysical models. Now more than 50 years of knowledge of bioelectrets, electret

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Fig. 9.5 Electric polarization in dielectric as suggested by Kohlrausch through (a) internal charge migration to dielectric surface, (b) charge migration within different layers of dielectric, (c) charge migration at molecular level, and (d) orientation of molecular diploes within the dielectric. [Redrawn from Goswami and Sen 2018]

effect was found in various biologically important molecules or polymers, viz., proteins, polysaccharides, polynucleotides, collagen, hemoglobin, DNA, and chitin (Mascarenhas 1980). The electret effect was observed in hydroxyapatite (HA), which forms about 60–70% of the bone mass of humans and animals. HA is thought to modulate bone formation and resorption, as well as promotes the regeneration of endothelial tissue (Bauer 2011). Depending on the amount of surface charges retained, electret based materials may deliver specific electrical signals to the tissue, giving rise to electrostatic fields and microcurrents to facilitate tissue regeneration processes. Therefore, various electret based biomaterials including natural and synthetic polymers were explored for range of TE applications such as bone, skin, artificial muscles and neural nerve, which are summarized in Table 9.3.

9.2.1.4 Photovoltaics Photovoltaic material is another class of electroactive materials, which can convert solar energy into electrical energy through photovoltaic effect and was demonstrated first in 1839 by Edmond Becquerel (Goetzberger et al. 2003). A photovoltaic material, semi-conducting in nature with two regions, namely n-type and p-type separated by pn junction (Fig. 9.6), is able to absorb a large spectrum of solar energy. Upon light absorption, electron–hole pairs are created. They migrate towards opposite directions towards each other and reach the pn junction, where an electric field is

Bone TE

Wound healing

Hydrothermal & Freeze drying method; grid controlled corona charging at 8 kV

Commercial nanocrystalline HAP/TCP; Corona poling at 5 kV

Lypholization; poling at 4 kV

Chitosan/hydroxyapatite (Ch/HA) nanocomposites

Hydroxyapatite/β-tricalcium phosphate (HA/TCP) nanocomposites Hydroxyapatite/silk fibroin (HA/SF) composite

Bone TE

Neural TE

Solution casting method (films) and lyophilization (channels); Corona poling at 8, 20 and 24 kV

Poly(lactic-co-glycolic) (PLGA)

Application Neural TE

Fabrication technique Extrusion based method; Corona poling at 14 kV

Electret based biomaterial Polytetrafluoroethylene (PTFE)

Table 9.3 TE applications of various electret based biomaterials

• Accelerated closure of a full thickness wound in porcine with poled HAP/SF gel. • Poled HAP/SF gel showed enhanced wound healing, re-epithelization, and matrix formation than the unpoled and pure SF gel. • Poled HAP/SF promoted maturation of fibroblast cells.

Outcome • After 4 weeks of implantation in a 4 mm mice sciatic nerve gap model, the cable area, blood vessel area and myelinated axons were significantly more in the regenerated nerves on positively and negatively charged PTFE tubes as compared to the uncharged PTFE tubes (diameter ¼ 0.9 mm). • PTFE tubes elicited minimal immune response. • Enhanced neurite outgrowth in mouse neuroblastoma cells grown on poled PLGA film compared to unpoled control film. • PLGA guidance channels with outer diameter 4 mm and internal diameter 2 mm, were implanted in 1 cm rat sciatic nerve gap models, which after 4 weeks demonstrated poled channels displayed regenerated nerves with greater conduction velocity and numbers of axons as compared to the unpoled guidance channel. • Improved primary rat cranial osteoblasts adhesion, proliferation, and differentiation capacity on composite electret membranes when compared to those on the uncharged membranes. • Improved osteoblast-cell adhesion, proliferation, and ECM formation on negatively poled nanocomposites.

Okabayashi et al. (2009)

Tarafder et al. (2011)

Qu et al. (2014)

Bryan et al. (2004)

References Valentini et al. (1989)

322 R. Borah et al.

Polypropylene/5-fluorouracil (PP/5-FU) patches

Gridcontrolled corona charging of PP film of thickness 13 μm at 15 kV; 5-FU patches were fabricated over the poled PP film

Wound healing

• In vitro scar permeation study showed PP/5-FU patch promoted higher permeation and retention of 5-FU through and in scar skin for hypertrophic scar (HS) inhibition. • In vivo study demonstrated significant reduction in collagen type I, collagen type III, TGF-β1 and HSP47 when PP/5-FU patches were applied onto the wound. Yuan et al. (2018)

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Fig. 9.6 Photovoltaic mechanism depicting light mediated migration of electron–hole pairs to opposite polarities in a traditional photovoltaic cell leading to production of electric current. [Redrawn from Tandon et al. 2018]

generated (Fig. 9.6). Photovoltaic devices consist usually of composite mixtures of semiconductor nanoparticles with conjugated polymers, wherein one component acts as electron donor and the other as electron acceptor (Goetzberger and Hebling 2000).

9.2.1.4.1 Photovoltaic Materials in Tissue Engineering Various semi-conducting materials showing photovoltaic activity are found to possess important features of an ideal biomaterial and hence, emerging TE strategies also include photovoltaic biomaterials for providing ES for tissue regeneration. The light absorption generated electric field, as described above, modulates the bioelectrical environment of cells or tissue, which controls ion influx processes through the plasma membrane. In a particular report, it was stated that the generated electric field induces Ca2+ ion translocation through voltage-gated calcium channels, which upregulates of cystolic Ca2+ leading to elevated activation of calmodulin (Jin et al. 2011). The elevated activation of calmodulin drives the nucleotide synthesis and cell proliferation. Photovoltaic polymer poly-3-hexyl-thiophene (P3HT) with phenyl-C61-butyricacid-methyl ester (PCBM) was assessed successfully to light mediated ES of neuronal activity of primary hippocampal neurons (Ghezzi et al. 2011). Similarly, P3HT based photovoltaic implants were reported to stimulate action potentials in explanted rat retinas (Ghezzi et al. 2013) and embryonic chick retinas (Gautam et al. 2014) through photoelectric stimulation. The light induced electrical energy generation was demonstrated by subcutaneous implantation of commercially available nonresorbable solar cells for powering pacemakers in vivo (Haeberlin et al. 2014, 2015). Subsequently, a bioresorbable and biocompatible silicon and magnesium based thin film solar cell was demonstrated for in vivo power supply (Kang et al. 2015). A group of researchers of USA in a breakthrough attempt used photovoltaic subretinal implants with 70 μm pixels for localized ES of retinal neurons when illuminated by near-infrared light (Lorach et al. 2015). This paved away the potential of photovoltaic biomaterials for stimulation of other tissues. However, photovoltaic biomaterials as TE scaffolds were scarcely explored for regeneration of nerve, bone, skin, and wound healing. Some of the interesting studies involving photovoltaicsbased biomaterials are summarized in Table 9.4.

Photolithography

Electrospinning

Photolithography

Poly(3-hexylthiophene)/Polycaprolactone (P3HT/PCL)

Monocrystalline silicon (Si)

Patterning technique

Fabrication technique Spin coating

Silicon (Si) microcell

Photovoltaics-based biomaterial β-Carotene/N,N0 -dioctyl-3,4,9,10perylenedicarboximide (β-carotene/ PTCDI-C8) and poly(3-hexylthiophene)/ phenyl-C61-butyric acid methyl ester (P3HT/PCBM) Poly(3-hexylthiophene) (P3HT) and the phenyl-C61-butyric acid methyl ester (PCBM) based organic photovoltaic patch

Wireless power supply for implantable medical devices

Skin TE

Bone TE

Wound healing

Application Neural TE

Table 9.4 TE applications of various photovoltaics-based biomaterials Outcome • The fabricated photovoltaics devices were able to provide NIR light induced electric field of 220–980 mV/mm. • Enhanced neurite extension by 64% and also effected direction of extension. • The disposable photovoltaic patches delivered visible light induced ES to skin wound in mice. • In vivo study showed that the patch promoted. • Cutaneous wound healing via enhanced hostinductive cell proliferation, cytokine secretion, andprotein synthesis. • Visible light induced photocurrent was successfully used to stimulate the intracellular calcium transients in osteoblast cells. • Under light induced ES, P3HT/PCL nanofibers demonstrated enhanced proliferation and healthier morphology of human dermal fibroblasts. • The Si based photovoltaic energy harvesting device was bioresorbable and induce no immune response, while capable to generate 60 μW in vivo.

Lu et al. (2018)

VargasEstevez et al. (2018) Jin et al. (2011)

Jang et al. (2018)

References Hsiao et al. (2016)

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9.2.1.5 Carbon Based Nanomaterials Carbon based nanomaterials possess highest electrical conductivity in the family of electroactive materials. Based on their structures, carbon based nanomaterials can be 0D (fullerenes, particulate diamonds, and carbon blacks), 1D (carbon nanotubes (CNTs), carbon nanofibers (CNFs) and diamond nanorods), 2D (graphene, graphite sheets, and diamond nanoplatelets), and 3D (nanocrystalline diamond (NCD) films, nanostructured diamond-like carbon (DLC) films, and fullerite (Lin et al. 2016). Among all, CNTs and graphene are the most attractive carbon allotropes for various technological applications due to their unique mechanical, thermal, and exceptional electrical properties. Graphene with single layer of a polycyclic aromatic hydrocarbon network sheet is the basic structural origin of other carbon allotropes, where sp2 hybridized carbon atoms are arranged in a honeycomb grid sheet (Fig. 9.7). Three of the four outermost valence electrons (2 s, 2px, 2py, and 2pz orbitals) in carbon atoms form covalent bonds with three neighboring carbon atoms, while the remaining electron in pz orbital (perpendicular to the sheet) forms pi (p) bond through sideways overlapping, which is highly mobile and this gives rise to high electrical conductivity (Wang and Weng 2018). Graphene sheet can be rolled up into a hollow cylindrical structure to get 1D CNT with the hexagonally arranged carbon atoms remains unchanged. The electrical conductivity of graphene and CNTs are comparable to the metallic conductors such as silver and copper, which are of order 107.

Fig. 9.7 Structures of different carbon based nanomaterials as indicated. [Redrawn from LloydHughes and Jeon 2012]

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9.2.1.5.1 Carbon Based Nanomaterials in Tissue Engineering Carbon based nanobiomaterials have some unique characteristics in regard to their potential use as TE scaffold. These include their size resembling with several biological components such as collagen, ultrahigh mechanical strength, and electrical conductivity. Nanotopography and electrical conductivity of CNTs mimic the native ECM (Eivazzadeh-Keihan et al. 2019). Carbon based nanobiomaterials offer the strongest material on earth till date and hence, they can be used for development of mechanically robust and durable biomaterial scaffolds. Thus, TE applications of carbon based nanobiomaterials are focused on exploring their mechanical strength and stiffness, high electrical conductivity, and complex physical properties. For instance, nanotopography and stiffness of carbon based nanobiomaterials are capable of modulating cellular activities including cell adhesion, proliferation, migration and differentiation. Likewise, these nanobiomaterials induce favorable cellbiomaterial interactions owing to their intrinsic electrical conductivity and were shown to boost cellular communication among electrically excitable cells such as neurons (Huang et al. 2012). Moreover, they can be modified with desired functional groups or molecules to improve desired cell-biomaterial interactions and also be tethered with other natural/synthetic biomaterials to boost their biocompatibility, biodegradability, bioactivity for TE applications. Researchers across the world explored various techniques such as coating, hydrogel blending, wet/dry-spinning procedures, and 3D printing to make 2D or 3D carbon nanobiomaterials based scaffolds for wound healing, neural, cardiac, bone, and cartilage TE. Few salient studies of carbon based nanobiomaterials in diverse TE areas are summarized in Table 9.5.

9.2.2

Magnetoresponsive Biomaterials

Similar to ES, magnetic stimulation (MS) has proved to positively effect biological functions at cellular and molecular level (Qian et al. 2019). Pulsed MS induces increased blood flow in capillary bed, serum ceruloplasmin expression, and improves angiogenesis. It has been established that MS effects ion influx through plasma membrane, various important protein and growth factor synthesis/secretion related to tissue regeneration (Fig. 9.8) (Qian et al. 2019). For example, low level electromagnetic field was shown to modulate cellular activities by influencing ionic transport across cellular membrane and action potential (Lacy-hulbert et al. 1998). Another study showed increased intracellular calcium concentration mediated tissue regeneration (Grassi et al. 2004). Magneto-responsive biomaterials contains active magnetic component within biomaterial network that can be manipulated spatiotemporally via an external magnetic field. This class of smart materials rely mostly on composites constituted by magnetic particles whose size allows them to become embedded into a polymer matrix to confer a magnetic response. The magneto-responsive behavior of scaffolds is especially controlled with magnetic nanoparticles of iron (Fe), nickel (Ni), cobalt (Co), and their oxides having a size less than 100 nm. The incorporation of magnetic

Cardiac TE

Bone TE

Dielectrophoresis and UV cross-linking

Template assisted method

Electrospinning of polyacrylonitrile (PAN) followed by carbonization at a 1000  C

Carbon nanotube/gelatin methacryloyl (CNT/GelMA) hydrogels

Polydimethylsiloxane/ multiwall carbon nanotubes (PDMS/ MWCNTs) 3D composites Carbon nanofibers (CNFs)

Cardiac TE

Neural TE

Electrospinning

Silk fibroin/reduced graphene oxide (SF/rGO) microfibers

Neural TE

Application Neural TE

UV cross-linking

Fabrication technique Chemical vapor deposition; ropelike structure with a diameter of 1 mm was prepared

Polycaprolactone fumarate/carbon nanotubes (PCLF/CNTs)

Carbon nanomaterial based biomaterial Carbon nanotubes (CNTs) ropes

Table 9.5 TE applications of various carbon based nanobiomaterials Outcome • As a viable substrate, CNT rope supported neural stem cell (NSC) growth and neurite outgrowth occurred favorably in the direction of the spiral topography on the CNT rope. • Electrical stimulation through CNT ropes accelerated the neurite outgrowth and early differentiation of NSCs into mature neurons. • Enhanced PC12 cell proliferation, neural differentiation, neurite outgrowth, cell migration, and intracellular connections on PCLF/CNT sheets upon ES (100 mV/mm and 20 Hz for 2 h/day). • SF/rGO microfibers supported PC12 cell viability and adhesion. • Electrical stimulation through SF/rGO promoted faster neural differentiation than those obtained by using nerve growth factor (NGF). • Aligned CNT/GelMA hydrogels offered enhanced the cardiac differentiation of the mouse embryoid bodies (EBs) compared with the pure GelMA and GelMA-random CNT hydrogels. • EBs activity was further enhanced under application of ES through aligned CNT/GelMA hydrogels. • 3D PDMS/MWCNT scaffold exhibited mechanical and conductive properties similar to the native heart muscle. • Further, it provided a suitable environment for enhanced viability, structural, and electrophysiological maturation, and proliferation of cardiomyocytes. • The fabricated electrospun CNFs is cytocompatible and suitable for cell culture and proliferation. • ES significantly enhanced the proliferation and the osteogenic activity of the bone cells.

Samadian et al. (2020)

Martinelli et al. (2018)

Ahadian et al. (2016)

AznarCervantes et al. (2017)

Zhou et al. (2018)

References Huang et al. (2012)

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Freeze drying

Electrospinning

Electrospinning

Reduced graphene oxide/ chitosan/silk fibroin (rGO/Ch/SF)

Chitosan/poly(vinyl alcohol)/graphene oxide (Ch/PVA/GO) composite nanofibers

Polycaprolactone/gelatin/ multi-walled carbon nanotubes (PCL/Ge/ MWCNTs)

Spider silk/carbon nanotubes (silk/CNTs)

Graphene oxide (GO) synthesized by modified Hummers method followed by deposition over cellulose paper Electrospinning

Graphene/cellulose (G/C) scaffold

Cartilage TE

Cartilage TE

Wound healing

Wound healing

Bone TE

• G/C electrodes possessed lower impedance and higher charge injection capacity than gold (au) electrodes, with high stability. • G/C scaffolds combined with ES supported enhanced ADSC proliferation, mineral deposition and ALP. • Expression compared to control samples without ES. • Silk/CNTs electrospun fibers combined with ES demonstrated elevated activity of diabetic dermal fibroblasts (DDF) for enhanced production of collagen with low COLI/ COLIII ratio and inhibited synthesis of matrix metalloproteinases (MMPs) leading to accelerated wound healing. • rGO/Ch/SF scaffold demonstrated radical scavenging ability, intracellular anti-oxidant activity in vitro and in vivo. • ES through rGO/Ch/SF offered improved adhesion and proliferation of C2C12 cells. • rGO/Ch/SF demonstrated improved in vivo wound regeneration. • Incorporation of GO increased the tensile strength of the nanofibers. • Ch/PVA/GO nanofibers promoted growth of mouse chondrogenic cells indicating the potential for cartilage TE applications. • Addition of MWNTs led to an increase in the hydrophilicity and tensile strength of the electrospun nanofibers along with the bioactivity. • Offered enhanced viability of adult chondrocytes. Zadehnajar et al. (2020)

Cao et al. (2017)

Tang et al. (2019)

Chi et al. (2019)

Li et al. (2020)

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Fig. 9.8 Scheme of cellular response elicited by magnetic stimulation (MS) through electroactive biomaterials based scaffolds for improved tissue regeneration and function (Qian et al. 2019)

nanoparticles in cells/tissues/scaffolds allows for magnetic force-based manipulation of these components to build more complex systems. In addition, integration of magnetic nanoparticles in scaffolds followed by the application of tensile or compressive forces using a magnetic field has been shown to induce functionality in certain cells. In contrast to ES, MS enables actuation at a distance on nanoscale and cell level. Furthermore, the magnetic field can penetrate deep into tissues, reaching a single cell and acting directly on its organelles; unlike the electric field, which is shielded by the membrane potential. For these reasons MS is gaining importance and intensively investigated in applications including tissue regeneration, targeted drug delivery, cancer therapy agent, etc. Among these different possible applications, this article mainly emphasizes on the applications of magneto-responsive scaffolds for different types of TE including bone, cardiac, cartilage, neural, etc. Different magneto-responsive scaffolds were prominently investigated in recent years for TE due to its ability to deliver direct mechanical stimulation to individual cells. Scaffolds based on the biological components such as bacterial cellulose, chitosan, or silk fibroins were proven to enhance cell proliferation, and differentiation under appropriate MS. These scaffolds not only provide a biocompatible environment for cell growth but also trigger desire cellular activities under MS. Hydroxyapatite (HA) due to its excellent biological activity, good biocompatibility, and bone conductivity has been considered as an obvious choice for bone replacement material. HA-based magnetic composites have also been investigated for bone (Torgbo and Sukyai 2019), cartilage (Huang et al. 2018) TE as well as for growth of human mesenchymal stem cells (D’Amora et al. 2017). Studies show that the combined effect of HA-based magnetic substrate and magnetic field exposure enhances cell proliferation, cell viability, and stimulates gene expression. In addition, magneto-active three-dimensional (3D) porous scaffold possessing a proper bone mimicking morphology has also been investigated for the adhesion and proliferation of preosteoblasts. It has been found that the application of magnetic stimuli increases the cell viability on the scaffolds, inducing a solid spiderlike network of cells, with the growth of cells on the scaffolds (Fernandes et al. 2019).

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Similar studies have also been conducted by developing 3D matrix of collagen hydrogel with magnetic nanoparticles to promote neural growth and cartilage TE. Investigation shows that these magnetically responsive 3D scaffolds can effectively induce the growth of neural cells and directed to form neural networks. Table 9.6 summarizes some of the recent findings in the field of TE using magneto-responsive scaffolds.

9.2.3

Thermoresponsive Biomaterials

One of the important/notable physical stimuli of which a relatively broad variation in its intensity can be withstood by body physiology is temperature (Doberenz et al. 2020). Interestingly, a small variation in temperature is able to cause alteration in size or shape of a unique class of materials, mainly polymers, which are known as thermoresponsive materials or polymers (Cabane et al. 2012). These polymers undergo a change in their miscibility or solubility at a critical temperature through a dramatic transition in the hydrophobic and hydrophilic interactions between their chains and the aqueous media (Cardoso et al. 2017). It leads to the dislocation of intra- and intermolecular hydrophobic and electrostatic interactions, causing the polymer chains to collapse, shrink, or expand. Intermolecular forces such as hydrogen bonding and hydrophobic forces in aqueous solution play a major role in the formation of micelle, hydrogel shrinking, and the physical cross-linking of thermoresponsive polymers. At critical temperature, thermoresponsive polymers change from monophasic (become completely soluble) to biphasic or vice versa. Thermoresponsive polymers, which dissolve completely to become monophasic above the critical temperature and show a phase separation below the critical temperature, are classified as thermoresponsive polymers with upper critical solution temperature (UCST). While polymers, which exhibit opposite behavior are regarded as thermoresponsive polymers with lower critical solution temperature (LCST). Another class of thermoresponsive polymers has been reported, known as thermally induced shape-memory polymers (SMPs) with non-UCST and non-LCST features but undergo changes in their shape and size under temperature fluctuations (Kim and Matsunaga 2017). Some common examples of thermoresponsive polymers are poly (N-isopropylacrylamide) (PNiPAAm), poly(N-vinylcaprolactam) (PNVC), poly (2-oxazoline)s (POxs), poly (L-lactic acid)-poly(ethylene glycol)-poly(L-lactic acid) (PLLA-PEG-PLLA), poly(ethylene oxide)-poly(propylene oxide)-poly (ethylene oxide) (PEO–PPO–PEO), etc. Applications of thermoresponsive polymers in TE applications are motivated by their thermally induced hydrophobic/hydrophilic properties to induce controlled cell attachment and detachment (Nagase et al. 2018). Compatibility of thermoresponsive polymers in TE is encouraged by another important fact that there is no harmful effect on cells and proteins over a temperature variation of 0–42  C (Doberenz et al. 2020). PNiPAAm is the most widely investigated thermoresponsive biomaterial with LCST behavior at 32 C, which is close to physiological condition (Yamada et al. 1990). PNiPAAm was explored as coating on cell culture dishes for collecting

Bone TE

Cartilage TE

Cartilage TE

Co-precipitation method followed by ultrasonic irradiation Co-precipitation synthesis of magnetic nanoparticles followed by electro-gelation

Co-precipitation

Ultrasonic dispersion freezethawing cross-linking molding process

Bacterial cellulose/ Fe3O4/hydroxyapatite Silk fibrion/Fe3O4

Collagen/hyaluronic acid/polyethylene glycol/magnetic nanoparticles Poly(vinyl alcohol)/ nano hydroxyapatite/ Fe2O3 nanoparticles

Bone TE

Neural TE

Application Neural TE

Embedding magnetic particles in collagen followed by solidification under magnetic field

Fabrication technique Oxidative hydrolysis synthesis of magnetic nanoparticles followed by lyophilization and mixing

Collagen hydrogel/ magnetic nanoparticles

Magnetoresponsive biomaterial Chitosan/ glycerophosphate/iron oxide nanoparticles

Table 9.6 TE applications of various magnetoresponsive biomaterials

• BMSCs show uniform growth on the surface of the magnetic nanocomposite hydrogel and high rates of proliferation. • BMSC growth is also enhanced by the addition of Fe2O3 and also significant stimulated chondrocyterelated gene expression.

Outcome • Nanocomposites able to support cell adhesion and spreading and further promote proliferation of SCs under magnetic field exposure. • Moreover, a magnetic field applied through the scaffold significantly increases the gene expression and protein secretion. • The magnetic elements have aggregated into magnetic particle strings along the magnetic lines within the gel. • These lines served as physical cues for neurons that developed in close proximity to the particles, leading to elongated and directed growth pattern. • Biocompatible and promote osteoblast attachment and proliferation. • Physical conjugation of basic fibroblast growth factor (bFGF) to Fe3O4 nanoparticles significantly enhanced the viability and growth of SaOS-2 cells on the scaffold. • Both human serum albumin coating and bFGF conjugation improves alkaline phosphate activity, total protein synthesis, and collagen synthesis. • The synthesized matrix exhibits similar microstructure and chemistry as hyaline cartilage and is cytocompatible with BMSCs in vitro after 24 h of culture period.

Huang et al. (2018)

Zhang et al. (2015)

Torgbo and Sukyai (2019) Karahaliloglu et al. (2017)

AntmanPassig and Shefi (2016)

Reference Liu et al. (2014)

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seeded cells and layers of cells just by regulating the temperature without using enzymes like trypsin. Traditional enzymatic degradation methods for cell separation reduce the cell function by affecting receptors, transport proteins and ECM and thus, integrity between confluent cell layers becomes weak leading to reduced efficiency for therapeutic applications. In the contrary, thermoresponsive biomaterials can provide intact cell sheet through non-enzymatic cell separation with retention cellular structure and function (Cooperstein et al. 2015). These intact cell sheets can be used as a fresh cell culture dish, applied to wound sites and host tissues, without requiring any sutures (Matsuda et al. 2007). Therefore, thermoresponsive biomaterials give spatial distribution of cells by layering sheets derived from various cell types or by layering monolayer cell sheets, creating 3D tissue constructs. Thermoresponsive biomaterials may be used as hydrogel, injectable gelling material, 3D printing or cell layer development by biomaterial surface modification.

9.2.4

Photoresponsive Biomaterials

Inspired by natural phenomena such as photosynthesis, researchers have been using light driven reactions to control biological functions and as a result, clinical implication of phototherapy using low level lasers, light-emitting diodes, and natural light, has increased in the last few years (Jin et al. 2011). Light, which is an electromagnetic radiation, is found to induce various regeneration associated molecular biology reactions such as increase in the cytosolic Ca2+ level in cells. Phototherapy has been proven to reduce inflammatory reactions, promote cell proliferation, and growth factor secretion (Desmet et al. 2006). Several researchers demonstrated light stimulation mediated accelerated wound healing, axonal regeneration, and spinal cord repair (Rochkind et al. 2002). These findings motivated scientists and researchers to explore photoresponsive biomaterials for various TE applications. Photoresponsive biomaterials, with light-sensitive molecules (chromophores) in them, when irradiated by light, are able to reversibly and frequently switch their physical and/or chemical properties, such as geometrical structure, refractive index, dielectric constant, conformation, solubility, and surface hydrophilicity, etc. in real time and spatiotemporal manner. Light stimulation through a photoresponsive biomaterial is a relatively straightforward, non-invasive technique to modulate dynamic cell microenvironment. Progress of such biomaterials in TE areas are summarized in this section. A photoresponsive culture surface composed of poly(N-isopropylacrylamide) (PNIPPAAM) with spiropyran chromophores as side chains was demonstrated to promote cell adhesion, when irradiated by ultraviolet (UV) light (wavelength: 365 nm) (Edahiro et al. 2005). Cells remained attached to the irradiated surface even after subsequent cooling and washing indicating better cell attachment due to UV irradiation. Acrylate based light-sensitive liquid crystalline elastomers (LCEs) were developed to assist cardiac muscle contraction (Ferrantini et al. 2019). The contraction was modulated in terms of light intensity, stimulation frequency, and time to on/off ratio in order to fit different contraction amplitude/time courses,

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including those of the human heart. Furthermore, LCE strips were successfully mounted in parallel with cardiac trabeculae, to improve muscular systolic function, with no impact on diastolic properties. Photoresponsive polysaccharide-based hydrogels obtained from radical polymerization was assessed for cartilage TE (Giammanco et al. 2016). These hydrogels become softer and more porous upon irradiation, presenting changes in their swelling and transport properties. Moreover, chondrogenic ATDC5 cells grown on the hydrogels showed a greater than two-fold increase in the production of sulfated glycosaminoglycans in the gels irradiated for 90 min compared to the dark controls. Poly(ethylene glycol) (PEG) hydrogel based micropatterned smart template was developed by spin coating method for culture of epithelial cells offering good cell adhesion and extended cell morphology (Gong et al. 2013). The study described the photoresponsive PEG hydrogel micropatterned smart template, which displayed transparency based photolithography to induce reversible control of cell adhesion with UV irradiation in defined areas. A 3D printable UV responsive cross-linking system based on polypeptides incorporating glutamic acid, isolycine, and nitrobenzene (NB) protected cysteine groups in a random and block copolymer was reported (Murphy et al. 2019). According to the report, the polypeptide with block architecture was more desired mechanical properties, gelled at lower concentration (3.0 wt %), and could easily deposit more than ten layered structures with high fidelity and resolution through 3D extrusion printing. In vitro cytotoxicity evaluated with human dermal fibroblasts cells revealed no toxic effect with fibroblasts.

9.2.5

Chemical Stimuli Responsive Biomaterials

Chemical stimuli responsive biomaterials respond to external chemical triggers such as pH, redox, and solvent. Since, these chemical stimuli are some important features of body physiology, chemical stimuli responsive biomaterials were also explored for various TE applications, which has been discussed in brief in this section. pH responsive materials contain ionizable groups for which they are able to accept or donate protons under any change in pH in the environment (Cardoso et al. 2017). Any pH change generate charges, which induces ionic interactions through electrostatic repulsion among them and ultimately causes physical or chemical changes in the material such as swelling, shrinking, dissociation, degradation, or membrane fusion and disruption (Gil and Hudson 2004). Researchers are motivated by the intrinsic pH variations present in living tissues to develop pH responsive biomaterial scaffold for various TE applications. For example, an injectable tissue scaffold based on branched nanofibers of peptide amphiphiles (PAs) with serine and histidine peptides conjugated to a single fatty acid tail, were shown to switch from solution state to hydrogel form at a pH above 6.5, which is within the physiological pH range (Lin et al. 2012). Another study demonstrated pH responsive C2-cyclohexane based low molecular weight hydrogels guided cell detachment with mild reduction in pH of the culture medium (Dou et al. 2012). Subsequently, a series of pH responsive tissue scaffolds composed of dimethylaminoethyl

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methacrylate (DMAEMA) and 2-hydroxyethyl methacrylate (HEMA) were shown to improve the oxygen and nutrient transport through expansion in response to a local pH change (You et al. 2015). The DMAEMA/HEMA composite scaffolds supported enhanced cell deposition and survival in vitro and subcutaneous implantation in rats showed upregulation of pro-healing genes indicating enhanced angiogenesis, granulation tissue formation, and tissue remodeling. Redox responsive materials possess redox sensitive group and they respond to any change in redox gradient of their surrounding environment by changing the oxidation state of the redox sensitive group (Cardoso et al. 2017). Application of redox responsive biomaterials in TE applications is inspired by the natural existence of redox potential in living tissues and glutathione/glutathione disulfide couple are the reducing agents available in abundance in animal cells. Redox responsive biomaterials under varying redox environment undergo changes in structure and shape. Therefore, TE applications of redox responsive biomaterials are mainly focused on utilizing the redox mediated degradation and drug/growth factor release properties. For instance, poly(ethylene glycol) (PEG) based cryogel containing disulfide-containing building blocks displayed the characteristics of a potential tissue scaffold such as biocompatibility and porosity (Dispinar et al. 2012). The cryogel demonstrated stability in physiological condition, but it degraded within few hours in presence of a reducing agent (glutathione), while the degraded by products did not affect cell viability. PEG based scaffold with redox mediated degradability and growth factor release features, was evaluated successfully in a rabbit radius critical defect for bone TE application (Yang et al. 2013). Same group also reported redox mediated degradable PEG based injectable hydrogel for bone regeneration (Yang et al. 2014).

9.2.6

Biological Stimuli Responsive Biomaterials

Biomaterials responsive to stimuli inherent to living tissues or cells are always advantageous. It is highly favorable for biomaterials to possess specific adaptive behavior in vivo. Alterations in conformation and degree of self-assembly of several important biomacromolecules in presence of specific chemical species in their surroundings, inspired scientists to develop innovative biomaterials that are responsive to biomacromolecules present in living systems. For that biomaterials are designed in such a way that it contains a functional group, which specifically interacts with biomacromolecules or sometimes, in conjugation with specific biological components. Although biological stimuli responsive biomaterials have not been studied extensively for TE applications, there are few evidences of using enzyme or glucose responsive biomaterials as potential tissue scaffolds. For example, an injectable self-healing hydrogel composed of phenylboronic acid and cis-diol modified PEG was demonstrated to release protein therapeutics in response to glucose, while also evoking no immune response in vivo (Yesilyurt et al. 2016). In another report, kartogenin, a chondrocyte differentiation inducing agent, was loaded into poly(lactic-co-glycolic acid)/hyaluronic acid (PLGA/HA) hydrogel for

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inducing differentiation of mesenchymal stem cells into chondrocytes (Shi et al. 2016). The kartogenin loaded hydrogel was demonstrated to play a major role in cell homing including recruitment of host’s endogenous cells in vivo without needing any cell transplantation. An injectable microporous annealed particle (MAP) gel based on PEG/vinyl sulphone for accelerated wound healing was demonstrated, wherein the microgel was cross-linked to cysteine-terminated matrix metalloprotease-sensitive peptide sequences for cell controlled biodegradability and resorption (Griffin et al. 2015).

9.3

Conclusions and Future Outlook

The current chapter provides a discussion on various types of stimuli responsive biomaterials in regard to their exploitation as potential tissue scaffolds with a special emphasize on physical stimuli responsive biomaterials such as electroactive and magnetoresponsive biomaterials. These biomaterials were explored extensively due to their potential to manipulate the intrinsic bioelectrical cues of the native tissue. TE is a more complex process and biomaterials are required to mimic the dynamic environment of the native tissue to support the natural regeneration processes. Hence, electroactive or magneto-responsive biomaterials discussed in the present chapter, have greater evidences as potential smart tissue scaffolds as compared to the chemical and biological stimuli responsive biomaterials including photoresponsive and thermoresponsive biomaterials. Moreover, CP and carbon based nanobiomaterials have emerged as superior smart biomaterial scaffolds among other electroactive biomaterials due to their intrinsic electrical conductivity, which is an important bioelectrical cue present in tissues and ES through such scaffolds were demonstrated for faster tissue regrowth and effective functional recovery both in vitro and in vivo. One of the major limitations of piezoelectric and electret based biomaterials is the requirement of poling the scaffold for dipole alignment sometimes for several hours above their glass transition temperature in presence of a high electric field of the order of kV (Shastri et al. 2000). It is only after the poling process for which piezoelectric and electret based biomaterials are usable for ES for finite length of time. Additionally, the electromagnetic signal, which is utilized by the systems such as photovoltaic, magnetoresponsive, and photoresponsive biomaterials, does not remain localized on the damaged area but gets penetrated to the surrounding areas of the injury site. In contrast, electroactive CP and carbon based nanobiomaterials offer focused ES with remarkable control over the level and duration of the stimulation. Chemical (pH and redox) and biological (glucose and enzyme) responsive biomaterials were scarcely explored as TE scaffold as compared to the formers. Besides the stimuli responsive feature, smart biomaterials should be flexible enough to be integrated with advanced biofabrication techniques such as photolithography, microcontact printing, 3D bioprinting, micromolding, and microfluidicassisted patterning etc., to be precisely mimic structure and other physical properties of the natural tissues (Mohamed et al. 2019). Future research should also focus to

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optimize the biophysical signal parameters within safe limit for living tissues to modulate the cell microenvironment. A successful technology to reach the end user needs to demonstrate robust clinical safety and efficacy for acquiring regulatory approval. Although, CP and carbon based biomaterials have demonstrated minimal immune response and biocompatibility, their one of the major constraints for use in TE is their non-degradability. Therefore, it is important to undertake strategies such as blending with natural or synthetic FDA approved other biomaterials to regulate the degradability feature. Acknowledgement RB gratefully acknowledges DST, Govt. of India, for the financial support through the Inspire Faculty Project (DST/INSPIRE/04/2018/000402). JU is extremely grateful to SERB for the financial support (ECR/2017/000628). BBR thankful to DBT, Government of India for financial assistance (BT/PR31908/MED/29/1401/2019) and LV Prasad Eye Institute, Hyderabad.

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Part III Applications of Biomaterials

Biomaterials for Hard Tissue Engineering: Concepts, Methods, and Applications

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Manju Saraswathy, Venkateshwaran Krishnaswami, and Deepu Damodharan Ragini

Abstract

Global tissue engineering market growth is expected to reach USD 28.9 billion by 2027 and witnesses a CAGR (compound annual growth rate) of 14.2% from 2020 to 2027. Rapid technological advancement in hard tissue engineering is expected to provide an effective solution for chronic conditions such as bone and joint disorder, severe injury, oral diseases, etc. Hard tissues engineering in the bones and teeth regeneration requires multiple contributing factors such as cells (embryonic stem cells, adult stem cells, induced pluripotent stem cells, fibroblast, etc.), smart biomaterial-based scaffolds (ceramics, polymers, composites, etc.), and growth factors (e.g., granulocyte colony-stimulating factor (G-CSF), interleukin (IL-8), tyrosine kinase-3, stromal cell-derived factor-1 (SDF-1), vascular endothelial growth factor (VEGF), angiopoietin-1 (ANG-1), macrophage inflammatory protein-2 (MIP-2), etc.). Advances in the development of biomaterials have provided attractive alternatives to hard tissue repair and replacement by regenerating tissues. Smart biomaterials provide exciting potential in hard tissue engineering by providing osteoinductive, osteoconductive, triggering/stimulating effects on cells and tissues to promote effective regeneration. In this chapter, we provide a brief description towards the importance of tissue engineering in the field of hard tissue regeneration, recent advances of biomaterials and strategies to fabricate biomaterials scaffolds for bone and tooth regeneration, significance of M. Saraswathy (*) · D. D. Ragini Division of Dental Products, Department of Biomaterial Science and Technology, Biomedical Technology Wing, Sree Chitra Tirunal Institute for Medical Sciences and Technology, Trivandrum, Kerala, India e-mail: [email protected] V. Krishnaswami Centre for Excellence in Nanobio Translational Research, Department of Pharmaceutical Technology, Anna University, Tiruchirappalli, Tamil Nadu, India # The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 B. Bhaskar et al. (eds.), Biomaterials in Tissue Engineering and Regenerative Medicine, https://doi.org/10.1007/978-981-16-0002-9_10

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3D bioprinting in hard tissue engineering, etc. Also, the chapter highlights current clinical trials in hard tissue engineering and portrays major challenges and future outlooks. Keywords

Bone regeneration · Tooth · Biomaterials · 3D bioprinting · Smart polymers · Shape memory polymers

10.1

Introduction

Hard tissue regeneration that used to repair and replace damaged tissues such as the bones, and the teeth, advances significantly with advances in multiple tissue engineering strategies. The hard tissues, also called calcified tissues contain unique cell types and composed of both inorganic and organic matrices. For example, the bone tissues consist of osteoblasts, bone lining cells, osteocytes, and osteoclasts (Zhang et al. 2018a). The organic phase of bone contains 90% of the collagenous proteins (type 1 collagen) and 10% of non-collagenous proteins (e.g., osteocalcin, osteonectin, osteopontin, fibronectin, etc). Whereas, the inorganic phase consists of phosphate and calcium ions in the form of hydroxyapatite crystals (Ca10(PO4)6 (OH)2)). The collagenous and the non-collagenous matrix proteins organize to form the scaffold for hydroxyapatite deposition and impart typical stiffness and resistance to bone tissue (Liu et al. 2016). On the other hand, the tooth is a highly complicated organ composed of both hard tissues and soft tissues with unique characteristics and functions. Hard tissues in tooth include the enamel, cementum, dentin, and alveolar bone (Yousef 2014). Major cells types involved in dental tissue formation is ameloblasts (form enamel), odontoblasts (form dentin), cementoblast (form cementum), osteoblast and osteoclast (form alveolar bone). Enamel is the hardest tissue in the human body that contains the highest percentage of minerals (96%) (Changyu et al. 2019). In contrast to enamel, dentin is soft and flexible and able to absorb energy and resist fracture. Cementum helps to cover the tooth root and provides proper attachment to the periodontal ligament. The structure and cell composition of hard tissues (e.g. bone and tooth) are depicted in Fig. 10.1. The concept of tissue engineering is based on the functional triad of cells, scaffolds, and biomolecules to induce cellular differentiation and tissue formation. This chapter is more focusing on biomaterials used for scaffold preparation in hard tissue engineering. In general, scaffold plays a major role in imparting mechanical support, and shape for tissue construction as seeded cells expand and organize. Also, the scaffold acts as a substitute for extracellular matrix, delivery vehicle for cells and growth factors, and triggers cell attachment, migration, and proliferation and thus affecting the efficacy of tissue regeneration (Pina et al. 2019). An ideal scaffold in tissue engineering should be biodegradable, biocompatible, bioactive, and, more importantly, should impart necessary stimulation for tissue regeneration (Baino et al. 2015). Biomaterials are the main component of tissue engineering scaffold that are

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Fig. 10.1 Structure of (a) natural bone (Gao et al. 2019), (b) natural tooth (Mourao et al. 2015)

available in multiple types (e.g., polymers, ceramics, and metals) (Sharma et al. 2014; Place et al. 2009). Functional requirements are key factors that determine the types of biomaterials to be used in a particular tissue engineering application. Special requirements of scaffolds in hard tissue engineering are mainly mechanical strength and porosity that demand the use of composite materials composed of both organic and inorganic components (Pina et al. 2019; Prasadh and Wong 2018). A multitude of fabrication techniques has been developed for these biomaterials to afford a range of potential shapes, size, porosity, and architecture in hard tissue engineering (Lee et al. 2018).

10.2

Biomaterials for Bone Tissue Engineering

Bone and associated diseases in people over 50 years old is a major clinical challenge to date. The younger generation shows a high regenerative capacity of bone fractures and heals without the need for major intervention (Marsell and Thomas 2011). However, large bone defects observed in severe fractures after accidents or bone tumor resections lack the template for proper regeneration and require surgical intervention (Tatullo et al. 2019). Among various clinical treatments available, the use of autografts is considered as the gold standard for bone repair and regeneration (Zheng et al. 2019; Brown and Cato 2020). Autologous transplantation involves the harvest of “donor” bone from a non-load-bearing site in the patient (typically an easily accessible site like the iliac crest) and transplant into the defect site. However, the use of autograft/allograft in hard tissue regeneration is restricted by its limited availability and donor site morbidity, infectious risk, and immune rejection (Gao et al. 2019; Oryan et al. 2014). Bone tissue engineering that induces tissue repair and regeneration via a natural mechanism of action is an effective alternative for autologous transplantation as it mimics the structure and properties as closely as possible (Chocholata et al. 2019). Bone tissue engineering mainly focuses on stem cells, growth factors, and scaffolds to enhance bone formation and repair. Besides, the establishment of sufficient vascular system is also crucial to satisfy the

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nutrient supplement and removal of waste during bone tissue regeneration (Abou Neel et al. 2014). Biocompatibility is the fundamental requirement for bone-tissue engineering biomaterial like any other tissue engineering concept. Surface characteristic of a biomaterial is an important factor that promotes adsorption of desired proteins in which cells can bind via receptor-mediated cell adhesion. The smart scaffolds that have bioactive peptides in it possess the inherent cell-binding capability. The interactions between cell receptors and proteins enable the deposition of ECM proteins and minerals within the scaffold that enhance the cell adhesion and differentiation. These materials can promote either orthotopic or ectopic bone formation. Additional material characteristics are required for ectopic bone formation compared to that of orthotopic bone formation as it involves the laying down of new bone material at the site where bone tissue would not otherwise present. Ectopic bond formation requires the osteoinductive properties in a material in which the material has the intrinsic capacity to stimulate osteogenesis via the recruitment and differentiation of stem cells into osteoblasts or pre-osteoblast that is the initial cellular phase of a bone-forming lineage (Miri et al. 2016; Kroese-Deutman et al. 2008). Naturally occurring osteoinductive materials include demineralized bone matrix (DBM) and specific bone morphogenetic proteins (BMPs) which form bone within the skeleton as well as extra skeletally (Pilipchuk et al. 2015). Whereas, the osteoconductive materials enable the deposition of mineralized tissue on their surface and thereby promote direct bonding to the bone (Alves et al. 2010). Osseointegration is the direct structural and functional connection between living bone and the surface of a loadbearing material in which new bone is laid down directly on the material surface to impart mechanical stability (Parithimarkalaignan and Padmanabhan 2013). In addition to bioactivity, hardness (mechanical properties) and biodegradation are the two most important characteristics for a bone substitute material. The bone substitute materials are required to withstand compressive loads experienced at the bone-forming site to prevent the collapse of the growing tissues. In general, the mechanical properties of these materials should preferably match those of native bone to avoid the stress shielding effect. Stress shielding can lead to a reduction in bone density called osteopenia harmonized with the Wolff’s law (Noyama et al. 2012; Elliott et al. 2016; Joshi et al. 2000). The higher mechanical strength of the material (compared to the surrounding tissue) leads to weakening of the healthy surrounding tissues as the majority of loading forces will be borne by those materials (e.g., metal implants). As known explicitly biodegradation is an integral part of tissue engineering. In an ideal condition, the rate of scaffold degradation should match the rate of mineralized tissue deposition, such that the gradual decrease in mechanical support provided by the degrading scaffold is compensated for the gradual increase in mechanical support provided by the new tissue (Bose et al. 2012; Matsumoto et al. 2019). However, the degradation products must also be biocompatible and must not alter the local environment such as local pH. Scaffold requirement of effective bone tissue regeneration is shown in Table 10.1. Another key player in bone substitute material is the porosity. Interconnected pores having a diameter ranging from 100–300μm facilitate infiltration of new blood

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Table 10.1 Scaffold requirement of effective bone tissue regeneration. Adapted with modification from Donnaloja et al. (2020) Properties Cytocompatibility

Biodegradability

Bioactivity

High porosity

Mechanical features Tunable properties Processability

Description The scaffold or its released products should not elicit inflammation or toxicity in vivo The degradation rate of the scaffold should match the rate of tissue regeneration by external-enzymatic/ biological process. Scaffold should resorb after fulfilling the purpose Scaffold should interact with the tissue according to osteoinductive and osteoconductive principles Interconnected pores induce cell adhesion, cell distribution, migration, and thereby enhance bone tissue ingrowth. In addition, increased surface area of porous scaffold provides site for the formation of chemical bond between the bioceramics and host bone. On the other hand, the porosity should not affect the mechanical stability Scaffold should reproduce elastic and fatigue strength of the bones tissue site Scaffold should have customizable properties. Easy manufacturing Scaffold should be easy to be fabricated and sterilized. Easy clinical manipulation is a key factor

References Chandra (2020), Bharadwaz and Jayasuriya (2020), Pina et al. (2015)

vessels, cell adhesion, differentiation, and migration throughout the construct. Although it decreases the overall mechanical properties of the scaffolds, pores are necessary for the entry and continued residence of cells, nerves, and blood vessels (Ostrowska et al. 2016; Hannink and Arts 2011). SEM images of interconnected porous structure of human trabecular bone and hydroxyapatite scaffold are shown in the figure (Fig. 10.2). Various types of bone substitute materials have been developed to fabricate scaffolds that include bioactive ceramics, bioactive glasses, polymers, and composites. Different types of scaffolds and strategies to enhance the treatment of bone tissue defects and diseases are shown in the figure (Fig. 10.3). In general, the choice of material for bone tissue regeneration depends on multiple factors.

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Fig. 10.2 (a) SEM images showing interconnected porous structure of human trabecular bone (b) SEM image of hydroxyapatite scaffold. Interconnected pores are clearly visible. Adapted from Doi et al. (2012)

Fig. 10.3 Different types of scaffolds (porous matrix, nano-fiber mesh, hydrogels, and microspheres) used to deliver bioactive molecules. This can be combined with a number of physicomechanical strategies to enhance treatment of various bone tissue defects and diseases. Adapted with permission from Yague et al. (2015)

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10.2.1 Polymers and Hydrogels Among different biomaterial types, polymers are more promising because of their unique characteristics such as cytocompatibility, biodegradability, flexibility of design of their blocks, and various other tunable functionalities such as crystallinity, degradation kinetics, chemical compositions, thermal transition, etc. (Asghari et al. 2017; Puppi et al. 2010). Both synthetic polymers and natural polymers are available abundantly in the fabrication of various scaffolds. Natural biodegradable polymers include collagen, gelatin, cellulose, hyaluronate, chitin, alginate, etc. (Zou et al. 2019; Akilbekova et al. 2018; Jazayeri et al. 2016). Whereas, synthetic polymers include poly(lactic acid): PLA, poly(glycolic acid):PGA, poly(lactic-co-glycolide): PLGA, poly(e-caprolactone):PCL, polyhydroxyalkanoates: PHA, etc. (Gunatillake and Raju 2003; Kumar et al. 2019; Ghassemi et al. 2018). For example, several bone tissue engineering products in the market contain collagen. Collagen is the main protein component of natural bone and contains amino acid sequences to which cells readily attach (Marques et al. 2019; Lin et al. 2019; Ferreira et al. 2012). Natural and synthetic polymeric materials suitable for bone tissue regeneration and their main characteristics are shown in the table (Table 10.2). However, the softness and low mechanical properties limit the use of polymeric scaffold in bone tissue regeneration (Liu et al. 2014). Crosslinking of polymers, either physical crosslinking or chemical crosslinking can address these issues at a larger extent. Chemical crosslinking enhances the mechanical properties and stability of the particular polymer system. Crosslinking networks are an important component of hydrogel systems that affect a wide range of scaffold properties such as mesh size, percentage swelling, and elasticity (Bai et al. 2018; Puppi et al. 2010; Chenxi et al. 2019). The three-dimensional hydrophilic network of hydrogel possesses mechanical strength, encapsulates bioactive molecule/cells, and can provide nutrient environments suitable for endogenous cell growth. Hydrogel formulated from natural polymers has several advantages in bone repair as it mimics natural ECM of the bone (Zhao et al. 2014; Yang et al. 2020). For example, collagen-based hydrogel was studied for the bone defect in the dorsal nasal bone of the rats (Lindsey et al. 1996). The study demonstrated that 6 weeks after the implantation a thin bone layer was formed on the surface of the defect. Another study delivered bone morphogenetic protein (BMP)-2 loaded hyaluronic acid (HA) gel to the cranial defect site of rats, and 75–100% of the BMP was released within the first 24 h (Patterson et al. 2010). HA gel BMP combination promoted higher bone formation in the defected area of rats than the treatment without HA gel. Dual network hydrogel structures were also studied in this regard, in which physically/chemically cross-linked anisotropic swimming bladder collagen fibril is the first network, and neutral, biocompatible poly (N, N0 - two methacrylamide) (PDMAAm) is the second network (Mredha et al. 2017). In vivo experiments show that the dual network hydrogel improved the stability of the gel and the strength of the binding to the bone. Kim et al. designed a bionic system for local delivery of drugs made from hyaluronic acid (HA) and vinyl phosphonic acid (VPAc) cross-linked biomineralized hydrogels (Kim and Jeong-Sook 2014). By

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Table 10.2 Natural and synthetic polymeric materials suitable for bone tissue regeneration and their main characteristics Scaffold Collagen

Gelatin

Silk fibroin

Chitosan

Alginate

Hyaluronic acid

Poly (caprolactone)

Advantages • Similar to ECM. • Enzymatic biodegradability. • Cytocompatibility and cellbinding properties. • Versatility in being processed in different physical forms such as microparticle and nanoparticle. • Possible injectability. • Capable of delivering other biological component • Cytocompatibility. • Biodegradability. • Osteoconductivity • Cytocompatibility. • Flexible processability. • High mechanical properties. • Ability to guide formation of hydroxyapatite • Cytocompatibility. • Biodegradability. • Cell-binding. • Differentiation and migration properties. • Antibacterial properties. • Mucoadhesivity • Cytocompatibility. • Easy gelling (ionic crosslinking with metal ions such as calcium) • Cytocompatibility. • Biodegradability. • Enzymatic biodegradability. • Viscoelasticity. • Easy manipulation. • Easy chemical functionalization • Biocompatibility. • Biodegradability. • High mechanical strength

Disadvantages • Low mechanical strength. • Difficulty in handling

References Bharadwaz and Jayasuriya (2020), Karadas et al. (2014), Pina et al. (2015)

• Poor mechanical properties. • Low stability in physiological conditions • Limited biological adhesion. • Immunogenicity

Barbani et al. (2012), Baheiraei et al. (2015)

Paşcu et al. (2013), Rockwood et al. (2011)

• Poor mechanical strength and stability. • Rapid in vivo degradation rate

Schwartz et al. (2011), Mirahmadi et al. (2013)

• Difficult to sterilize. • Low cell adhesion • Poor mechanical strength. • Very rapid degradation

Lee and Mooney (2012), Khosravizadeh et al. (2014) Schwartz et al. (2011), Lee et al. (2013), Shang et al. (2014)

• Reduced bioactivity. • Hydrophobicity. • Low cell adhesion. • Slow degradation rate

Bharadwaz and Jayasuriya (2020), Lee et al. (2013), Moeini et al. (2017)

(continued)

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Table 10.2 (continued) Scaffold Poly (L-lactic acid)

Advantages • Biocompatibility. • Thermal stability. • Biodegradability

PLGA

• Wide range of degradation rate based on the composition. • Tunability

Disadvantages • Reduced bioactivity. • Possible adverse tissue reaction for acid degradation product • Sub optimal mechanical properties. • Poor osteoconductivity

References Yang et al. (2005), Cacciotti et al. (2014)

Tahriri et al. (2016), Masaeli et al. (2016)

regulating the crosslinking density, mineralization degree, and ionic strength, the system could control the water content, degradation rate, speed of drug release, and could successfully deliver the protein drugs that would promote bone repair and regeneration. In all these studies resorbability and excellent integration potential of these hydrogels with surrounding tissues reduced the possibility of an inflammatory response and postsurgical complexities. Hydrogels formulated using synthetic polymers are another promising candidate for bone tissue regeneration. Unlike natural materials, synthetic polymers have basic structural units, so the properties of polymers (such as porosity, degradation time, and mechanical properties) can be adjusted for specific applications (Kretlow and Antonios 2007; Liu and Peter 2004). Hydrogels composed of poly(aldehyde guluronate) and adipic acid dihydrazide were studied as cell carriers to implant primary rat cranial osteoblasts into the backbone defect in mice. Nine weeks later, mineralized bone tissue formed at the defect (Lee et al. 2001). Synthetic polymers have extensive mechanical stiffness and controllable degradation rate. It is reported that the pendant cyclic ester modification of PCL can modulate the slow drug release. The degradation of amphiphilic PCL-PEG-PCL hydrogel resulted from the strong hydrophobicity and crystallinity of PCL segments (Wang et al. 2012). The composition of synthetic copolymers affects the structure and properties of the gels. In the preparation of poly (vinylphosphonic acid-co-acrylic acid) (PVPA-co-AA) bone graft substitute, increasing PVPA content generated hydrogels with great swelling capacities, high porosities, and adjustable mechanical and cell adhesion properties (Dey et al. 2018). Although synthetic materials have the above advantages, their success is limited by their own inherently poor biological activity, acid by-products, and other shortcomings. Therefore, synthetic materials can be conjugated with biological and chemical entities to improve the comprehensive properties of hydrogels. Another study grouped PEG hydrogels based on their physical/chemical modification (Thoma et al. 2017). Each group was implanted onto six loci of the rabbit skull. After 6 weeks of observation, they found that chemical and/or physical modification had a significant effect on PEG hydrogel

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matrix stability, degradation time, and integration into the surrounding soft tissues and hard tissues.

10.2.2 Hybrid Scaffolds in Bone Tissue Engineering Composite materials (hybrid materials) are a combination of organic and inorganic phases offer appropriate mechanical properties and flexibility to the scaffold. Inorganic-organic composites mimic the composite nature of real bone and during such combination; the alkalinity of the inorganic filler neutralizes acidic autocatalytic degradation of polymers such as PLA (Fragal et al. 2016; LewandowskaŁańcucka et al. 2015). Enhanced bone regeneration potential of inorganic biomaterials such as the bioactive glass (e.g., SiO2-CaO-P2O5) and bioactive ceramics or bioceramics (e.g., hydroxyapatite; Ca10(PO4)6(OH)2)) has well demonstrated in the literature with their high osteoconductive and bone-bonding ability (Gleeson et al. 2010; Kim et al. 2006; Yang et al. 2011). Hydroxyapatites (HAP) of different shape and size have been used extensively in bone tissue engineering because of its biological properties such as biocompatibility, osteoconductivity, and angiogenesis. HAP release calcium and phosphorus ions at the implant site and promote bone mineralization (Daculsi et al. 2003; Jung et al. 2010). A great assortment of techniques has been used for fabrication of porous scaffolds based on composites such as calcium phosphate/polymer, and bioactive glass nanoparticles (SiO2-CaO-P2O5)/polymer via solvent-casting and in situ precipitation. Other commonly used techniques include sol–gel processing, solid freeform fabrication, emulsion freeze-drying, porogen leaching, fiber bonding, gas foaming, electrospinning, microsphere sintering, phase separation, 3D-plotting technique, or a mixture of these techniques (Hutmacher and Cool 2007; Soundarya et al. 2018). Schematic representation of different fabrication techniques is illustrated in the figure (Fig. 10.4). Although these fabrication methods can produce highly porous scaffolds, these methods have limited control over scaffold architecture and pore interconnectivity. Numerous novel manufacturing techniques have been developed to process porous scaffolds with large void volumes. Each technique has particular advantages and disadvantages. For example, disadvantages include the use of toxic solvent and lack of uniform architecture and poor strength, while the advantages are ease of fabrication, superior structural strength, the ability to incorporate and deliver bioactive molecules. The selection of a scaffold fabrication technique is, therefore, a question of setting priorities to determine the vital requirements. HAP/collagen composite stimulates the formation of new bone tissue. Fukui et al. implanted composites consisting of nano-HAP and collagen into the mandibles of rabbits. The study demonstrated the formation of greater amount of newly formed bone tissue in the composite environment than in the control, and faster implants replacement with host bone tissue (Fukui et al. 2008). Composite material based on hydroxyapatite and collagen has been tested on dog tibia bones of 20 mm defect (Kikuchi et al. 2001). It was demonstrated that the newly formed bone tissue gradually penetrated and the defect was completely cured after 8 weeks. Both the

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Fig. 10.4 Common scaffold fabrication techniques (a) solvent casting-particle leaching process, (b) gas foaming, (c) freeze-drying, (d) phase separation, (e) electrospinning. Adapted from (Puppi et al. 2010; Shu et al. 2018)

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study confirmed the involvement of these composite in bone remodeling process. Recently three-dimensional, porous, human-like collagen (HLC)/nano-hydroxyapatite (n-HAP) scaffold cross-linked using diepoxyoctane demonstrated the excellent mechanical and superior biological properties for bone tissue regeneration (Liu et al. 2020).

10.3

Applications of Tissue Engineering in Dentistry

The socioeconomic need to treat/replace damaged or non-functional tissues with pioneering approaches, designs, and technologies forms the major challenge in dental research. Dental tissue regeneration includes regeneration of alveolar bone, periodontal ligament, enamel, dentin, and even the whole tooth (Abou et al. 2014; Tahriri et al. 2020). For the last few years interest in dental tissue engineering has significantly increased globally with multifactorial applications that focus more on biomaterials design/processing, surface characterization, and functionalization for improved cell–material interactions (Galler et al. 2012a, b). Cell injection, cell induction, and cell-seeded scaffold are the common tissue engineering strategies utilized for dentistry. Cell injection of intelligent/stem cells results in tissue formation and affords tissues to get regenerated (Ana et al. 2018). Challenges involved in cell injection include phenotype maintenance and immunological rejection (Volponi et al. 2010). Cell induction therapy recruits circulating body cells to regenerate dental tissue (Elham 2017). Osteoinduction of osteoconductive material regenerates damaged dental tissues. Injection of signaling molecules such as growth/differentiation factors (fibroblasts growth factors 2 and 9, transforming growth factors b1, vascular endothelial growth factors, recombinant human growth/differentiation factor-5, and bone morphogenetic protein) regenerates the damaged dental tissues. The cell-seeded scaffold strategy depends on the isolation of appropriate cell population from individual biopsy, especially mesenchymal stem cells (MSC). Human oral mucosa/gingiva-derived MSCs having potent immunomodulatory and antiinflammatory properties are a potential cell source for MSC-based therapies in dental research (Zhang et al. 2012). These isolated stem cells will be expanded in the lab culture and seeded into either natural or synthetic scaffold of choice (Jiarong et al. 2011). The cells adhere to the scaffold, proliferate, differentiate, and form particular tissue type. At the same time the scaffold breakdowns completely after fulfilling its intended function during the tissue formation.

10.3.1 Tooth Regeneration Tooth loss and tooth defects that occur due to various conditions including dental caries, periodontal disease, and genetic disorders are common clinical diseases in stomatology and affect individual vital activities and psychosocial well-being. Compared with the conservative treatment strategies, tooth regeneration has unique advantages as it involves natural mechanisms of tooth development and biological

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Fig. 10.5 Tissue engineering concepts of tooth regeneration (a) scaffold-based approach, (b) cell aggregation-based approach, (c) stem cell homing-based approach. Adapted with modification from Steindorff et al. (2014)

healing processes (Volponi et al. 2018). Recently whole tooth regeneration in vivo has become a hot topic in dental research. Tooth development is a complex process that consists of multiple tissues including the enamel, dentin, cementum, and pulp that involve multiple factors such as cytokines, growth factors, and other bioactive molecules. Different tissue engineering concepts are tried even in the tooth regeneration including scaffold-based approach, cell aggregation-based approach, and stem cell homing-based approach (Fig. 10.5). Bioengineered incisor tooth germ was successfully developed using cells isolated from the epithelium and mesenchyme of the dental germ (Nakao et al. 2007). Another study demonstrated the whole tooth regeneration in a swine model using early-stage tooth germs using gelatin– chrondroitin–hyaluronan-tri-copolymer scaffold (Wu et al. 2019). The study resulted in the formation of dentin/pulp-like complex structures after 36 weeks of implantation. However, the regenerated teeth were much smaller in size compared to that of the normal teeth. Considering the complexity of tooth structure, and opportunities associated with it, multiple research groups focus on developing aesthetic and morphologically acceptable tooth structures.

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10.3.2 Bone Regeneration in Dental Application Even though, bone has a strong capacity for self-repair, conditions such as complex trauma, tumor, infection, and congenital disorders that cause large bone defects and resorption demand for alternate methods. Autografts are the gold standard material for bone defect management. However, there are a lot of limitations associated with autografts including donor site morbidity and limited availability. The use of other conventional surgical methods like allografts, and xenografts are also having limitations that include immune rejection and potential transmission of infectious diseases. Because of the limitations inherent with conventional bone graft strategies, tissue engineering has become an area of promising approach for bone repair and regeneration. Bone tissue engineering strategies have great potential to address most of the current clinical need. The fundamental concept of bone tissue engineering is to combine stem cells seeded onto biocompatible 3D scaffolds with appropriate growth factors to produce functional bone structures. For example, functional regulation of osteoblast lineage cells in health and osteoporosis were studied and evaluated the effect of strontium and its role in regulating bone remodeling. Another study demonstrated the effect of nuclear factor-κB ligand (RANKL) in periodontal bone resorption and factors to regulate RANKL expression (Chen et al. 2014). The effect of interleukin10 (IL-10) in the treatment of bone loss diseases was evaluated and showed that IL-10 inhibits bone resorption and can be used as a potential therapeutic strategy in periodontitis and other bone loss diseases (Zhang et al. 2014). Optimizing and refining the use of scaffolds is another important aspect for bone tissue engineering. Natural materials such as polysaccharides (starch, alginate, chitin/chitosan, hyaluronic acid derivatives, soy, collagen, fibrin gels, silk) and synthetic polymers such as poly(lactic acid) (PLA), poly(glycolic acid) (PGA), and their copolymers, inorganic materials such as bioactive glasses, tricalcium phosphate, hydroxyapatite and their combinations are commonly used in cell transplantation and scaffolds for dental bone tissue engineering because of their superior mechanical properties and degradation rate control (Galler et al. 2010). Novel calcium phosphate cement containing gold nanoparticles were studied for its osteogenic induction ability on human dental pulp stem cells (Xia et al. 2018). The results showed that the incorporation of gold nanoparticles improved human dental pulp stem cells behavior on calcium phosphate cement with better cell adhesion (about twofold increase in cell spreading) and proliferation, and enhanced osteogenic differentiation. Multifunctional nanosized mesoporous bioactive glass/poly(lactic co-glycolic acid) composite-coated CaSiO3 scaffolds were studied for bone tissue regeneration that demonstrated improved mechanical strength, apatite-mineralization activity, cytocompatibility, and drug-delivery properties for bone tissue engineering application (Shi et al. 2014). Another study demonstrated the direct effect of substrate surface properties to cell adhesion on different biomaterials via evaluating the cellular responses of biomimetic calcium phosphate coatings and alkaline-treated titanium surfaces (Yu and Wei 2013). Scaffolds of different proportions of low and medium molecular weight alginates with and without hydroxyapatite nanoparticles for bone tissue regeneration were also studied (Barakat 2019). The results showed

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complete degradation of the scaffolds within 14–21 days depending on the composition. Healing properties of embroidered polycaprolactone-co-lactide (PCL) scaffolds coated with collagen/chondroitin sulfate were studied in sheep and demonstrated enhanced de novo bone formation at the proximal and distal ends of the tibia (Rentsch et al. 2014). The same is used as skull bone implants for large in vivo defects.

10.3.3 Enamel Regeneration Enamel is the outermost covering of teeth which constitutes 96% of minerals the uniquely organized nanostructured material (96% crystalline calcium phosphate) which is generated by ameloblasts (Beniash et al. 2019). The control, orientation, and elongated growth of enamel rods during the mineralization process supported by a factor called amelogenin. Ameloblastin, enamelin, and tuftelin proteins also control apatite nucleation and growth of enamel at different stages of amelogenesis. Enamel regeneration may be achieved by direct solution mineralization, protein/ peptide-induced mineralization, hydrogel driven mineralization, and precursor assembly (Shao et al. 2019). Recently development include a liquid solution comprising calcium phosphate ion clusters which effectively mineralize the external surface of damaged tooth enamel and afford protection. Shao et al. developed tooth enamel regeneration material composed of calcium phosphate ion clusters to act as a precursor layer to induce the epitaxial crystal growth of enamel apatite and to mimic the biomineralization of crystalline amorphous frontier of hard tissue. They emphasized that even though several materials reported for the restoration of tooth enamel, they responded temporarily only due to the imperfect combination of materials utilized as compared to that of the native enamel. The study insisted that the calcium phosphate ion clusters may integrate into native enamel thereby provides a permanent effect for enamel erosion.

10.3.4 Dentin and Dental Pulp Regeneration Dental caries remains one of the most prevalent in young adult and childhood diseases, while the phrase “root canal” is probably the most dreaded term in dentistry. There are several ways in which one can potentially engineer lost dentin and dental pulp. Recent studies reported that even if the odontoblasts (cells that produce dentin) are lost due to caries, it may be possible to induce formation of new cells from pulp tissue using certain BMPs (Kim et al. 2013; Wang et al. 2019; Kawashima and Takashi 2016). These new odontoblasts can synthesize new dentin. Tissue engineering of dental pulp itself may also be possible using cultured fibroblasts and synthetic polymer matrices. Further development and successful application of these strategies to regenerate dentin and dental pulp could 1 day revolutionize the treatment of our most common oral health problem including dental cavities.

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Regeneration of dental pulp tissue may induce blood into the root canal. Understanding on vascular network within the root canal is important for revascularization, or reestablishment of root development (Olcay et al. 2020; Rombouts et al. 2017). Regenerative endodontic procedures will be performed with disinfection process coupled with chemical irrigation as well as the stimulation of pulp-dentin regeneration through bleeding induction (Latham et al. 2016; Sismanoglu and Pinar 2020; Chanyong et al. 2018). Type I collagen derived from animal source and recombinant collagen are widely used biomaterials for tissue engineering for dental pulp regeneration (Prescott et al. 2008). Collagen scaffolds of native fibrillar forms and denaturized forms (sponges, sheets, plugs, and pellets) are used for dental pulp regeneration. During dental pulp regeneration, blood-derived growth factors such as platelet-derived growth factor, fibroblast growth factor, vascular endothelial growth factor, and insulin growth factor act as a dominant signaling protein supporting vasculogenesis and angiogenesis which may promote blood vessel formation enhancing neovascularization (Lin et al. 2019; Saghiri et al. 2015; Guo et al. 2020). Dental pulp-derived stem cells incorporated with bioactive factors in scaffold were studied to promote cell proliferation, differentiation, and angiogenesis using enzyme-cleavable hydrogel made from self-assembling peptide nanofibers to encapsulate dental pulp stem cells (Galler et al. 2012a, b). Dentin composed of hydroxyapatite, tricalcium phosphate, octacalcium phosphate, amorphous calcium phosphate, and dicalcium phosphate dehydrate. It has been found that the demineralized dentin matrix (DDM) consists of transforming growth factor β, bone morphogenetic proteins, vascular endothelial growth factor, fibroblast growth factor-2, platelet-derived growth factor, and insulin-like growth factor-1 (Pashley et al. 2003; Kang et al. 2017). DDM supports osteogenic and dentinogenic differentiation of bone mesenchymal stem cells. Bioactive molecules present in the dentin matrix supports dentin formation naturally especially for trauma and infection which may support the survival, apoptosis, and differentiation of human dental pulp stem cells.

10.4

Biomaterials Used in Dentistry

Biomaterials including both natural/synthetic play a vital role in dental tissue engineering (Sarang et al. 2014). The cellular microenvironment required for optimal dental tissue regeneration was obtained by biomaterial fabrication utilizing frameworks such as scaffolds, matrices, or constructs with interconnected pores. Tailor-made polymers with varying molecular weights are used in dental tissue regeneration to afford interconnected porosity, large surface area, adequate mechanical strengths, varying surface characterization, and varying geometries essential for dental tissue regeneration. The most prevalent structural protein found in the extracellular matrix of various connective tissues like bone, cartilage, tendon, muscle, skin, etc., is the collagen. Due to the structural and chemical similarities (cellular adhesion, cellular migration, and cell growth) collagen has been widely utilized in dental regeneration

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(Pankajakshan et al. 2020; Li et al. 2019). Osteodentin formation has been observed with type I collagen. Whereas, alginate exhibits low mechanical stiffness and uncontrolled in vivo degradation rate. Increasing calcium content and crosslinking density may enhance the mechanical strength of alginate (Sancilio et al. 2018). Dentin pulp and periodontal regeneration may be supported with alginate hydrogels (Moussa and Aparicio 2019; Yuan et al. 2011). Alginate also supports the delivery of growth factors like TGF β for tooth regeneration. The natural biomaterial fibrin (shrinkable/biodegradable/stiffer) obtained by polymerization of fibrinogen (De et al. 2020). Manipulating the fibrinogen/thrombin components the physical characteristics of fibrin-based scaffolds can be modified (De Melo et al. 2020). The biocompatible and low immunogenic hyaluronic acid which may enzymatically be degraded by hyaluronidase into nontoxic products is suitable for pulp regeneration, dental pulp proliferation, and invasion of vessels from amputated dental pulp (Inuyama et al. 2010; Alshehadat et al. 2016). The synthetic polymer polyethylene glycol offers advantages of non-toxicity, biocompatibility, low immunogenicity, and biodegradation. Elastic modulus of polyethylene glycol may be varied by altering the molecular weight/concentration of the polymer. Block copolymers of polyethylene glycol and PEGylated polymers are widely utilized for formation of vascularized tissue and dental tissue regeneration (Chang et al. 2017; Zein et al. 2019). The natural biopolymer chitosan is obtained from chitin. Chitosan possesses biocompatible, biodegradable, antimicrobial and possesses tissue healing and osteoinductive properties. Chitosan and its combination are effective for periodontal regeneration (Tang et al. 2020). Due to the diverse physical properties such as cell attachment and proliferation, silk is widely used for soft dental pulp formation (Zhang et al. 2019). Synthetic polyester polymers such as polylactic acid (PLA), polyglycolic acid (PGA), polylactide-co-glycolide (PLGA), inorganic calcium phosphate materials (hydroxyapatite or beta-tricalcium phosphate), and compositions of silicate and phosphate glasses at different grades are used for dental tissue regeneration which may support to generate tissues similar to dental pulp and dentin. The surfacemodified polymers are also utilized for cementum generation with suggested potential in periodontal regeneration. The application of ceramics in dental tissue regeneration is tremendous. Biocompatible ceramics in dental tissue regeneration include alumina, zirconia, bioactive glasses, glass ceramics, and bioresorbable calcium phosphates. For example, zirconia-based bioceramics promote cell proliferation, color compatibility with that of existing teeth, and differentiation in osteogenic pathways (Guarino et al. 2020). Bioactive glasses and glass-ceramics support osteogenic genes expression, stimulate neovascularization/angiogenesis, enzymatic activity, and differentiate mesenchymal stem cells (Cornelissen et al. 2017). The biocompatible, osteoconductive, and bioresorbable calcium phosphate which resembles with that of inorganic part of major normal/pathological calcified tissues majorly used for tooth replacement and maxillofacial reconstruction at the molar ratio of 0.5–2.0. The overall nanocomposite scaffold composed of poly (lactic-coglycolic acid)/nano bioactive glass-ceramic/cementum protein 1/chitin-poly (lacticco-glycolic acid)/fibroblast growth factor 2/platelet-rich plasma-derived growth

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factors is widely used for differentiation of human dental follicle stem cells (Sowmya et al. 2015; Rad et al. 2019).

10.5

Dental Stem Cells in Hard and Soft Tissue Engineering in Dentistry

Mesenchymal stem cells (MSCs) play an essential role in organ development and postnatal repair in tissue engineering (Zheng et al. 2019). A variety of studies demonstrated the potential of both endogenous and exogenous MSCs for bone and tooth regeneration. Common cell sources studied so far for the hard tissue regeneration are bone marrow derived MSCs, adipose-derived MSCs, and dental stem cells including dental pulp stem cells and periodontal ligament stem cells (Arigbede et al. 2012). However, adverse effect of diseased microenvironment creates challenges in establishing safe, effective, and simple stem cell-based approaches for tissue regeneration. Dental stem cells have the capacity to regrow tooth tissues (Miran et al. 2016). The dental pulp within the teeth also contains a type of stem cell termed as odontoblasts (Kawashima and Takashi 2016). Dental pulp stem cells naturally differentiate in to odontoblasts with dentin forming capacity. Dental pulp stems cells have high expression of CD90, CD73, CD105, CD29, CD13, and CD44 surface antigens, and lack the expression of hematopoietic markers. Here in the stem cells are undifferentiated cells which possess the capacity to renew themselves upon differentiation and development into different cell lineages. Stem cells may offer alveolar bone regeneration and solve large periodontal tissue defects thereby substitute the lost tooth. The pluripotent stem cells elicit its major role in dentistry for regenerative treatments due to its pluripotency and unlimited self-renewal capacity. Regeneration of periodontal tissue, salivary glands, missing jaw bone, and tooth loss has been achieved by pluripotent stem cells (Chankee et al. 2014). Dental pulp based induced pluripotent stem cells from adult human may form deciduous teeth which may get differentiated into neural crest cells supporting dental bone-based regeneration. Combinatorial effect of stem cells and enamel matrix derivatives may enhance periodontal regeneration and the cementum formation of the periodontal ligament and alveolar bone. Self-regeneration, healthy tissue differentiation, and damage tissues repair may be achieved by adult stem cells. Dental tissues stem cells of dental pulp and periodontal tissues hold the capacity to regenerate to form reparative dentine. Periosteum derived stem cells possess the ability to differentiate into adipocytes, osteoblasts, and chondrocytes. Mesenchymal stem cells have the capacity to produce mineralized tissues with complex structures similar to dentin, dental pulp, and periodontal ligament. Dental pulp banking has extreme richness in MSCs which may help stem cells based dental tissue regeneration. Technologies such as autologous bone grafts, cell-sheet technology, scaffold-free technology, scaffoldfree engineered cell-made constructs, etc., are also used for dental tissue regeneration. Dentinogenic peptide impregnated injectable hydrogel for the support of dental pulp stem cells was studied (Nguyen et al. 2018). The study demonstrated the potential application of self-assembled peptide for guided dentinogenesis.

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Advanced Tissue Engineering Strategies

10.6.1 3D Printing in Hard Tissue Engineering 3D printing technology has gained attraction in the field of bone regenerative medicine by overcoming engineering challenges of conventional tissue engineering strategies. It creates scaffolds in a layer by layer fashion from the 3D computer-aided design models (Fig. 10.6). Fabrication of 3D-printed scaffold with cells and growth factors is a great challenge that maintain cell survival and cell growth during or after printing process. Polycaprolactone and polycaprolactone/graphene scaffolds were fabricated using an extrusion-based additive manufacturing system. The graphene nanosheets were distributed uniformly and the scaffolds were coated with a P1-latex protein and seeded human adipose-derived stem cells. Results showed that scaffolds containing protein and graphene increased proliferation and cell adhesion (Caetano et al. 2018). Polydopamine modified polylactic acid-based 3D scaffold showed significant cell growth, cell adhesion and increased the secretion of the extracellular matrix (Kao et al. 2015). In another study a porous scaffold of polycaprolactone was fabricated using FDM technique (Dong et al. 2017). The rabbit bone marrow mesenchymal stem cells and bone morphogenetic protein-2 were encapsulated in the thermo-sensitive chitosan hydrogel and injected into a three-dimensional polycaprolactone scaffold. The scaffold with cells showed growth and cell proliferation, and increased osteogenesis. Subcutaneous implantation of MSCs seeded

Fig. 10.6 3D printing is a part of additive manufacturing technology used to create scaffolds in a layer by layer fashion from the 3D computer-aided design model (Kellner et al. 2014)

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scaffolds in nude mice was also performed. Results showed that the polycaprolactone/chitosan hybrid scaffold could repair bone damage. Recently, cryogenic 3D printing technique was used to fabricate dual delivery scaffold for improved bone regeneration using tricalcium phosphate and osteogenic peptide (OP) containing water/poly(lactic-co-glycolic acid)/dichloromethane emulsion inks (Wang et al. 2020). The scaffold showed highly porous structure that is mechanically similar to the human cancellous bone. Collagen I hydrogel containing angiogenic peptide (AP) was then coated on scaffold to impart angiogenic potential. It is demonstrated that a sequential release of both AP and OP. Tooth is one of the strongest and hardest tissues in the body, consisting of hard tissue such as enamel, dentin, cementum, alveolar bone, etc., along with other soft tissues. Personalization of restoration and replacement of damaged tissue in the treatment are important concerns. Continuous wear and tear in the teeth occur due to application of pressure in the mouth while performing personal routines and bacterial attack. The scaffold design in complex dental defects is a major challenge and 3D printing can contribute effectively to address the complexities in dental tissue regeneration. Polycaprolactone/mineral trioxide aggregate scaffold was fabricated using electrohydrodynamic jet 3D printing technique (Wu et al. 2016). 3D-printed polycaprolactone/β-tricalcium phosphate scaffolds were fabricated using a custom made 3D bioprinting unit for dental tissue regeneration (Park et al. 2017). Another study reported multiple 3D-printed structures with the release of active factors in order to imitate different structures of the tissue to enhance the tissue regeneration potential of these scaffolds. Dental pulp stem implanted subcutaneously in mice resulted in the formation of the collagen-I-rich tissue. Three-dimensional scaffold of hydroxyapatite and tricalcium phosphate was constructed by 3D printing technique and subsequently seeded the human dental pulp stem cells and stem cells from the apical papilla (Hilkens et al. 2017). 3D-printed scaffolds implanted in mice demonstrated the regeneration of vascularized pulp-like tissue and mineralized tissue formation in all stem cell constructs by histological and ultrastructural analysis.

10.6.2 3D Bioprinting in Hard Tissue Engineering The conventional 3D printing approach involves the predefined layered printing of scaffolds followed by cell seeding and perfusing the construct before implantation. However, this method suffers from a lack of uniform spatial and temporal distribution of cells and growth factors in the construct (Mandrycky et al. 2016; Zhang et al. 2017; Željka et al. 2019). As a result 3D bioprinting has evolved, which has precise control of biological architectures and mechanical properties (Kačarević et al. 2018). Bioink is a material that resembles the extracellular matrix in the form of hydrogels or viscous fluids containing cells encapsulated within the matrix (Pati et al. 2014; Shen et al. 2017). These bioinks are deposited with precision and polymerize or cross-link either before or after deposition to form the scaffold (Ji and Murat 2017). Biphasic calcium phosphate (BCPs) scaffolds were fabricated along with hydroxyapatite and

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tricalcium phosphate (TCP) as the ideal composition for the repair and replacement of significant bone defects. The achieved structural accuracy of the BCPs scaffolds was higher than 96.5%. A study reported the use of human nasal inferior turbinate tissue-derived mesenchymal stromal cells (hTMSCs) and demonstrated the interaction of 3D bioprinted scaffolds containing PCL, PLGA, and b-tricalcium phosphate (b-TCP) (Pati et al. 2015). Another study reported the effect of surface modification of 3D-printed PLA scaffold with cold plasma to enhanced cell attachment and proliferation of human fibroblasts, osteoblasts, and bone marrow mesenchymal stem cells (Wang et al. 2016). It was shown that under optimal CAP conditions, both hydrophilicity and nano-scale roughness played a significant role in enhancing the biological properties of the printed constructs for bone regeneration. The three main types of 3D bioprinting technologies are inkjet-based 3D bioprinting, laserbased 3D bioprinting, and extrusion-based 3D bioprinting. Inkjet-based 3D bioprinting approaches involve ink chambers with different nozzles corresponding to piezoelectric, thermal, and acoustic actuating units. In order to actuate the unit a short pulse of electric current is required in case of inkjet-based 3D bioprinting technology. The advantages associated with inkjet-based 3D bioprinting are this technique is fast, economical, and readily available with higher resolution. Herein the printing process involves jetting of fixed volume of fluid over the platform through ink droplets (Bishop et al. 2017; Kocak et al. 2020). Another study demonstrated the development of bioink using collagen type I or agarose along with sodium alginate and incorporated chondrocytes to construct in vitro 3D-printed cartilage tissue. They observed the improved mechanical strength of sodium alginate/collagen type I and sodium alginate/agarose groups compared to sodium alginate group. The study concluded that 3D bioprinted sodium alginate/collagen type I with favorable mechanical strength and biological functionality may act as a promising tool for cartilage tissue engineering (Yang and Huang 2019). Reinforced 3D-printed hydrogel constructs were developed for cartilage tissue engineering by formulating composite bioinks containing alginate and short sub-micron polylactide fiber. The study demonstrated that three-fold increase of young’s modulus up to 25.1  3.8 kPa may be obtained for pristine alginate constructs with the addition of polylactide short fibers. Further bioinks seeded with human chondrocytes and cultured in vitro for up to 14 days to assess the performance in cartilage tissue engineering. The study confirmed that the fabricated composites materials may represent a valid solution for tissue engineering applications (Kosik Kozioł et al. 2019). Nanocomposite hydrogel construct based on polyethylene glycol diacrylate/ laponite nanoclay/hyaluronic acid sodium salt bioinks was developed using two channel 3D bioprinting method (Kim et al. 2019). The bioink facilitates the 3D-bioprinting and enables the efficient delivery of oxygen and nutrients to growing cells. The construct exhibited excellent osteogenic ability in the long term, due to the release of bioactive ions (magnesium ions, Mg2+ and silicon ions, Si4+), which induces the suitable microenvironment to promote the differentiation and acts promising for bone tissue regeneration in terms of cell engraftment, survival, and long-term function (Zhai et al. 2018).

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The nozzle free high precision-based laser 3D bioprinting approaches involve direct writing using laser at the pulse involved in the localized heating to deposit a layer of energy absorbing biomaterial. Laser-based 3D bioprinting approaches have been used in cell patterning involving single cell and cell suspensions. In order to support the patterns of printed cells the laser-assisted bioprinting approaches utilize biological layers of hydrogel termed as biopapers. Self-assembled osseous sheets can serve as living biopapers to support the laser-assisted bioprinting of human endothelial cells and thereby guides the tubule like structure formation specific for autologous bone repair applications (Kawecki et al. 2018). Three-dimensional biological structures composed of living cells and hydrogel through the direct and accurate printing of cells with an inkjet printing system were demonstrated using alginate hydrogel obtained through the reaction of a sodium alginate with calcium chloride solution (Nishiyama et al. 2009). Another study demonstrated the sustained release of bone morphogenetic protein-2 (BMP-2) in bioprinted scaffolds for osteogenicity in mice and rat (Poldervaart et al. 2013). Gelatin microparticles were used as a delivery vehicle to facilitate the sustained release of BMP-2 and dispersed in hydrogel-based constructs. Osteogenic differentiation and bone formation were studied for the construct seeded with goat multipotent stromal cells. The study showed that released BMP-2 significantly enhanced the osteogenic differentiation both in vitro and in vivo. Recent study demonstrated the effect of 3D bioprinting of alginate/gelatin hydrogel scaffold on the proliferation and differentiation of human dental pulp stem cells (hDPSCs) (Yu et al. 2019). The study demonstrated osteogenic/odontoblastic differentiation of hDPSCs with the enhanced formation of bonelike nodules.

10.7

Shape Memory Polymers in Hard Tissue Engineering

Shape memory polymers have emerged as an excellent class of biomaterials in hard tissue engineering due to their excellent shape memory performance, and other tunable characteristics. Shape memory polymer scaffolds can change their shape in response to a specific stimulus such as temperature change, light, pH, and electric or magnetic field (Zhang et al. 2018b; Lendlein and Kelch 2005). Shape memory polymers shown multifactorial applications in hard tissue engineering include support materials for bone grafts, augmenting healing and possess the ability to conform to irregular-sized defects in bone (Montgomery et al. 2017). Scaffolds having shape memory behavior can be predesigned, deformed to be conveniently implanted into the target site (e.g., bone defects) via minimally invasive surgery and then retain its original shape to completely occupy the defects. The initial implant has a small size that can easily be deployed in the body using the minimally invasive technique with the least damage to host tissues. The presence of particular stimulus condition make the implant to regain larger shape to fill the bone defect precisely (An et al. 2015). Another study fabricated BMP2-loaded shape memory porous nanocomposite scaffold that consists of chemically cross-linked poly(ε-caprolactone) and hydroxyapatite nanoparticles for the repair of bone defects. The study showed that almost

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complete new bone formation and trabecula-like structures in the BMP2-loaded scaffold treated group in vivo (Lin et al. 2011). 4D-printed hierarchy scaffolds were created using a series of biocompatible shape memory polymers. Here 4D printing refers to the 3D printing of shape memory polymers (SMPs) having time-dependent shape transformation when exposed to environmental stimuli after printing (Miao et al. 2016). Recently, 4D shape memory polyurethane scaffolds were fabricated via 3D printing. For example, water-based biodegradable shape memory polyurethane was demonstrated as the main component of the 3D printing ink for fabricating bone scaffolds containing 500 ppm of superparamagnetic iron oxide nanoparticles that promote osteogenic induction and support to curb the shape (Wang et al. 2018). They concluded that shape memory properties eliciting 3D-printed polyurethane scaffolds with biodegradability, and osteogenic effect may be utilized for minimally invasive surgical procedures as customized-bone substitutes for bone tissue engineering. It has been shown that time-dependent morphing recovery of the 4D scaffold significantly elongates the cells and nuclei and thereby the activity of those cells could be effectively directed via multiple mechanical stimuli (Miao et al. 2016). As piezoelectric materials enhanced tissue formation by providing an electrically active microenvironment without the need for external power sources for electrical stimulation, efforts were made to develop shape memory polymers and scaffolds that can be triggered using electrical stimulus. A recent study demonstrated the fabrication of 3D fibrous scaffold of piezoelectric poly(vinylidene fluoridetrifluoroethylene) (PVDF-TrFE). The dynamic compression of the piezoelectric scaffold at 1 Hz frequency with 10 % deformation induced greater MSCs chondrogenic differentiation (Damaraju et al. 2017). Another study demonstrated an electrospun PVDF-TrFE fiber scaffold containing zinc oxide nanoparticles for the boosted adhesion and proliferation of human MSCs (hMSCs) and enhanced the blood vessel formation in a rat model (Augustine et al. 2017).

10.8

Tissue Engineering Challenges in Dentistry

Tissue engineering has much potential application in dentistry. However, there exist many challenges to practice in the clinical scenario. Even though there is tremendous literature available in the area of dental tissue engineering, very few tissueengineered products have reached in clinical trials. The main challenges in dental tissue engineering include the complexity of oral tissues and the lack of details of tissue-specific problems (Grawish et al. 2020). To exemplify the complexities of the oral tissues, enamel and dental pulp regeneration can be considered (Zafa et al. 2015). Scientists are unable to stimulate ameloblasts to secrete enamel tissue in vitro with structure and properties similar to that of natural enamel (Zafar et al. 2015). The vascularized network of dental pulp lacks collateral circulation hence immune system finds difficulty to eradicate infected cells (Nanci 2012) with the difference in the mechanical properties of dentin and predentin, poor rate of dentin formation (approximately 4μm/day for primary dentin and 1–3.5μm/day for reparative dentin)

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creates a greater challenge as it will take several months to regenerate 1 mm of dentin in vivo (Linde and Goldberg 1993). The challenging tasks of the mechanical properties (hardness, elastic modulus), physical properties (porosity, surface area), and biological properties (biocompatibility, degradation rate) of functional biomaterial scaffolds to that of natural dental tissue are a huge hurdle for dental researchers. Progressive degradation, stability, shelf life, immunogenicity, and delivery methods are also additional challenging tasks for dental tissue-engineered products. Even after fully understanding the structural and functional complexities of dental tissues, one needs to consider other challenges before the commercialization of such products happened. For example, translation of the knowledge to a patient population at large, ethical concerns of stem cell research, treatment cost, and regulatory aspects of tissue-engineered products (Tahriri et al. 2020). Extensive research and multidisciplinary collaboration may drive clinical trial and resolves ethical issues. The involvement of multidisciplinary fields (health sciences, engineering, and basic sciences) may support the output on dental tissue engineering products. Recently 3D printing-based technologies support to perform surgical reconstruction of maxillomandibular defects by three-dimensional virtual techniques.

10.9

Current Clinical Trials in Dentistry

Currently, there are only a few registries for clinical trials on tissue-engineered dental products and in most cases the results are still unavailable (Yamada et al. 2020). For example, the University of California, Los Angeles sponsored a clinical study in 2019 that aims to develop improved therapy for teeth that require root canals due to tooth infection or tooth inflammation. Participants were divided into two groups. First group received traditional root canal therapy. The second group received the test treatment, which involves harvesting of pulp tissues from the same tooth or other teeth. The results not yet published (https://clinicaltrials.gov/). Another clinical trial on dental stem cells and bone tissue engineering product (CELSORDINO) was sponsored by Central Hospital, Nancy, France in 2017. The study aims to develop a new pre-vascularized tissue-engineered bone construct, using human cells of a simple and non-invasive tissue source, dental pulp. In addition, the trial focus towards the boosting effect of the conditioned medium on cell differentiation and production of a vascularized bone construct. Another clinical trial aims to study the effect of encapsulated mesenchymal stem cells (RanoKure) for dental pulp regeneration sponsored by Universidad de los Andes, Chile in 2017. The results showed the efficacy of the treatment 1 year after the intervention. The effect of different surgical techniques for socket preservation procedure studying soft and hard tissue outcomes was done in 2018. The study evaluated membranes, crosslinked membrane in secondary intention healing, and a non-cross-linked membrane in primary intentional healing. In the study protocol, 30 subjects that require tooth extraction were randomly allocated to either control group (allograft covered with a non-cross-linked collagen membrane with primary closure) or experimental group

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(allograft covered with cross-linked collagen membrane left exposed) (https:// clinicaltrials.gov/).

10.10 Concluding Remarks and Outlook Tissue engineering offers as a new era for both bone regeneration and dental tissue regeneration. This article reviewed recent developments in hard tissue engineering that cover both bone regeneration and dental tissue regeneration. The use of individual biomaterials in hard tissue engineering such as polymers, bioceramics, and hydrogels is not a smart option as it needs optimal mechanical properties and porosity. Hybrid biomaterials (composites or smart materials) based scaffolds play a major role in recent developments in hard tissue engineering. Mostly, smart approaches developed in bone tissue engineering apply to dental tissue regeneration and vice versa. Acknowledgement The authors gratefully acknowledge the Director and Head BMT Wing, Sree Chitra Tirunal Institute for Medical Sciences & Technology, Trivandrum, Kerala, India for providing the facilities and for the kind permission to publish the chapter.

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Biomaterials for Soft Tissue Engineering: Concepts, Methods, and Applications

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Chelladurai Karthikeyan Balavigneswaran and Vignesh Muthuvijayan

Abstract

Soft tissues connect, support, or surround other structures and organs of the body, including skeletal muscles, tendon vessels, and nerves supplying these components. Also, organs such as the heart, brain, liver, and kidney are considered as soft tissues. Acute and chronic injury may cause transient or permanent damage to organs and soft tissues. If the damage is severe, the natural physiological repair and restoration mechanisms are not possible. The repair or regeneration using tissue engineered (TE) scaffolds has been considered as a clinical solution. TE approach involves the replacement of damaged parts using grafts made from natural or synthetic or composite polymers. Choosing the polymer with appropriate biological, physicochemical, and mechanical properties is the key to make a successful TE scaffold, and it is still a challenging task. Moreover, the fabrication technique and choice of cells or growth factors for encapsulation to develop the graft also play a crucial role. Therefore, in this chapter, we have highlighted the grafts developed for engineering soft tissues such as blood vessels, skin, cartilage, intervertebral disc, tendon, and skeletal muscle. We have restricted our focus on electrospun scaffolds, and injectable hydrogels prepared using polymers include collagen (Col), chitosan (CS), hyaluronic acid (HA) alginate (Alg), poly (caprolactone) (PCL), poly(lactic acid) (PLA), poly(glycolic-lactic acid) (PLGA), and their composites. This chapter will help the readers to understand the choice of materials and fabrication techniques for developing successful TE scaffolds for soft tissue engineering applications.

C. K. Balavigneswaran · V. Muthuvijayan (*) Tissue Engineering and Biomaterials Laboratory, Department of Biotechnology, Bhupat and Jyoti Mehta School of Biosciences, Indian Institute of Technology Madras, Chennai, Tamil Nadu, India e-mail: [email protected] # The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 B. Bhaskar et al. (eds.), Biomaterials in Tissue Engineering and Regenerative Medicine, https://doi.org/10.1007/978-981-16-0002-9_11

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Keywords

Mechanical properties · Soft tissue engineering · Electrospinning · Injectable hydrogels

11.1

Introduction

Soft tissues connect, support, or surround other structures and organs of the body, which includes skeletal muscles, skin, ligament, vessels, and nerves. In addition, organs such as the heart, liver, and kidney are considered as soft tissues. Acute and chronic injury may cause transient or permanent damage to organs and soft tissues. If the damage is severe, the natural physiological repair and restoration mechanisms are not possible. The best way of treating the damage is to use auto-, allo-, or xenogeneic grafts. The availability of donors and immunocompatibility limits the usage of allo- or xenografts. Therefore, alternative approaches for restoration and repair are required. With the advances in medical and natural sciences, Langer and Vacanti first defined the term “tissue engineering” (TE) in the 1990s, which focused on exploiting body’s own ability to regenerate, with the use of materials (Cima et al. 1991; Vacanti and Langer 1999; Shinoka et al. 1995). The three-dimensional (3D) scaffolds mimic and provide the support as an artificial extracellular matrix (ECM). A native ECM is a 3D network that interacts with the cells, provides structural support, transfers mechanical forces, and allows the transport of chemical signals. The global market for tissue engineering is expected to reach US$94.2 billion by 2022 (Ye et al. 2018). The TE scaffolds can be fabricated from synthetic, natural, or composite macromolecules using techniques such as solvent casting (Liao et al. 2002; Janik and Marzec 2015), particulate leaching (Liao et al. 2002; Janik and Marzec 2015), gas foaming (Janik and Marzec 2015), electrospinning (Khorshidi et al. 2016), injectable hydrogels (Hunt et al. 2014), and 3D printing (Lee and Yeong 2016). Controlling the porosity and surface properties of engineered scaffolds is essential for the specific biological function. Achieving the required properties of TE scaffolds for soft tissues is challenging as the physicochemical, mechanical, and biological properties significantly vary between different types of tissue such as blood vessels, tendon, skin, cartilage, and individual organs. The natural and synthetic polymers or a combination of hybrid polymers also exhibit different properties. Thus, choice of polymer and its properties to match for a specific application is a key consideration for soft tissue regeneration. Therefore, in this chapter, we have provided an overview of general physicochemical properties required for the 3D scaffolds to be used for soft tissue regeneration and different polymers used for blood vessels, skin, cartilage, IVD, tendon, and skeletal muscle tissue engineering.

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11.2

383

The Properties of Scaffolds for Soft Tissue Engineering

TE scaffolds provide the platform to guide the formation of new tissue. To mimic the properties of natural soft tissues using TE scaffolds, understanding the following properties of polymer is mandatory. (1) The polymer should be biocompatible and biodegradable. (2) The functionality, interconnectivity, porosity, pore size, crosslinking density, and degradation rate are the critical properties that determine the appropriateness of the polymer for soft tissue engineering. (3) The mechanical properties, such as tensile strength and strain, flex strength and strain, elasticity, and compression strength of the 3D scaffold, should mimic the targeted soft tissue to attain the required biological properties (Mondschein et al. 2017). Hence, balancing the physicochemical, biological, and mechanical properties, as shown in Fig. 11.1, is essential to replicate the targeted soft tissue.

11.2.1 Biological Properties The biocompatibility of a TE scaffold is crucial as all the cellular functions are dependent on the biomaterial. The factors affecting the biocompatibility of scaffolds include choice of polymers, structure, and chemistry used for functionalization (Asghari et al. 2017). Synthetic polymers offer excellent physicochemical properties to fabricate TE scaffolds. Especially, polyethers and polyesters are the most commonly used synthetic polymers. On the other hand, natural polymers mimic the host ECM. Polysaccharides and polypeptides are the widely used natural polymers. We are limiting the scope of our chapter to natural polymers such as collagen (Col), gelatin (Gel), alginate (ALG), and hyaluronic acid (HA); and synthetic polymers Fig. 11.1 Properties required for biodegradable TE scaffolds

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such as poly(ɛ-caprolactone) (PCL), poly(lactic acid) (PLA) and poly(lactide-coglycolic acid) (PLGA); and their composites due to the advantages they possess.

11.2.2 Physicochemical Properties As physicochemical properties influence cytotoxicity, surface properties, and degradation of polymeric scaffolds, a thorough characterization after fabrication is necessary to understand their potential use in soft tissue engineering.

11.2.2.1 Cytotoxicity The most common causes for the cytotoxicity that lead to cell necrosis or apoptosis include the initiator used, and organic solvent residue from polymer synthesis; crosslinker used for fabricating scaffolds, and the products after polymer degradation (Oryan et al. 2018). 11.2.2.2 Fabrication Techniques There are many techniques available to fabricate 3D scaffolds to be used for soft tissue engineering such as solvent casting (Liao et al. 2002; Janik and Marzec 2015), particulate leaching (Liao et al. 2002; Janik and Marzec 2015), gas foaming (Janik and Marzec 2015), electrospinning (Khorshidi et al. 2016), and injectable hydrogels (Hunt et al. 2014). We have limited our chapter with electrospinning and injectable hydrogels as much studies have been done using these techniques. 11.2.2.3 Surface Properties of TE Scaffolds The surface of the TE scaffold is the primary site for the cells to interact. Therefore, the surface area of the TE scaffold should be large enough for efficient cell interaction, and it should be studied seriously to develop an efficient scaffold system. The surface includes both chemical and topographical characteristics (Mondschein et al. 2017). In addition to the surface properties, the degradation mechanism and degradation rate of the TE scaffolds are very important factors while designing a new TE scaffold (von Burkersroda et al. 2002).

11.2.3 Mechanical Properties The stability of TE scaffolds in the physiological environment is dependent on water absorption at the polymeric material interface, strength, elasticity, and degradation of the polymer. The stability of scaffolds is a crucial factor as the TE scaffolds should temporarily resist the loads and promote the regeneration of tissue. Therefore, in the view of mimicking in vivo tissue properties to establish the appropriate biomechanical functions, TE scaffolds ideally should match the native tissue’s mechanical properties (Wegst and Ashby 2004). The tensile, compression, and flexural properties of natural tissue vary upon age, gender, and ethnic group all for similar tissue. The biomechanical properties of soft tissues such as cardiac tissue including

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Table 11.1 Mechanical strength of various soft tissues Property Tensile Human tissue type (s) Valves and vasculature Skin Cartilage

Strength (MPa) 0.2–2.5

1–20 2.9–40

Intervertebral disc Annulus 80–20 fibrosus Nucleus 3–10 fibrosus Tendon 40–100

Strain at break (%) 15–150

Modulus (MPa) 0.1–1.3

30–70

0.3

15–120

4.5–24

20–30

18–45

Nerurkar et al. (2010)

20–30

0.138

Nerurkar et al. (2010)

50–345

Herman (2007), Mondschein et al. (2017), Butler et al. (1986), Blanton and Biggs (1970)

9–15

Reference(s) Albanna et al. (2012), Baudis et al. (2009), Mondschein et al. (2017), Mondschein et al. (2017), Herman (2007) Herman (2007), Mondschein et al. (2017), Kim et al. (2019), Schneider (1983) Herman (2007), Mondschein et al. (2017), Kerin et al. (1998), Kempson (1982), Yang et al. (2017)

blood valves and heart muscle (Tiburcy et al. 2014); connective tissues include ligaments (Quapp and Weiss 1998), tendons (Blanton and Biggs 1970), and cartilage (Yang et al. 2017); and epithelial tissues include intestines and skin (Schneider 1983) have been well studied (Herman 2007; Wegst and Ashby 2004). Table 11.1 summarizes the mechanical strength of the various soft tissues reported in the literature. Ideally, the mechanical strength of cartilage will be highly more than the neurons. The pore volumes/size/shapes, orientations, and interconnectivity of the TE scaffolds are found to be influencing the mechanical properties and structural integrity of the TE scaffolds. It has been reported that the cell morphology, molecular expression of proteins, and the cellular functions are significantly affected by the elasticity of the TE scaffolds (Engler et al. 2006; Buxboim et al. 2010). The choice of polymer is important to replicate the bulk mechanical properties of the target tissue. The tuning of scaffold stiffness will be easier with synthetic polymers due to the chemical functional groups. The possible way to control the mechanical strength of TE scaffolds could be by adding nanofillers to tailor its mechanical properties (Cheung et al. 2007) and increasing crosslink density (Peterson et al. 2016).

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Application of TE Scaffolds in Soft Tissue Engineering

11.3.1 Vascular Tissue Engineering 11.3.1.1 Structure of Blood Vessels Blood vessels are conduits, which deliver oxygen, and nutrients to various parts of organs and removing the waste products from tissues. It is responsible for withstanding in a wide range of pressures and shear stresses, regulating blood flow and permeability, and resisting thrombosis under basal conditions. They also play a vital role in immunological responses. Blood vessels can be categorized into three types based on the size, such as (1) microvessels (6 mm in diameter). Blood vessels are of a concentric layered structure. Each layer of the blood vessel is made of specific cells and biomolecules such as collagen and proteoglycans. Figure 11.2 shows the structure of a blood vessel. The innermost layer is called intima, and it has a monolayer of specialized endothelial cells (ECs) that form a tight non-thrombogenic barrier between the lumen of the vessel and the rest of the wall. This layer prevents unwanted clot formation, protects infection, and inflammation from the underlying tissue. This also acts as a medium for sending the signals to the muscular component of the vessel wall. The next layer is the basement membrane, which consists of collagen IV and laminin. The muscular layer of the artery is the next. It is also called as media, and is embedded with smooth muscle cells (SMCs). It is majorly constituted by collagens Type I and III, and lesser amounts of other proteins and proteoglycans are also present. The collagen matrix and the SMCs are generally aligned circumferentially, or in a spiral pattern along the axis of the vessel. The contraction or dilation of the vessel is coordinated by media, which is stimulated by the signals from the ECs of the lumen. The final layer is adventitia, which composed of a loose collagen matrix embedded with fibroblasts. This layer serves

Fig. 11.2 The transverse (a) and longitudinal (b), cross-sectional (c) views of the structure of the blood vessel (Wang et al. 2019)

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as a substrate for a vascular supply to the artery wall. It anchors the blood vessel to the surrounding tissue, and provides additional structural support (Wang et al. 2019).

11.3.1.2 Need for Vascular Tissue Engineering The damage of blood vessels may occur in many ways, including injury, disease, or inflammation, which causes aneurysms and dissections. However, the most common reason for vessel failure is occlusion due to atherosclerosis. This inflammatory disease causes a build-up of plaque beneath the intimal layer of the vessel wall. The plaque is formed due to the infiltration of monocytes into the intima, which increases the proliferation, migration, and secretory activity of the SMCs. Further, it narrows down the blood vessel lumen, and decreases the blood flow to the tissues, when the plaque grows and calcifies. The plaque rupture and subsequent clot formation cause infarction of the downstream tissue. The increased incidence of atherosclerotic disease demands a strong need to treat blocked vessels, especially for blood vessel replacement (Wang et al. 2019). The first vein allograft interposition to treat aneurismal arterial segment was reported in the early 1900s, but host immune rejection limited its successful application. In the late 1940s, autografts of saphenous vein were successfully used for femoropopliteal bypass surgery to treat diseased small-diameter vessels (Menzoian et al. 2011). The saphenous vein is the most commonly used autograft, but internal mammary arteries and radial arteries can also be used. Although autograft veins and arteries are considered to be a opt vascular substitute, it is not suitable for approximately one-third of patients, and additional surgical revascularization is required within 10 years of the initial surgery. Therefore, alternative grafts are needed specifically for cardiovascular diseases (Seifu et al. 2013; Chang and Niklason 2017). Vascular tissue engineering aims to establish live conduits, matching the properties of the native blood vessels. The engineered vascular grafts should mimic the properties, such as burst pressure  1700 mmHg (equal to that of the saphenous vein); ability to withstand against cyclic physiological loading in vitro without any dilatation for 30 days; the ability of the autologous EC to form a monolayer (equivalent to the intimal layer of native vessels); and consistency during the fabrication of concentric structure with 6 mm in case of small vessels. In addition, the grafts should have the ability to support the proliferation of endothelial cells, fibroblasts, and vascular SMCs. In recent years, vascular surgery has emerged from “replacement” to “regeneration” of the vascular tissue. Hence, cells can be used as one of the sources for regeneration, which can be extracted from either the patient or the donor, but the use of autologous cells is preferable to control the risk of infection and immune rejection. In addition, stem cells such as MSCs and hematopoietic stem cells have also been considered for the treatment (Seifu et al. 2013; Chang and Niklason 2017; Stegemann et al. 2007). 11.3.1.3 Tissue Engineered Vascular Graft The vascular TE strategies can be broadly divided into two groups, regardless of the source of the cells being used: self-assembled cell-sheet technology and 3D scaffold-

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based approach (decellularized matrix or natural or synthetic or hybrid polymer) (Seifu et al. 2013). Since the focus of our chapter is the biomaterials-based approach, we will be discussing only different scaffolds used and injectable hydrogels cannot be used for vascular tissue engineering. Therefore, this section will concentrate on nanofibers-based scaffolds as vascular grafts. 11.3.1.3.1 Electrospun Scaffold-Guided Vascular Grafts Electrospinning is the more suitable technique for making vascular grafts as tailoring the structure, and tunability of the mechanical properties can be easy during the fabrication process. Functionalizing the synthetic polymer matrix with natural polymers having inherent cell-binding sites can support to form a monolayer of ECs in the lumen, and proliferation of other cell types in the matrix of the graft’s wall. Further, aligned nanofibers will help polarizing cells in a certain direction to provide the anisotropy encountered in certain organs, including blood vessels (Hasan et al. 2014). Many researchers have fabricated acellular electrospun scaffolds with poly-Llactic acid-co-poly-e-caprolactone (PLLCL) for their use in vascular tissue engineering (Wise et al. 2011; He et al. 2009). He et al. fabricated tubular scaffold from PLLCL, and studied their effect in rabbits for 7 weeks. Their results showed that the scaffolds maintained its structural integrity, and there was no immune rejection. Further, ECs cultured collagen-coated PLLCL scaffolds showed that the ECs proliferated uniformly throughout the scaffolds for the period they studied (1–10 days) (He et al. 2009). De Valence et al. evaluated the grafts fabricated using PCL-based nanofiber and microfiber in a rat model for abdominal aortic replacement. Their results revealed good mechanical properties, rapid endothelialization, and patency after 18 months of their study. However, the regeneration of the vessel wall was poor (de Valence et al. 2012). The nanofibers fabricated using PCL/tropoelastin achieved compliance with the elastic modulus, permeability, and burst pressure of the human internal mammary artery. The implantation of acellular graft in a rabbit model showed enhanced endothelialization with reduced platelet attachment in PCL/elastin conduits compared to PCL graft (Wise et al. 2011). Further, bilayer electrospun scaffolds of PCL-collagen blend demonstrated to produce a strong and multilayer structure equivalent to that of native blood vessels (Ju et al. 2010). Also, PCL/collagen scaffolds were also fabricated to support cell growth, and maintain patency in a rabbit aortoiliac bypass model. In vivo studies showed that these scaffolds were able to retain their structural integrity over 1 month, and biomechanical strength of blood vessels after a month was comparable to the native artery (Tillman et al. 2009). Matsumara et al. for the first time, reported the use of bone marrow cells (BMCs) seeded onto PLCL/PLLA scaffold for vascular tissue engineering. They fabricated the scaffolds using electrospinning followed by freeze-drying technique. They implanted the scaffolds into the inferior vena cava of dogs, and investigated for 2 years. Their results showed that the cells proliferated, differentiated into SMCs, and partly contributed to the endothelialization of the tissue engineered graft. Also, no stenosis was observed (Matsumura et al. 2003). The small-diameter blood vessel

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fabricated using PGA (polyglycolic acid) scaffolds was cultured with rat cardiac fibroblast cells and evaluated for their biocompatibility in rats by intramuscular implantation. The results showed that PGA was biocompatible (Boland et al. 2004). Zhu et al. showed that aligned PCL fibers coated with fibrin induced the alignment of hSMCs along the direction of PCL/fibrin fibers. In addition, the aligned fibers promoted SMC modulation into a contractile phenotype and biological function (Zhu et al. 2010). Recently, Ardila et al. reported TGFβ2 loaded PCL/gelatin nanofibrous scaffolds modulating the SMCs. The fabricated tubular scaffolds supported the infiltration and growth of vascular SMCs (Ardila et al. 2019).

11.3.2 Skin Regeneration Skin is the largest organ of the human body, accounting for about 15% of the total body weight in adults. The integumentary system comprises the skin, and its appendages primarily serve as a protective layer against the external environment, such as physical, chemical, mechanical disturbances, and pathogenic microorganisms. It helps in thermoregulation and hydration retention that prevents substantial body fluid loss. It acts as an active site of immune defense and vitamin D synthesis. It also performs other crucial functions, such as self-healing, and sensory detection (Bhardwaj et al. 2018; Vijayavenkataraman et al. 2016).

11.3.2.1 Structure of Skin The skin has a very complex multi-layered structure that is mainly composed of three layers: epidermis, dermis, and subcutaneous hypodermis layer. The outer epidermis serves as a barrier against the external environment, and helps for retention of hydration. It is composed of keratinocytes and melanocytes. The dermis layer is a highly vascular thick connective tissue that constitutes the bulk portion of the skin. The mechanical strength of the skin is influenced by the constituents of the dermis layer, viz., the fibroblasts and extracellular matrix (ECM) components such as glycosaminoglycans (GAGs), elastin, and collagen. In addition, it also contains various appendages such as sweat glands, sebaceous glands, and hair follicles, along with blood vessels (Vijayavenkataraman et al. 2016). The well-vascularized hypodermis or subcutaneous layer is the bottom layer, which is made of adipose tissue, contributes to the thermoregulatory and mechanical properties of the skin. It acts as a cushion between the skin and other skeletal structures (Bhardwaj et al. 2017). The layer between epidermis and dermis is called epidermal junction. Various molecular or cellular level signaling processes occur via communication between epidermal and dermal layers (Fig. 11.3). 11.3.2.2 Need for Skin Tissue Engineering Cutaneous wounds can generate a continuum of injuries, which increases the wound depth in the skin. The increase in the number of trauma and pathophysiological conditions represent a major healthcare problem globally. The normal wound healing process is a very well-organized process that consists of series of events

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Fig. 11.3 Structure of skin (MacNeil 2008)

such as hemostasis, inflammation, proliferation, and ECM remodeling (Metcalfe and Ferguson 2007). The four healing phases involve interactions between various types of cells, bioactive factors, and a supporting platform. The recruitment of blood platelets and immune cells will happen in hemostasis, and inflammation to control the blood loss, and clearing of pathogens, respectively. The recruited immune cells secrete chemokines, and growth factors, which is useful during the proliferation phase. Wound contraction happens in the proliferation phase, which involves events such as granulation development, angiogenesis, and re-epithelialization. This phase is regulated by cross-talk between macrophages, and other cells embedded in the skin such as fibroblasts, ECs, and keratinocytes. The final phase is remodeling, wherein the contracted wound forms either a functional skin or a non-functional scar tissue. This normal healing process is dysregulated in pathophysiological conditions. The majority of non-healing wounds are associated with trauma or disease conditions like diabetes. The large trauma wounds due to burns and accidents result in loss of significant amount of skin, where self-healing of the skin fails. Depending on the type of chronic wound (pressure sores, ulcers, or diabetic wounds), the healing process could be hindered at any of the four phases (Goodarzi et al. 2018). Skin injury is one of the major health concerns due to the complications from accidents or co-morbidities such as diabetes or infection. According to WHO statistics, 58 million people are affected by fatal injuries, among which 5 million people die each year, and several millions of people require proper treatment and care (Organization 2010). Partial-thickness wounds due to big accidents generally do not require grafting. They can be treated using antimicrobial dressings, which will help the regrowth of epithelial appendages to cover the wounds. However burns, which do not heal within ~3 weeks, and full-thickness burns require autograft or allograft for the treatment, wherein a replacement of the epidermal barrier occurs using autologous keratinocytes. Transplantation of keratinocytes can be achieved by conventional split-thickness skin grafts (STSG). Autologous STSG may provide permanent wound closure in a short-time compared to allogeneic STSG. The

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delivery of growth factors and extracellular matrix components along with STSG results in rapid wound closure. The STSG has been shown to reduce granulation tissue, minimize the scar, and produce functional skin. The meshed or unmeshed SPSG can be fabricated based on application (Goodarzi et al. 2018). Although auto and allografts show noticeable effects in would closure after the trauma injury, the lack of donors and additional pain caused in the patient in case of autograft limit their successful application. This shows the huge demand for artificial skin grafts.

11.3.2.3 Tissue Engineered Skin Grafts The majorly studied techniques for fabricating TE scaffolds for skin tissue engineering are injectable hydrogels and electrospinning. 11.3.2.3.1 Injectable Hydrogels for Skin Tissue Engineering As the collagen is the major constituent of ECM of skin, collagen has been extensively used to produce injectable hydrogels responsive to temperature. Many studies have been reported with collagen as a matrix for skin tissue engineering. In a study, Helary et al. used collagen as a dermal substitute (Helary et al. 2012). The type of collagen did not influence on proliferation of fibroblasts and vascularization in the view of replacing the damaged skin. Further, epithelization and vascularization in the membrane group (Col-I/Col-III) were of the same compared to the group treated with an STSG (Wehrhan et al. 2010). Moreover, combining collagen with gelatin improved cell viability, and accelerated the wound healing. This has also been found to promote the neovascularization of burn wounds (Chan et al. 2015). It has been reported that the local pH of chronic wound in the range between 5.4 and 8.9 depending on the location of the wound, degree of necrosis, and local oxygen availability. This deviation enabled pH inducible gelation to be utilized as a potential treatment modality for skin associated malignancies (Dimatteo et al. 2018). Qu et al. reported to use quaternized chitosan (QCS), and benzaldehydeterminated Pluronic®F127 (PF127-CHO), which formed the pH-dependent injectable hydrogel via dynamic Schiff-base chemistry. They also found that the developed hydrogel dressings were stretchable, and the strength was equivalent to the modulus of human skin. Further, they showed good adhesiveness and immediate self-healing to bear the deformation. In addition, the loaded curcumin significantly accelerated wound healing in vivo by producing higher thickness of granulated tissue, collagen disposition, and upregulated vascular endothelial growth factor (VEGF) in a full-thickness skin defect model (Qu et al. 2018). In a similar study, the hydrogels were formed by Schiff-base chemistry using oxidized alginate (OAlg) and carboxymethyl chitosan (CMCS) polymers. To achieve the hydrogels to be antibacterial, Chen et al. incorporated gelatin loaded microspheres into the hydrogel matrix. The composite injectable hydrogel matrix inhibited the growth of bacteria, such as E.coli and S.aureus (Chen et al. 2017). Ionic cross-links are another kind of method to form hydrogels using the interactions between localized charges on the polymer backbones, and counter ions present in solution to build salt bridges. This produces mechanically stable gels. Pereira et al. reported the Alginate/Aloe vera thin-film hydrogels prepared using

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calcium chloride crosslinking for skin tissue engineering in the view of aloe vera to support cell proliferation (Pereira et al. 2013). In another study, a novel bilayer hydrocolloid film based on ALG was reported as wound dressing material. The ibuprofen loaded bilayers healed the wound efficiently in vivo (Thu et al. 2012). The hydrogels responsive to photo light have also been studied for skin tissue engineering. In a study, Zhao et al. demonstrated the efficacy of photosensitive methacrylated gelatin (GelMA) hydrogels for epidermal reconstruction. These GelMA hydrogels vary in mechanical properties based on gelatin concentration in the pre-polymer solution. The mechanical strength of the developed GelMA hydrogels were equivalent to the natural epidermis, which made them as ideal candidates for treating chronic wounds (Zhao et al. 2016, 2017). Feng et al. reported the formation of hydrogels using the combination Konjac glucomannan/heparin for the localized activation of macrophages, and sequestration of pro-angiogenic signals. The heparin in the gel allowed to establish an efficient binding with pro-angiogenic growth factors VEGF, platelet-derived growth factorBB (PDGF-BB), and basic fibroblast growth factor (bFGF). The human monocytes and murine macrophages adhered hydrogels exhibited the upregulated production of pro-angiogenic signals in vitro. In vivo results also demonstrated vascularized skin with the higher expression levels of VEGF and PDGF sequestration (Feng et al. 2017). In another study, the hydrogel was prepared using alginate, chitosan with Alpha-tocopherol (vitamin E) (Chit/Alg/400 IU Vit E). Vitamin E was used for better wound healing performance. The authors found that the hydrogel-based dressings had better wound closure compared to the gauze-treated wound (the control) (Ehterami et al. 2019). The stem cells or autologous cells-loaded hydrogels have also been found to accelerate wound healing. Eke et al. developed a dermal substitute using GelMA/ HA-MA containing adipose-derived stem cells (ADSCs) to improve the regeneration of skin by stimulating vascularization. The in vivo results revealed ADSCs encapsulated hydrogels increased vascularization up to three-fold compared to their acellular counterparts (Eke et al. 2017). Similarly, Schmitt et al. loaded MSCs along with VEGF and bFGF into alginate hydrogels gels for better wound healing. The authors found both VEGF and bFGF were released in a sustained manner for a six-week duration. This enhanced the proliferation of MSCs, which is useful for skin regeneration (Schmitt et al. 2015). 11.3.2.3.2 Nanofibrous Scaffolds for Skin Tissue Engineering Asymmetric membranes have been fabricated to reproduce skin anatomy for the enhanced healing process. This asymmetric membrane exhibits general properties such as dense network, hydrophobic microporous top layer, which prevents bacteria penetration as well as interconnected pores that allow the exudate absorption, gaseous exchange, and cell migration. The detailed review of electrospun nanofibers as skin substitutes is reviewed (Sundaramurthi et al. 2014; Miguel et al. 2018). Many studies have been reported to use the fibrillary structure of collagen Type I (Col-I) for fabricating the nanofibers (Law et al. 2017). Further, it has also been found that electrospun collagen favors cell attachment and proliferation, decreases wound

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contraction, and improves early-stage wound healing when compared to freeze-dried collagen scaffolds (Powell et al. 2008; Rho et al. 2006). The gelatin (Gel) was able to produce fibers with the size ranging from 5–10 μm, which favored the efficient growth of fibroblasts and keratinocytes. This revealed Gel was an efficient dermalepidermal composite skin substitutes in vitro. Moreover, glutaraldehyde improved the structural stability of nanofibrous scaffolds (Zhang et al. 2006). It has been observed that HA has a positive influence on all the phases of wound healing. In this view, HA nanofibers have also been used as a superior wound healing matrix compared to other wound dressing materials (Uppal et al. 2011). Synthetic polymer PCL was also found to support fibroblast cells, which mimicked a complex organization of Col-I. It has also been found that the high-density fibers affected cell migration due to the fishnet effect (Xie et al. 2010). Venugopal et al. reported a nanofibrous membrane of PCL, and collagen was fabricated, wherein PCL provided the mechanical strength, and Col provided the support for cellular function. Thus the authors suggested PCL/Col nanofiber for the treatment of skin defects and burn wounds (Venugopal et al. 2006). Nanofibers from blending mixtures of poly(L-lactic acid)–co–poly(ε-caprolactone) (PLPCL) and Gel have been shown to provide an efficient matrix for skin tissue engineering applications (Jin et al. 2012). The onsite delivery of stem cells/xenogenic cells and bioactive factors like genes, drugs, proteins/peptides using the PLPCL/Gel matrix showed accelerated wound healing in animal models (Jin et al. 2012; Bishi et al. 2013). Chen et al. fabricated nanofibrous scaffolds from chitosan-graft-poly (ɛ-caprolactone) (CS-g-PCL). The enhanced cell proliferation of L929 on CS-g-PCL suggested its use for skin tissue engineering (Chen et al. 2011a). CS/PVA nanofibers have been reported to have good cytocompatibility, and also enhanced wound healing in vivo (Zhou et al. 2008). The naturally inherent properties of Gel and CS were used to fabricate Gel/CS electrospun scaffolds. These were reported to support fibroblast proliferation efficiently (Dhandayuthapani et al. 2010). Antimicrobial activity can also be achieved by incorporating antibiotics within nanofibers. Liao et al. studied the antimicrobial activity of PCL/cellulose/dextran electrospun nanofibrous mats loaded with tetracycline hydrochloride. Their results revealed that nanofibers loaded with tetracycline alone inhibited the growth of bacteria (Liao et al. 2015). Heyu Li et al. loaded ciprofloxacin (commonly used to treat skin infections) into PLACL thermoresponsive electrospun mats showed to inhibit the growth of E.coli and S.aureus. Further, this also promoted better wound healing compared to commercial gauzes (Li et al. 2017). The inherent antimicrobial activity of CS was also used in various studies for skin tissue engineering applications (Ding et al. 2014). The antibacterial activity CS is achieved due to the interactions between amino groups of CS and the electronegative functional group present on the bacteria’s cell wall. This interaction inhibits the membrane permeability, cell leakage, and ultimately to the death of the cell (Miguel et al. 2018). In a study, CS/PLA electrospun membranes reported inhibiting the growth of S.aureus and E.coli (Ignatova et al. 2009). Antunes et al. reported the use of deacetylated/ arginine-modified CS to fabricate full-thickness skin. They were also found to inhibit the growth of E. coli and S. aureus due to the surface charge presented by

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arginine-modified CS (Antunes et al. 2015). A nanofiber from a composite of collagen/chito-oligosaccharides was demonstrated as a dermal substitute because of the enhanced proliferation of human skin fibroblast and antibacterial activity (Wang et al. 2011). Electrospun gelatin/silver nanoparticles were used as a wound dressing material with efficient proliferation along with antibacterial activity in vitro (Rujitanaroj et al. 2008). Miguel et al. produced an asymmetric nanofibrous matrix to use as a full-thickness skin substitute. They fabricated the matrix from blending a mixture of PCL, CS, and AV. The top and bottom layer of the scaffold were fabricated using PCL (epidermis), and CS with Aloe Vera (AV) (dermis), respectively. The top layer provided good mechanical strength with low porosity, whereas the bottom layer allowed the growth of fibroblasts to develop the ECM and helped in the healing process. It also showed antimicrobial property against S.aureus and E. coli at the wound site because of the CS presence (Miguel et al. 2017). The strategy of using plant extracts for treatment of burns and wounds was also studied. Jin et al. reported the nanofibers fabricated from plant extracts of Indigofera aspalathoides, Azadirachta indica, Memecylon edule (ME) and Myristica andamanica along PCL. PCL/ME nanofibers were found to support higher proliferation compared to other matrix. Further, they observed epidermal differentiation of ASCs on PCL/ME scaffolds which suggested their potential as substrates for skin tissue engineering (Jin et al. 2013).

11.3.3 Cartilage Tissue Engineering 11.3.3.1 Structure of Cartilage Cartilage is present between the bone surfaces, which is a highly specialized connective tissue characterized by its unique mechanical properties that provide wear resistance under high loading. Its frictionless surface facilitates smooth sliding movements of the bone. The cartilage is avascular, aneural, and lymphatic. It consists of chondrocytes embedded in ECM, which is present in the trachea, bronchi, nose, ear, larynx, knee joints, and intervertebral discs. The solid ECM matrix of this connective tissue is primarily composed of an integrated network of Col-II and proteoglycans (PGs), which are responsible for its mechanical properties (GhasemiMobarakeh et al. 2013). The layered organization of cartilage can be classified into four zones based on the function and structure: the superficial zone, the middle or transitional zone, the deep zone, and the zone of calcified cartilage. The top 10–20% of the thickness of the cartilage surface is a superficial zone or tangential zone, which helps for gliding movements and resists shear. The highest Col content is observed in this region, wherein Col is densely packed and arranged parallel to the articular surface. The chondrocytes in this region are flattened, elongated, and are also parallel to the surface. The next layer is the middle zone, which constitutes up to 40–60% of the articular cartilage volume. Here, Col fibrils are loosely packed and aligned obliquely to the surface; chondrocytes appeared to be more rounded with lower cell density. The next layer to the middle zone is a deep zone, which encompasses 30% of the

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cartilage. This zone has the highest concentration of PG; Col fibrils are found to be perpendicularly arranged, and the lowest water content, wherein chondrocytes are observed as large in size and oriented perpendicular to the articular surface. The mechanical strength of the first three zones of the articular cartilage is found to be in an increasing order ranging from the superficial zone to deep zone. The final layer is a calcified cartilage zone, which is found adjacent to the underlying cortex of the bone. The mineralization of calcium phosphate and endochondral ossification occurs in this zone. After complete mineralization, chondrocytes undergo apoptosis; osteoclasts invade the calcified cartilage, and the bone is laid down by the osteoblasts (Ghasemi-Mobarakeh et al. 2013). Col-II is the major constituent (approximately 90%) of articular cartilage, and small amounts of other types of collagen-like types I, III, V, VI, IX, and XI are also found. The two major classes of PGs, viz., large aggregating molecules, and smaller proteoglycans account for 4–7% of the cartilage weight. These are homogeneously distributed in the cartilage. The concentration of PGs is lowest in the superficial zone and increases towards the cartilage–subchondral bone interface. The synovial fluid contributes 60–80% of total cartilage weight, and the water content is high in the superficial zone, and decreases towards the deep zone. Cartilage is classified into three types—fibrocartilage, elastic cartilage, and hyaline cartilage. The fibrocartilage is a dynamic tissue that is located within fascicles or in endo- or epitenon. They are fibrocartilaginous, where the tendons and ligaments are subject to compression. The cells in the fibrocartilage are packed with intermediate filaments, which involve in transducing mechanical load. Fibrocartilage is rich in Col-I and present in tissues that undergo tensile forces such as the intervertebral disk. Elastic cartilage consists of elastic fibers of Col, which is found in the external ear, epiglottis, and the upper portion of larynx. The elastic nature allows us to return back to its original shape when there is deformation. The articular surface of joints, which is composed of hyaline cartilage made of Col-II and PGs, functions as a shock absorber. The mechanical strength of the hyaline cartilage is found as 2.9–4.0 MPa depending upon the permeability of interstitial fluid in the tissue, and the balance between the solid matrix and fluid phases. The remodeling of collagen fibrils during maturation forms the inhomogeneous structure of the cartilage. Further, this 3D collagen architecture is mainly influencing the dynamic and tensile strength of articular cartilage, and it prevents tissue deformation in the direction of the Col fibrils with respect to the zone. PGs provide the static stiffness of the cartilage, and they resist compression and fluid flow throughout the tissue, thereby influencing permeability. Moreover, half of the compressive stiffness of the articular cartilage is attributed to the electrostatic repulsion and osmotic swelling interactions of the aggrecans (Ghasemi-Mobarakeh et al. 2013; Radhakrishnan et al. 2017).

11.3.3.2 Need for Cartilage Regeneration The regeneration or repairability of cartilage is poor due to avascularization, sparse, and highly differentiated cell population, and slow matrix turnover. The response of cartilage to injury depends on the severity of the injury, and the effect may range

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from fibrillation to development of an osteochondral defect. Osteoarthritis (OA) is an age-related clinical syndrome associated with degeneration of articular cartilage, which affects the normal cartilage structure and function. It affects daily activities of about 23% of the aged adult population. Globally, 9.6% of men, and 18% of women over 60 years of age have been diagnosed with symptomatic OA. The disease causes joint tenderness, varying degrees of inflammation, joint pain, occasional effusion, reduces the easy mobility between joints and degeneration of the joint. Chemically, breakdown of the Col network, reduction in PG concentration, increase in water content, softening of the cartilage tissue result in OA. Biologically, fibrillation of cartilage surface occurs, which extends to the deeper zones for the fibrillated fissures growth resulting in the torn of the cartilage surface and releasing cartilage fragments into the surrounding joint space. As this progresses, the cartilage becomes thinner direct bone-to-bone contact occurs (Radhakrishnan et al. 2017; Lü et al. 2015).

11.3.3.3 Tissue Engineered Cartilage The conventional clinical treatments for articular cartilage defects include debridement, spongialization, drilling, microfracture, osteochondral, periosteal/perichondrium transplantation, and so on. Although these strategies have exhibited some efficacy in restoring the damaged articular cartilage, they have some limitations. Debridement causes cell apoptosis, and changes the mechanical loading way. The bone marrow and the bone marrow mesenchymal stem cells (BMMSCs) filled in drilling, and microfractures form fibrocartilage rather than the natural hyaline cartilage, which results in OA. Spongialization removes the subchondral bone, but exposes the cancellous bone, resulting to large invasion. Transplant rejection occurs in the case of allograft cartilage, and the long-term prognosis of periosteal and perichondrium graft is unsatisfied because the stiffness of neocartilage was lesser compared to the natural cartilage. On the other hand, transplantation of autologous cartilage is effective in repairing the partial articular defects, but large cartilage defects cannot be treated, and it brings a secondary injury (Zhang et al. 2019). Therefore, there is an urgent need to develop a promising alternative approach articular cartilage regeneration. 11.3.3.3.1 Injectable Hydrogels The main advantage of injectable hydrogels for cartilage repair is: it is minimally invasive, and encapsulating with stimulatory growth factors will stimulate the regeneration of defect cartilage. Natural polymers such as Col, Gel, CS, and their derivatives have been studied to prepare injectable hydrogels as scaffolds for cartilage tissue engineering. By tuning the components, concentration, and conducting chemical modification, and using chemical cross-linkers, these polymers would form a hydrogel in the physiological environment, especially at 37  C (Eslahi et al. 2016; Zhang et al. 2019). Col-I and Col-II are the main components in connective tissues and hyaline cartilage. Col-II plays a significant role in maintaining the proliferation, function of chondrocytes, and inducing chondrogenesis of MSCs (Yuan et al. 2016; Kim et al. 2015; Lee et al. 2003). On the other hand, there are reports wherein chondrocytes

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cultured in Col-I gel showed a higher rate of cell proliferation and gene expression (Nöth et al. 2007). Moreover, it was found that biomolecules such as bone morphogenetic protein-2 (BMP-2), TGF-β, dexamethasone, insulin, and cytokines stimulate the cartilage regeneration. Noth et al. reported that the MSCs cultured on Col-I gel encapsulated with BMP-2, and TGF-β1 underwent distinct chondrogenic differentiation (Nöth et al. 2007). Further, by adding cytokines, and culturing on Col-I gel in vitro for 21 days, MSCs exhibited enhanced production of cartilage matrix, and cartilage-related gene expression (Yokoyama et al. 2005). In the case of Gel, the viability of chondrocytes in GelMA-based hydrogels was maintained for 8 weeks, and ECM accumulation of chondrocytes was also observed in vitro. Further, incorporation of HA, Gel/HA hybrid gel showed high mechanical strength and enhanced chondrogenic gene expression (Schuurman et al. 2013; Levett et al. 2014). Col gels not only promoted cartilage repair in vitro but also facilitated cartilage regeneration in vivo. Transplantation of MSCs encapsulated Col scaffolds into large cartilage defects showed natural cartilage-like tissue after 48 weeks, while the control group did not show any improvement in the regeneration (Wakitani et al. 1998). Moreover, MSCs (Jiang et al. 2017) and primary chondrocytes (Pulkkinen et al. 2013) in Col gels also exhibited good efficiency in cartilage repair. Wang et al. reported that the gelatin-hydroxyphenylpropionic acid conjugated hydrogels modulated chondrocyte functions, including cell proliferation, expression of chondrogenic genes, and production of chondrogenic ECM in vitro. Further, the gross appearance of repaired cartilage tissues was observed in vivo (Wang et al. 2014a). It has also been reported that the CS/GP gels support high cell proliferation of chondrocytes in vitro. The subcutaneous injection of bovine articular chondrocytes encapsulated in CS/GP gels in vivo showed higher secretion, and expression of extracellular cartilage matrix, and cartilage-specific genes, respectively (Chevrier et al. 2011; Chenite et al. 2000). The autologous chondrocytes (Schneider et al. 2011) or MSCs (Enea et al. 2015) in Col gel have also been studied against the cartilage defect area of patients. The clinical results, including knee pain scores, and histological grading scores, exhibited that the Col gel improved regeneration of cartilage (Guenther et al. 2013). However, the low mechanical properties of Col and CS limited their successful use in clinical application. Therefore, natural polymers were modified or combined with other polymers to attain good stability and promote cell migration (Zhang et al. 2019). In a particular study, the electroconductive hydrogel was made using pyrrole-based oligomer-chitosan, in which they observed conductivity contributed to cell activities, and cartilage repair (Kashi et al. 2018). The CS/GP/hydroxyethyl cellulose hydrogel encapsulated with BMMSCs, and TGF-β3 exhibited better chondrogenic differentiation (Naderi-Meshkin et al. 2014). Huang et al. used a demineralized bone matrix (DBM) to enhance the mechanical strength of CS hydrogels. Their experiments revealed that CS/DBM gel promoted the chondrogenic differentiation with higher GAG, and gene expression of hyaline cartilage markers in vitro (Huang et al. 2014). Further, the in vivo study of cartilage defect repair in rabbits showed an improved distribution of allogenic chondrocytes and regeneration of cartilage tissues after 2 years of post-transplantation evidenced by MRI (Man et al.

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Fig. 11.4 Schematic demonstrates the use of BMMSCs seeded peptide nanofibrous scaffolds for cartilage regeneration. In vivo results showed the development of hyaline-like cartilage in the defected site (Sun et al. 2018)

2016). Sun et al. designed a hydrogel from DCM-derived scaffold functionalized with a self-assembling peptide Ac-(RADA)4-CONH2/Ac-(RADA)4GGSKPPGTSSCONH2 (RAD/SKP) to deliver BMMSCs to the articular cartilage defect (Fig. 11.4). Their results showed DCM-RAD/SKP facilitated the recruitment of endogenous MSCs within the defect site after a week of implantation. Further, in vivo results showed that the DCM-RAD/SKP helped to achieve hyaline-like cartilage in the defected site (Sun et al. 2018).

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11.3.3.3.2 Nanofibrous Scaffolds To reproduce the mechanical strength of natural articular cartilage, 3D woven fabric composite scaffolds have been developed, and incorporated with bioactive molecules, including cartilage derived matrix, and growth factors for promoting chondrogenesis, and cartilage formation. This approach provides structural integrity and prevents the cartilage from the complex physiological loading after implantation. These scaffolds could also mimic the anisotropic, nonlinear, and viscoelastic properties of articular cartilage. The nanofibers fabricated from natural, synthetic, and composite polymers have been investigated to repair the cartilage defect (Ghasemi-Mobarakeh et al. 2013; Bhattacharjee et al. 2015). PCL nanofibrous scaffolds were used to culture fetal bovine chondrocytes. The results revealed that the chondrocytes grown on PCL were supported efficiently with the spindle or round-shaped cells during culturing compared with flat surface observed on a 2D tissue culture plate. The rate of proliferation was found to be 21-folds increase compared to 2D (Li et al. 2003). Further, when the MSCs were cultured on PCL nanofibers, they found that the mechanical property of the scaffolds provided the support to maintain the phenotype of chondrocytes (Li et al. 2005). In a particular study, Wright et al. compared PDLA/PLLA and PDLA/PCL nanofibers. Primary canine chondrocytes cultured on PDLA/PCL produced both Col-II and PGs although mechanical strength of PDLA/PLLA-hydrogel scaffolds was more (Wright et al. 2014). Nanofibrillated chitosan (NC) was added to PCL to improve the viscosity, conductivity, and electrospinnability. PCL/NC scaffolds showed remarkable enhancement in both tensile strength and Young’s modulus (Fadaie et al. 2018). The nanofibers using synthetic polymers alone were not able to achieve efficient chondrogenesis. Therefore, they were modified with natural polymers. PCL nanofibers were coated with CS and incorporated with TGF-β1 to improve the cartilage regeneration in a rabbit model. It was observed that levels of GAGs content and cartilage yield were higher (Casper et al. 2010). Electrospun nanofibers of PLLA were modified with Gel to improve their biocompatibility with chondrocytes. In vitro studies reported that the fabricated scaffolds increased viability, proliferation, and differentiation of rabbit articular chondrocytes. Subcutaneous implantation of the cell-scaffold matrices with autologous chondrocytes confirmed the formation of ectopic cartilage tissue after 28 days of post-operation (Chen and Su 2011). Alginate grafted hyaluronate (Alg-g-HA) was used to functionalize PLA nanofibers to improve the compressive modulus, followed by cartilage regeneration in vivo (Mohabatpour et al. 2016). Recently, it was reported to trigger the chondrogenic differentiation of MSCs, and the cartilage regeneration in vivo using poly (ε-caprolactone)/polytetrahydrofuran (PCL-PTHF urethane), and Col-I from calfskin (PC). PC nanofibers supported chondrogenic differentiation in vitro, and cartilage regeneration in vivo due to specific blockage of the NF–kappa B signaling pathway to suppress inflammation (Jiang et al. 2018). In a very recent study, Silva et al. used Kartogenin (KGN—a small molecule known to promote MSCs chondrogenesis) with poly(glycerol sebacate) (PGS)/poly(caprolactone) (PCL) aligned nanofibers for cartilage regeneration. The KGF incorporated nanofibrous

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scaffolds significantly increased cell proliferation, chondrogenic differentiation, and chondrogenic markers gene expression levels (Silva et al. 2020).

11.3.4 Intervertebral Disc (IVD) 11.3.4.1 The Structure of the IVD IVD is a fibrocartilage that lies between bony vertebral bodies, provides flexibility, load transfer, and energy dissipation to the spine. The disc is observed as the largest avascular component of the human body, wherein the cells of adult lumbar discs are about 8 mm away from the nearest blood supply. The disc receives its nutrition through the gradient diffusion of solutes. The complex structure of IVD has three specific tissue components, such as annulus fibrosus (AF), nucleus pulposus (NP), and end plates. The AF comprises a series of loosely connected concentric layers (lamellae) of highly oriented Col-I tissue that enclose the NP. The collagen fibers within these lamellae are arranged parallel to each other at an angle of approximately 60 to the spine axis and are oriented in opposite directions in successive layers. The peripheral region of the AF is connected to the posterior longitudinal ligament of the vertebral body through Sharpey’s fibers. The inner area of AF is attached to NP horizontally, and the collagen fibers are spread vertically into the endplate (Fig. 11.5). The NP is a gelatinous matrix, which consists of a random network of Col-II fibrils along with highly negative sulfated charge PG. The end plates lie at the cranial and caudal interface between disc and the vertebral body. They consist of translucent, hyaline cartilage. These plates are directly connected to the inner lamellae of the AF without collagenous connections. The central region of the endplates is thinner, which connects to the vascular buds lying in the medullary space and facilitates diffusion of molecules into and out of the disc (Yang et al. 2016; O’Halloran and Pandit 2007; Nerurkar et al. 2010). Fig. 11.5 Structure of intervertebral disc. It is highly organized and oriented in concentric rings consisting of annulus fibrosus (AF), which is on the outside cover of the nucleus pulposus (NP). The middle (green) portion is the endplate (Yang et al. 2016)

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In healthy discs, approximately 70 and 20% dry weight of the AF and NP is made of Col. The 50% of the dry weight of the NP accounts for PG, and a tiny percentage of PG found in AF. The largest PG in the disc matrix is aggrecans (chondroitin sulfate and keratin sulfate). The chondroitin sulfate is a hydrophilic molecule, which gives translucence appearance and resilient compressive strength to the disc. The aggregates are formed in the central region of the NP, by the accumulation of aggrecan molecules on a hyaluronan strand. Further, elastin is also found throughout the IVD. However, it is more densely located parallel to the collagen fibers in the interlaminar space of the AF, and forms cross-bridges between the lamellae of outer AF. In the NP, these elastin fibers are randomly oriented, which helps for the restoration of the shape of the tissue after deformation (Yang et al. 2016). The IVD has a low density of cells compared to other tissues, which accounts for 1% of the total disc volume. The outer AF contains fibroblast-like cells that are embedded in collagen fibril and positioned themselves according to the orientation of collagen fibril. The inner AF and cartilaginous endplates are rich in chondrocytelike cells (Yang et al. 2016).

11.3.4.2 Need for the Disc Repair Disc degeneration is a common disorder among 97% of individuals of 50 years or older age. The back pain caused by degenerative disc disease (DDD) has a lifetime prevalence of 80%. The back pain is rarely life-threatening, but the medical expenses are more. Clinical treatments for DDD have been addressing the alleviation of symptoms, but the restoration of function is not focused. The surgical standard for the treatment of axial low back pain is fusion, and this works by eliminating motion across the joint space. This technique often fails to alleviate pain, also affects the adjacent discs. Total disc arthroplasty is a recently approved surgical option to maintain the segmental motion that relieves pain but does not restore disc height or its original load-bearing capacity, which affects its long-term success. Therefore, a strong need is there to treat the back pain. 11.3.4.3 Tissue Engineered Disc IVD has a complex structure that consists of three tissues: NP, AF, and the CEP. All of them are involved in the disease process and may require repair. This adds a level of complexity to biological disc repair. Therefore, the biomaterials approach involved for the repair includes a combination of scaffold fabrication techniques. 11.3.4.3.1 Nanofibrous/Hydrogel Scaffolds for Disc Repair Mimicking the collagen fiber architecture of native AF tissue is considered to be the “holy grail” of AF repair, and NP replacement has been extensively studied using on injectable hydrogels that can restore disc height and motion segment stability to an IVD with an intact AF (Bowles and Setton 2017). AF cells and MSCs seeded onto aligned nanofibrous scaffolds adopted a spread morphology, and the cells aligned parallel to the supporting nanofibrous scaffold. The alignment promotes the efficient growth of cells, and provides support for ECM deposition by AF precursor cells. Thus, electrospun scaffolds for fabricating the aligned fibers are considered an

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Fig. 11.6 Self-assembled collagen nanofibrous injectable hydrogel repaired the IVD regeneration (Uysal et al. 2019)

efficient template for AF repair/regeneration (Nerurkar et al. 2010). The bovine AF cells were cultured on aligned PCL nanofibrous scaffolds for 8 weeks in vitro. The authors found an increase in the tensile modulus from 25 to 50 MPa (Nerurkar et al. 2008). Shamshah et al. fabricated a nanofibrous scaffold using a blend of PCL and PLLA to mimic AF tissue. Their bilayer structure of PCL: PLLA fibers were aligned at 30 to resemble the native AF lamellar layers. This was made into a circular lamellae structure using a custom-built cell sheet rolling system (CSRS). The scaffolds cultured with bovine AF cells showed that the cells were viable and grown in the orientation of fibers, and increased deposition of Col-I (Shamsah et al. 2019a, b). Zhou et al. studied the influence of surface topography in the differentiation of AF-derived stem cells (AFSCs). They fabricated PLLA nanofibers with a different orientation and fiber diameter. AFSCs were nearly round, resembling the shape of chondrocytes-like cells in the inner region of AF, on scaffolds with small fibers. However, they became spindle-shaped on scaffolds with large fibers, mimicking the morphology of cells in outer AF. Meanwhile, the upregulated expression of the Col-I gene in large fibers, while enhanced expression of Col-II, and aggrecan genes were observed on scaffolds with small fibers. Further, culturing AFSCs on a heterogeneous scaffold by overlaying membranes with different fiber sizes resulted in formation of a hierarchical structure approximating native AF tissue. Uysal et al. use collagen-based self-assembled nanofibrous structures for IVD regeneration (Fig. 11.6). The collagen peptide based hydrogel was injected into the degenerated rabbit IVDs induced more glycosaminoglycan and collagen deposition compared to controls. Further the bioactive scaffold showed functional recovery of the IVD degeneration (Uysal et al. 2019). The implantation of autologous NP cells or stem cells seeded scaffolds delays degeneration in the disc experimental model. However, implantation of AF or NP will not reverse changes such as CEP calcification, and thus the restoration of structural functionality can be obtained, when all three tissues have degenerated. In these settings, the optimal approach would be to replace the entire disc with proper cells seeded scaffolds (Bowles and Setton 2017). Du et al. recently engineered a biomimetic AF-NP composite using the core/shell structure of PCL microfiber/ alginate hydrogel. The circumferentially oriented PCL microfibers were seeded

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with AF cells, and alginate hydrogel encapsulated with NP cells as a core. The authors observed AF cells spread along the orientation of PCL microfibers, and NP cells colonized in the alginate hydrogel similar to native IVD. Additionally, both the cells were observed in the AF and NP region without any migration. Further, in vivo implantation of the scaffolds showed progressive tissue formation indicated by deposition and organization of ECM (Du et al. 2019). In another study, Gloria et al. reported the multiphasic nucleus/annulus scaffold for IVD regeneration using PCL and Col/HA. Their design involved a PCL scaffold with tailored architectural features as AF and a cell-laden Col-low molecular weight HA as NP. Culturing of MSCs on the scaffolds showed efficient viability (Gloria et al. 2020). The core/shell of CS hydrogels as core and poly (butylene succinate-co-terephthalate) copolyester (PBST) electrospun fiber as the shell has also been reported to regenerate IVD. The in vivo results of cell-seeded scaffolds in New Zealand white rabbits showed PBST outer AF provided mechanical support for the whole IVD, and the bilayered scaffold simulated the natural structure of the IVD (Yuan et al. 2019).

11.3.5 Tendon Repair and Regeneration 11.3.5.1 Structure of Tendon Tendons are formed from the spontaneous aggregation of Col into complex architecture. The aggregation of multiple Col molecules results to form primary fiber bundle (subfascicle), a secondary fiber bundle (fascicle), a tertiary fiber bundle, and finally produce tendon. Tendon is composed of 95% of total Col-I, which is 60% of the total dry mass of the tissue. This organized hierarchical architecture of tendon prevents mechanical load and enabling locomotion. Tenocytes are present in the ECM matrix, which remodels its microenvironment in response to the mechanical load. This is mediated by the action of matrix metalloproteinase (MMPs), which is necessary for the regeneration, and function of the tendon tissue. The ECM of the tendon is composed of PGs, GAGs, glycoproteins, and collagens. The ECM of the tendon acts as a lubricant to surrounding tissues, which ensures the easy gliding of fibers during mechanical deformation (Walden et al. 2017). Approximately 95% of the tendon cells accounts for tenocytes, which are embedded in the aligned collagen fiber bundles, and the epitenon and endotenon. The tenocytes synthesize the collagen, essential for the hierarchical architecture of the ECM. The remaining 5–10% is constituted of progenitor cells, chondrocytes necessary for enthesis formation, synovial cells, and vascular cells, including SMCs and ECs needed for blood vessel development. The tendons undergo self-healing process, when it is subjected to excessive stain. The healing process involves intrinsic and extrinsic mechanisms after the injury. The intrinsic and extrinsic pathway is associated with the proliferation of fibroblasts from the epitenon and endotenon and recruitment of inflammatory cells and fibroblasts from surrounding tissues, respectively. This results in the migration of cells to the site of the lesion and forms a new matrix. Further, the healing process is characterized by three overlapping phases: (1) hematoma formation, (2) inflammatory response after injury, and (3) regeneration

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confirmed by overpopulated myofibroblasts, synthesis of ECM, and finally a remodeling of ECM. The remodeling phase alone takes up to 1 year, results in the formation of randomly aligned Col. However, the healed tendon tissue is not equivalent to its native, which gives raise to impaired function of the tissue. Also the newly formed tendon will be weaker and prone to rupture. This condition is known as tendinopathy, which needs a clinical solution. For the patient, this results in a reduction of mobility, an increase in pain, and morbidity (Walden et al. 2017).

11.3.5.2 Need for Tendon Repair Tendinopathy arises from excessive mechanical loading, in which the tendon is not able to withstand further tension. It is caused by multiple factors such as age, gender, disease, occupation, and physical training. Tendinopathy can be classified as either acute or chronic. Acute, and chronic, occurs due to excessive overload and degenerative condition persistent over time, respectively. Degenerative tendinopathy is characterized by a decrease in Col-I, and increase in Col-III, which indicates the weakening and decreased tensile strength of the tendon. This causes friction and pain (Walden et al. 2017). Current treatment for tendon injuries involves the management of pain rather than healing of the underlying damage. For example, non-steroidal anti-inflammatory drugs (NSAIDs) such as ibuprofen are prescribed, wherein inflammation is the key factor for the management. Also, exercise and mobilization therapy are also prescribed. Ruptures of more than 5 mm have limited healing capacity, and the surgery is the only option for the physically fit and active patients. Although surgery is considered to produce satisfying results, problems such as alleviation of pain, reduced mobility, morbidity, or high risk of re-rupture for patients limits the successful application. Overall, there is no accepted standard treatment option for tendon injuries, and there is an obvious unmet clinical need (Walden et al. 2017). 11.3.5.3 Tissue Engineered Tendon In order to reproduce the spatial and temporal signaling profile during the healing process of the tendon tissue, suitable TE scaffolds are needed as a supporting matrix to deliver proteins, genes, and cells in a sustained manner. TE scaffolds can be used as mechanical support, or as carriers for deliverable factors. The ideal TE scaffold aims to closely mimic the native ECM architecture, and biomechanical properties of native tendon tissue. The TE scaffolds needed for tendon regeneration can be fabricated from either natural or synthetic polymers (Walden et al. 2017). 11.3.5.3.1 Injectable Hydrogels Systems Hydrogels are a prominent biomaterial scaffold for the delivery of regenerative morphogens along with cells. They can be easily modified to reproduce the mechanical strength of the natural tendon with controlled degradation rate and injectability. In addition, the hydrophilicity and large water-absorbing capacity of hydrogels are also advantageous to reproduce tendon ECM. Hydrogels encapsulated with proteins and cells for their controlled delivery followed by regeneration have been

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investigated in tendon models (Qiu et al. 2011; Chen et al. 2011b). The injectable hydrogel can be used as a space-filling agent, sealant, and a barrier to form the adhesion sites to repair smaller tendon defects (Kuo et al. 2014) (Walden et al. 2017). In a study, decellularized ECM (dECM) from tendon based hydrogels was used to treat tendon injury. After implantation of ASCs loaded thermoresponsive hydrogels facilitated to form aligned Col fibers, and as well as invasion of cells in vivo (Farnebo et al. 2014). Col has also been used along with dECM to deliver MSCs for the tendon repair. This approach was effective in maintaining the extracellular scaffold matrix, and for expressing Col-I, and oligomeric cartilage matrix (Martinello et al. 2014). The thermos-responsiveness of CH/GP has also been used to develop new peritoneal ligament tissue in dogs (Ji et al. 2010). Recently, Yang et al. reported the delivery of tendon stem cells (TSCs) using chitosan/bglycerophosphate/collagen(CS/GP/Co) hydrogel for the tendon repair in Achilles tendon defect rat model. They found improved healing in terms of biomechanical and biochemical properties after 4 and 6 weeks post-injury (Yang et al. 2018). In the case of flexor tendon injury, peritendinous adhesions complicates the situation for the patients. There are various approaches that have been proposed using biomaterials. In a study, alginate was reported as an anti-adhesive material, which resulted in complete tendon healing with the longitudinal remodeling of collagen fibers (Namba et al. 2007). Similarly, HA also has been reported to reduce the peritendinous adhesions following partial-thickness tendon injury in rabbits. Liu et al. reported that Seprafilm® and Carbylan™-SX cross-linked HA-derived film promoted healing of a flexor tendon injury without the formation of fibrosis after 3 weeks, wherein Carbylan™-SX cross-linked gel was found to prevent the tendon adhesion formation efficiently (Liu et al. 2008). 11.3.5.3.2 Implantable Fibers System Implantable materials aim to reconnect the tendon within larger midpoint ruptures and suture it, after removing the debris of the tissue. This can be used for partial and full ruptures. The TE scaffold has the ability to bear the mechanical strain, and the healing occurs. Fiber-based TE scaffolds have been extensively used for the tendon repair due to ease of suturing at the site of injury, and ease of incorporating various biomolecules into pre-aligned structures resembling native tendon tissue. The in vitro study of MSCs cultured electrochemically aligned collagen fibers showed efficient differentiation to a tenogenic lineage and increased expression of tendonspecific markers such as scleraxis and tenomodulin (Kishore et al. 2012). Further, Col-I fiber scaffolds showed to be conducive. Additionally, infiltration of tenocytes, and white blood cells (WBCs) were observed in the collagen surface (Kay et al. 2001). In another study, collagen electrospun fiber implants, cross-linked using riboflavin, and UV and combined with bovine platelet gel were implanted to treat large Achilles tendon defects in rabbits (Moshiri et al. 2015). Schnabal et al. seeded the MSCs in aligned CS fibers and investigated the tendon regeneration. The MSCs produced phenotypic change, and they revealed a 50-fold increase in Col -I expression in vitro. In vivo study in rats exhibited an improved tendon healing response when MSCs seeded CS fibers were implanted (Schnabel et al. 2009).

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Fig. 11.7 The schematic diagram demonstrates the use of the biomimetic sheath membrane for tendon repair (Liu et al. 2012)

The synthetic polymers such as PLA, PCL, and PLGA have also been examined as TE scaffold for the repair of tendon injuries (Mouthuy et al. 2015). In a study, PCL/ silk fibroin was used to produce the fibers in the view PCL to provide the required mechanical strength, and the silk fibroin yarns to support the cell proliferation of tendon tissue. In another study, PGA electrospun fibers were used to deliver muscle-derived cells from mice to promote the formation of neo-tendon. Their results after 12 weeks revealed that muscle cells were able to produce stronger tendon tissue, with a weak expression of MyoD, thicker collagen fibers exhibiting increased maturity, and increased expression of tendon-specific markers when compared to control (Naghashzargar et al. 2015). Liu et al. fabricated a biomimetic bilayer core-sheath membrane consisting of HA–loaded PCL (HA/PCL), and PCL fibrous membrane as the inner and outer layer, respectively, as shown in Fig. 11.7. The implantation of the scaffold in a chicken model revealed that peritendinous adhesions were reduced, and tendon gliding was improved. Similarly, Deng et al. used PGA/PLA fibers as a scaffold for the implantation of ADSC to treat 3-cm long

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defects of Achilles tendon in rabbits. The scaffolds have PGA fibers aligned in the inner region, and that was covered by PLA as an outer layer to form a mesh-like structure. The ADSCs seeded PGA/PLA scaffolds were implanted into the rabbit, and the repair was studied after 12, 21, and 45 weeks of post-surgery. They observed neo-tendon tissue formation, which was comparable to the native tendon. The tensile strength of the scaffold treated tendon showed a 60% improvement. In addition, the cell-free scaffolds group produced fibrotic tissue, disorganized collagen, and increased inflammatory cells (Deng et al. 2014).

11.3.6 Skeletal Muscle Tissue Engineering 11.3.6.1 Structure of Skeletal Muscle Skeletal muscle is composed of myofibers, blood vessels, nerves, and ECM connective tissue (Fig. 11.8). The myofiber is the basic structural unit of a skeletal muscle. Many undifferentiated muscle cells (myoblasts) connect together and become elongated, and produce multinucleated myotubes. The myotubes mature and form myofibers, wherein the nuclei in central position move to periphery position. The myofibers are surrounded by connective tissue, and bundled together to form a skeletal muscle. The connective tissue forms the supportive framework, which are responsible to maintain muscle shape, and synergistic movement during contraction

Fig. 11.8 The structure of skeletal muscle (Simon 1998)

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and relaxation. The skeletal muscle consists of two fiber types, oxidative slow-twitch (type-I), and glycolytic fast-twitch (type-II, further classified into type-IIA and typeIIB) that are categorized based on their myosin heavy chain (MHC) isoforms. The metabolism of both fiber types varies, which is dependent on the composition of their sarcomeres. The type-I fibers are primarily responsible for less intense, but more prolonged activities such as long-distance running, whereas the type-II fibers are responsible for short, but high-intensity work such as sprinting. Thus the proportion of the two fiber types present in the individual muscle decides its contractile properties (Qazi et al. 2015; Gilbert-Honick and Grayson 2020). In addition to myofibers, and connective tissue, the skeletal muscle consists of a dense network of blood vessels to keep the muscles metabolically active by distributing oxygen and nutrients. The contraction and relaxation of muscle are mediated through neuromuscular junctions, which induce contractility by delivering impulses to the myofibers (Qazi et al. 2015; Gilbert-Honick and Grayson 2020).

11.3.6.2 Need for Skeletal Repair/Regeneration Approximately 40–45% of an adult human body mass accounts for skeletal muscles that generate forces, and enable voluntary movement and locomotion. The repair and regeneration of the skeletal muscle occur through three main phases. These include (1) The destruction/inflammatory phase: The rupture of myofibers due to injury undergoes necrosis. The rupture of fibers releases cytokines and growth factors such as tumor necrosis factor- α (TNF-α), fibroblast growth factor (FGF), insulin-like growth factor (IGF), interleukin-1β (IL-1β), and IL-6 released, which activate and recruit the inflammatory cells residing in the muscles. Then, inflammatory cells such as neutrophils enter to necrotic fibers regions through blood vessels. This cocktail of soluble factors along with inflammatory cells influence the behavior of satellite cells (SCs) during the repair and remodeling phase. (2) The repair phase: In this stage, initially phagocytosis of necrotic muscle fibers followed by removal of cellular debris occurs with the help of macrophages. The phagocytosis of ruptured fibers is caused by pro-inflammatory macrophages (M1), which further activates and recruitments muscle progenitor cells, including SCs. The anti-inflammatory macrophages (M2) promote proliferation and differentiation of SC-derived myoblasts to form new muscle fibers. Also, the invasion of nerves, blood vessels, and the deposition of scar tissue by fibroblasts occur during the repair phase. (3) The remodeling phase: This stage involves the reorganization, and merging of neo-myotubes with existing myofibers. The remodeled scar tissue starts its contractile function again (Qazi et al. 2015). The damage of skeletal muscles is common and is caused by contusions and strains during sports activities; or lacerations and trauma due to accidents. Moreover, almost every surgery also causes damage to skeletal muscle. The skeletal muscles have an inherent ability to regenerate by activating resident multipotent stem cells, known as SCs, along with the mechanisms mentioned above. Also, the rate of repair via SCs activation is dependent on the severity of the injury. Large surgical trauma, which results in the scar tissue formation and fatty muscle degeneration, is known as iatrogenesis. In case of severe trauma the regeneration process is hindered by the

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formation of fibrous scar tissues (Qazi et al. 2015; Gilbert-Honick and Grayson 2020). Therefore, loss in muscle functionality due to severe trauma needs a clinical solution.

11.3.6.3 Tissue Engineered Skeletal Muscle The phases of muscle regeneration are same irrespective of the type or complexity of the injury, but the mode of treatment may differ depending on the type, function, and location of the muscle tissue. The biomaterial-based TE has been evolved to solve the issues characteristic to that type of injury (Qazi et al. 2015). 11.3.6.3.1 Injectable Hydrogels for Skeletal Muscle Regeneration Vandenburgh et al. reported that cells cultured on Col hydrogel produced a dense assembly of highly contractile myotubes, equivalent to neonatal myofibers (Vandenburgh et al. 1988). The difference in contraction forces generated by myoblasts at various developmental stages has also been studied. The cells cultured in 3D collagen gels showed lesser contraction forces initially as myoblasts in the gel remained as rounded morphology. However, when the cells grew, the contraction force increased as the myoblasts formed the myotubes (Cheema et al. 2003). Guo et al. recently reported the hydrogels produced from dextran-graft-aniline tetramergraft-4-formylbenzoic acid and N-carboxyethyl chitosan and used it for skeletal muscle tissue engineering. The in vivo study in the volumetric mass loss model showed the hydrogels improved muscle regeneration (Guo et al. 2019). A number of other papers have also been reported that mammalian muscle cells grown in 3D Col gels exhibited integration of myoblasts, the formation of striated multinucleated myofibers, and myotendinous junction formation (Rhim et al. 2007; Smith et al. 2012). Rossi et al. designed a HA–photoinitiator (HA-PI) complex strategy to deliver either SCs or muscle progenitor cells (MPCs). The SCs encapsulated gels implanted mice showed a major improvement in muscle structure in terms of formation neo-myofibers, compared to the control. (Rossi et al. 2011). The growth factors-laden hydrogels have been found to improve skeletal muscle regeneration. The delivery of myogenic and angiogenic growth factors incorporated Alg hydrogel into ischemic mice hindlimbs promoted functional muscle regeneration by stimulating myogenesis, angiogenesis, and reinnervation (Borselli et al. 2010). In another study, Kuraitis et al. encapsulated SDF-1 in Alg microspheres. Further, they were incorporated into an injectable collagen-based matrix for the treatment of ischemic hindlimb muscle. This local delivery increased the migration of progenitor cells, and also improved the recruitment of angiogenic cells (Kuraitis et al. 2011). Goldman et al. reported laminin-111 incorporated HA-based hydrogel (HA + LMN) as a putative myoconductive scaffolds. The delivery of a reduced dose of minced muscle graft (50% of VML defect) within HA + LMN in a rat tibialis anterior muscle VML model resulted 42% improvement in peak tetanic torque production compared to unrepaired VML affected limbs. However, the strength was not improved (Goldman et al. 2018). Hagiwara et al. studied the implantation of only myoblasts, myoblasts+Gel hydrogel microsphere, and myoblasts+ basic

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fibroblast growth factor (bFGF) + Gel hydrogel microsphere in GFP-positive myoblasts transfected thigh muscle defect model in Sprague–Dawley (SD) rats. Four weeks after transplantation, GFP-positive myoblasts were found to be integrated into the recipient’s muscle and to contribute to muscle fiber regeneration in all groups with the significantly higher expression level of GFP in the Gel+bFGF +myoblasts implanted group and the survival rate of myoblasts in this group was also found to be better (Hagiwara et al. 2016). Recently, Wang et al. reported shape memory Alg hydrogels along with growth factor and myoblast to treat injured tibialis anterior muscles in mice. The growth factor and myoblast loaded hydrogel showed regenerated neo-myofibers with reduced fibrosis and improved muscle forces. (Wang et al. 2014b). 11.3.6.3.2 Nanofibrous Scaffolds for Skeletal Muscle Regeneration Electrospun aligned fibers replicate the anisotropic structural organization of elongated myofibers in skeletal muscle, and play a crucial role in morphogenesis. These fibrous guidance cues vary from nanometers to micrometers in size, and stimulate cytoskeleton alignment, striated myotube formation, and expression of myogenic proteins. In a study, CS/PVA nanofibrous scaffold was assessed for cell attachment and mechanical support during regeneration. The authors found that cells had good viability after 20 days of culture and suggested to use as an engineered muscular graft (Kheradmandi et al. 2016). Liu et al. studied PCL electrospun fibers coated with mussel-inspired poly norepinephrine (pNE) (PCL/pNE) for muscle regeneration. The PCL membrane with a smaller diameter of fibers, and a relatively larger surface area was found to promote cell proliferation both in vitro and in vivo. After implantation of the scaffolds into the defected site, the site was found with regenerated muscle layer as well as fibrous membranes were integrated with the surrounding tissues. The mechanical strength of the regenerated muscle was also found equivalent to the native tissue (Liu et al. 2017). Similarly, conductive scaffolds fabricated using adhesives such as PCL/DOPA studied for muscle regeneration. The aligned conductive fibrous matrices promoted myoblast differentiation effectively, which was evidenced by the upregulation of myogenic markers (Tang et al. 2019). On the other hand, fibers fabricated using dECM were also studied for muscle regeneration. In a study, Smoak et al. had outcomes with the limitations associated with dECM such as physicochemical properties. They tailored fiber orientation and degree of crosslinking of these dECM scaffolds, which resulted in achieving better mechanical properties with degradation kinetics (Smoak et al. 2019). The delivery of growth factors for muscle regeneration has been extensively studied (Creaney and Hamilton 2008). In a study, Liao and Leong reported growth factor-releasing co-axial electrospun fibers for efficient muscle regeneration. The sustained delivery of angiogenic or lymphangiogenic growth factors served to stimulate the lymphatic or vascular systems to enhance the FVIII transport from the implant site into the systemic circulation. Further, the construct was integrated well with the host tissue within 1 month (Liao and Leong 2011). The 5-azacytidine (5-AZA) incorporated PCL electrospun fibers were studied for in vitro myogenesis.

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Fig. 11.9 Aligned dECM Nanofibers and PLGA Struts for efficient maturation of Human Muscle Progenitor Cells (Lee et al. 2019)

The hMSCs were differentiated and formed mature myofibers in PCL/5-AZA scaffolds compared to PCL scaffolds (Fasolino et al. 2017). A review article by Pantelic and Larkin clearly describes the need for stem cells for efficient muscle regeneration (Pantelic and Larkin 2018). The aligned nanofibrillar collagen scaffolds regenerated muscles efficiently in ablated murine tibialis anterior muscle. They also significantly enhanced the density of perfused microvessels compared to the control group. (Nakayama et al. 2018). Honick et al. studied the ability of murine myoblasts seeded electrospun fibrin/alginate scaffolds to regenerate the muscle structure in VML defects. They found myoblast-cultured scaffolds regenerated the muscle with high myofiber and vascular densities after 2 and 4 weeks. In contrast, scaffolds without cells lacked muscle regeneration (Gilbert-Honick et al. 2018). Lee et al. reported the fibers fabricated using dECMMA with PLGA to promote skeletal muscle cell maturation. The multiscale dECMMA/PLGA composite scaffold significantly promoted the myotube formation of human muscle progenitor cells compared to control scaffolds, as observed in Fig. 11.9 (Lee et al. 2019). Recently, Narayanan reported the effect of topography on muscle regeneration. They fabricated aligned fiber scaffolds with fiber diameters ranging from 335  154 nm to 3013  531 nm using PLGA. The larger diameter having scaffolds supported the myogenesis efficiently in vitro. They performed the in vivo study in the dystrophin-deficient MDX mouse model. Implantation of the primary myoblasts cultured fiber mat resulted in the production of dystrophinpositive myofibers network in tibialis anterior muscles (Narayanan et al. 2020).

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Future Perspective

The strategies involved for soft tissue engineering have been evolved in last few decades, which has brought few successful grafts such as Integra for skin engineering; Collagraft® from Nuecoll Inc. (a mixture of collagen with HA and β-TCP), Agili-C™ (a mixture of aragonite and hyaluronic acid for the cartilage part) for cartilage tissue engineering, etc. The artificial grafts available in the market have been used to treat chronic wounds with or without cells. The TE grafts currently available in the market use dECM or Col or synthetic polymer PLGA. Although they have proven to be an efficient acellular graft, invasion of microorganisms into the graft and the host tissue is a serious risk during the usage. The current research using cell-laden grafts could be advantageous over the acellular grafts, yet the shelf-life, transportation, and infection of the graft are major issues for their usage. Moreover, well-established laboratories with advanced instruments are required to maintain them, which will make the cost of the product so high. The scaling up of graft production is also difficult. The lengthy clinical trials associated with government rules are the other obstacle in translating the products from bench to bed. In addition to the optimization of polymer properties, research should also be focused on the molecular interplay, and cross-talk between cells, stem cells, ECM, growth factors, miRNA, and other bioactive factors to avoid the complications associated with remodeling stages of grafting after implantation. Although few research groups are working on developing products after studying the cross-talk between cells, and molecular interplay, none of the products has come to the market, and the strategies have not been fully optimized yet. Therefore, smart strategies such as easily scalability, the use of growth factors along with stem cells should be applied to develop soft tissue engineering grafts.

11.5

Conclusion

Soft tissue engineering addresses the challenges associated with the repair and regeneration of certain tissues that cannot be treated using classical medical approaches. Integrating the knowledge of chemistry, materials science, and biology is the key to develop a successful biodegradable soft tissue engineering scaffold. Further, the fundamental understanding of all the cell differentiation is favorable for tissue engineers to design the appropriate TE scaffold. Most natural polymers are having an inherent ability to interact with cells because of the presence of RGD moiety or PGs in it and making them ideal candidates for tissue scaffolds. Still, the mechanical strength and the stability of the natural scaffolds limit its successful application. Synthetic polymers can be easily tailored through monomer selection, molecular weight targeting, and functionalization. However, the most significant disadvantage with the synthetic polymers is the absence of cell adhering moieties. Therefore, combining synthetic, tunable polymer mechanical properties with natural polymer’s biological properties gives to design

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the optimum TE scaffold with respect to the tissue of interested with the desired mechanical properties and biological properties. The architecture and porosity required for the TE scaffold can be achieved by choosing the appropriate fabrication technique. The polymeric nanofibers fabricated through electrospinning can be effectively used for polarizing cells, especially for vascular tissue engineering. The injectable hydrogels can be used to fill irregular defects of cartilage and tendon. However, the stability and the density of hydrogels limit the porosity of the scaffolds. Therefore, advanced techniques like 3D bioprinting or integration of technologies like electrospinning and 3D bioprinting are required to develop successful soft TE scaffolds. Acknowledgement Author C.K.B.V thank the Indian Institute of Technology Madras, India, for the financial support in terms of fellowship and infrastructure support. We also thank the Science & Engineering Research Board (SERB), Department of Science and Technology (Project no. ECR/2017/003064), Government of India for financial support.

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Biomaterials for Specialized Tissue Engineering: Concepts, Methods, and Applications

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Divya Sree Kolla and Bhavani S. Kowtharapu

Abstract

Biomaterials are the most commonly used vital constituents of tissue engineering (TE) and regenerative medicine. Even though a large number of regenerative therapies are presently available, improvement of novel regenerative strategies for individualized TE are highly warranted. Development and implementation of innovative, novel tissue-specific biomaterial scaffolds to enhance their regeneration capacity is the need of the hour. Biomaterials in hard and soft TE has been widely explored and well documented. TE strategies towards sensory and endocrine tissues have emerged using the existing knowledge of biomaterials, fabrication methods, and biology of tissues. The present chapter mainly deals with application of biomaterials and their role in renewal of sensory and endocrine tissues with special focus on recent advancements in nerve as well as pancreatic TE. Application of carbon-based nanomaterials, and current developments in nerve TE using biomaterials in the form of modified nerve guidance conduits, conductive hydrogels along with improvised piezoelectric, electro conductive neural scaffolds have been discussed. We also covered implementation of engineered scaffolds including islet cell surface alterations in pancreatic TE and provide contemporary approaches for the development of artificial tissues. Keywords

Nerve regeneration · Nerve guidance conduits · Pancreatic tissue engineering · Biomaterial scaffolds

D. S. Kolla · B. S. Kowtharapu (*) Brien Holden Eye Research Centre, L V Prasad Eye Institute, Hyderabad, India e-mail: [email protected] # The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 B. Bhaskar et al. (eds.), Biomaterials in Tissue Engineering and Regenerative Medicine, https://doi.org/10.1007/978-981-16-0002-9_12

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Abbreviations ANG APC BDNF BMSCs BVSM CGA CGO CGRP CMCS CNS CNTF CNTs CS CS-PEG DPSCs DRG ECM ePTFE ESCs EVs FGF1 FN GAP-43 GDNF Gel GelMA GMSCs GNs GO Gr HA IBMIR IGF-1 iPSCs KOS LbL LN LOCS MEF MF MNPs MS

Artificial nerve grafts Activated protein C Brain-derived neurotrophic factor Bone marrow-derived mesenchymal stem cells Bisvinyl sulfonemethyl Collagen-glycosaminoglycan matrix Carboxylic-graphene oxide Calcitonin gene-related peptide Carboxymethyl chitosan Central nervous system Ciliary neurotrophic factor Carbon nanotubes Chitosan Chondroitin sulfate-incorporated starPEG nanocoating Dental pulp stem cells Dorsal root ganglion Extracellular matrix Expanded polytetrafluoroethylene Embryonic stem cells Extracellular vehicles Fibroblast growth factor 1 Fibronectin Growth associated protein-43 Glial cell-derived neurotrophic factor Gelatin Gelatin methacryloyl Gingival mesenchymal stem cells Graphene-based nanocomposites Graphene oxide Graphene Hyaluronic acid Instantaneous blood-mediated inflammatory reaction Insulin-like growth factor-1 Induced pluripotent stem cells Keratose Layer-by-layer Laminin Linear ordered collagen fibrous scaffold Mouse embryonic fibroblasts Magnetic field Magnetic nanoparticles Magnetic scaffold

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Biomaterials for Specialized Tissue Engineering: Concepts, Methods, and. . .

MWCNTs nAg nBG NCG NGC NGF NSCs NT-3 NTFs OECs PAM PANI PCL PCLA PCLF PDLLA PDLSCs PEDOT PEG PGA PGAt PGFs PHB PHBV PLA PLCL PLDLA PLGA PLLA PLLA-CL PNS PPV PPY PSS PT PU PVA PVDF PVDF-TrFE rGO SA SC SCI SDS SF

Multi walled carbon nanotubes Nanosilver Nanobioglass Nanoporous cellulose gels Nerve-guidance-channels/conduits Nerve growth factor Neural stem cells Neurotrophin-3 Neurotrophic factors Olfactory ensheathing cells Polyacrylamide Polyaniline Poly(ε-caprolactone) Poly(caprolactone-co-lactide) Polycaprolactone fumarate Poly(D,L-lactic acid) Periodontal ligament stem cells Poly(3,4-ethylenedioxythiophene) Poly(ethylene glycol) Poly(glycolic acid) PGA tube Phosphate glass microfibers Poly-3-hydroxybutyrate Poly (3-hydroxy butyrate-co-3- hydroxyvalerate) Poly(lactic acid) Poly(D,L-lactide-co-ε-caprolactone) Poly(L/D-lactic acid) Poly(lactic acid-co-glycolic acid) Poly(L-lactic acid) Poly(L-lactide-co-ɛ-caprolactone) Peripheral nervous system Poly(p-phenylene vinylene) Polypyrrole Poly(styrenesulfonate) Polythiophene Polyurethane Polyvinyl alcohol Polyvinylidene fluoride Polyvinylidene fluoride-trifluoroethylene Reduced graphene-oxide Sodium alginate Schwann cells Spinal cord injury Salidroside Silk fibroin

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SIS SPIONs SWNT TE TM TMV UBM UCMSCs US VEGF

12.1

Small intestine submucosa Superparamagnetic iron oxide nanoparticles Single-walled carbon nanotube Tissue engineering Thrombomodulin Tobacco mosaic virus Urinary bladder matrix Umbilical cord-derived mesenchymal stromal cells Ultrasound Vascular endothelial growth factor

Introduction

Biomaterials are tissue compatible, reproducible, biodegradable, porous, mechanically tunable molecules that can interact with cell structure, function, behavior, morphogenesis by mimicking and controlling surrounding target tissue native extracellular matrix (ECM) microenvironment and have been used in the repair and treatment of almost all types of tissues. Improvements in the design and characteristics of biomaterials substantially enhance target tissue regenerative potential. Porosity, pore size alterations in the biomaterial scaffolds change mechanical properties by adjusting cell attachment, increase nutrient availability through diffusion, and removal of waste products. Controlled degradation of the biomaterial at the target site supports native, infiltrating cells to secrete ECM components to fill the empty spaces and thus allows complete regeneration of the target tissue. Alterations in the surface chemistry and biological activity of the biomaterials augment cell adhesion, proliferation, and differentiation at the target tissue. Even though applications and regenerating capacity of biomaterials are very promising in tissue engineering (TE), their usage is still considered as an evolving therapeutic option due to the property of being different from the target tissue along with the limitations they have at the target site. Polymer-based biomaterials, derived from natural and synthetic sources, include polysaccharides and proteins with innumerable applications in TE. Collagen, matrigel are protein-based biomaterials derived from animal sources. With low mechanical strength and high degradation rates, they lack structural supporting ability during tissue regeneration. Chitosan (CS) and alginate are examples of polysaccharides with large scope in TE for encapsulation of cells. Synthetic polymer-based biomaterials are made up of polymerization using homopolymer or copolymer reactions from biodegradable or non-biodegradable constituents and during manufacture, their stereochemistry, length and polymer assembly can be customized by fine-tuning various chemical and physical parameters. Further, due to the possibility of fabrication into 3D scaffolds during their processing, synthetic polymers gained demand in tissue regeneration applications. Biomaterial scaffolds

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provide temporary structural integrity by substituting the ECM and serve as transport vehicle for cells as well as growth factor carriers besides enhancing the interaction of cells and biomolecules. Hydrogels are a different type of water-swollen smart biomaterials with suitable mechanical strength resulting from a network of 3D crosslinking hydrophilic polymers. They are considered as functional scaffolds in TE owing to their effective encapsulation of drugs as well as bioactive molecules. Smart polymeric biomaterials are capable of changing their chemical and mechanical properties reversibly in response to exogenous stimuli includes thermal, pH, and ionic. They represent a new class of biomimetic hybrid scaffolds and promote drug delivery, wound healing, tissue regeneration. Hybrid biomaterials exhibit combined benefits of both synthetic and natural polymer materials and are presently in high demand in various TE and regenerative medicine applications. Recently, decellularized ECM-derived scaffolds have been widely explored as they mimic tissue microenvironment analogous to that of native ECM by retaining their intricate native cell niche and represent the closest scaffold to the nature by preserving its complex composition. Further, ECM-derived scaffolds are emerging as sources for biomaterial engineering for improved adherence of cells, differentiation, biocompatibility, and inducing essential cell-specific responses (Xing et al. 2019). Similarly, cell-generated scaffolds containing collagen and elastin are reported to provide tailored mechanical properties in the range of native tissues. Additionally, biomaterials of human origin have the advantage of serving as building blocks for therapeutic applications by precisely representing natural characteristics of the native human tissue and enable TE during regeneration. Exploring the usefulness of the human-derived biomaterials and designing them as future therapeutics for clinical use largely constitutes the future of biomaterials (Chen and Liu 2016).

12.2

Biomaterials for Nerve Tissue Engineering

Neurons are the primary structural and functional machineries of the nervous system and contain axons, soma, and dendrites. Dendrites assist in transmitting electrical signals to soma whereas axons drive the impulses away from the soma. Glial cells function as auxiliary cells for neuronal function in which astrocytes and oligodendrocytes are present in the central nervous system (CNS) and Schwann cells (SCs) are present in the peripheral nervous system (PNS). Generally, in comparison to the PNS, CNS display high degree of anatomical complexity. Damage caused to the CNS and PNS either by injury or disease progressively leads to the lethal consequences due to the intricate structure and inadequate regeneration capacity of the neural tissues. In the CNS, during pathological conditions and after injury, increased population of reactive astrocytes is mostly responsible for the scar formation that eventually prevents regeneration of the damaged neural tissue. PNS nerves are also susceptible to injuries owing to the extensive innervation of peripheral nerves all over the body. End-to-end neurorrhaphy is contemporary standard procedure to bridge smaller nerve discontinuities whereas autologous nerve grafts,

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harvested from the patient’s own body, is generally considered most consistent strategy for nerve damage exceeding 1 cm. Nerve autografts have also been used to treat long-lasting larger and extended motor nerve flaws. Even though allografts and xenografts were also implemented for nerve reconstruction, existing limitations of each method, for example, requirement of long-term immune suppression during allograft therapy and danger of cross-species disease transmission during the use of xenografts, necessitated development of substitute approaches. Engineering of neural tissue with the help of substrate materials to develop immunologically inert, biodegradable, biocompatible neural scaffolds contributes as a potential substitute to the grafts for the treatment of damaged nerves (Sensharma et al. 2017). The plasticity of developmental innervation and neural circuit formation gradually decreases with age and following neuronal damage due to the inactive, disoriented reinnervation cues. Development of novel tailored biomaterials that augment fabrication of tissue microenvironments may help to re-establish the native niche and enhance neuronal regrowth. Designing of regenerative biomaterials capable of preventing apoptosis, inflammation, scar formation, inducing axonal regrowth, neurogenesis, and curative drug delivery are already in progress. In nerve TE, biomaterials are primarily used in the form of conduits, scaffolds to repair and fill the injured gap.

12.2.1 Nerve Guidance Conduits With the advent of TE, management of nerve related injuries has taken a new dimension by eliminating the need for surgeries or suturing. Literature reports confirm that inclusion and usage of bioengineered synthetic guidance tubes or artificial nerve grafts (ANGs) or nerve-guide-conduits or nerve-guidance-channels (NGCs) is a safe and feasible alternative approach for the regeneration of damaged nerve tissues. Long-distance communication by neurons occur by formation of neural circuits. Reconstruction of the emerging reinnervation signs is possible due to the innovative neural conduits and ANG or NGCs developed through TE and presently serves as an alternative therapy for peripheral nerve injuries. NGC with a hollow tubular structure is considered as a neural conduit whereas a conduit with an inside filling material is termed as a nerve scaffold. An ideal NGC should act as a hollow guiding bridge to physically guide regenerating axons across the lesion, augment the natural secretion of trophic factors from the injured nerve endings and should be able to counteract the fibroblast intrusion at the site of injury, permeable to nutrients, flexible to endure mechanical stress, porous, neuroinductive, neuroconductive, and biodegradable. Since neuronal maintenance relies on the presence of suitable microenvironment for trophic factors, cytokines, hormones, and ECM components, multifunctional NGCs designed and manufactured through novel combinatorial bioengineering procedures, to mimic the natural neural tissue microenvironment, are indispensable to improvise clinical outcomes following nerve damage.

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12.2.1.1 Biological Conduits Biological tubulization referred to the utilization of non-nervous autologous tissues as a scaffold to bridge a nerve gap. Non-nervous biological tissue-based conduits were employed as alternatives to repair short severed nerve gaps with varying success rates and comprise arteries, veins, and skeletal muscle. Nerve tubes made from processed biological tissue materials such as collagen and laminin also successfully used for specific applications. The disadvantages associated in using blood vessels and muscles as conduits, include tissue reaction, fibrosis, scar infiltration, lack of mechanical strength and their efficacy is only restricted to short nerve defects. These restrictions prompted emergence of conduits made up of novel synthetic materials. Currently, empty epineural tubes were also tested as biological conduits. Their neural origin, low antigenicity, high laminin, and vascular endothelial growth factor (VEGF) expression supports SC attachment and axonal elongation and functions better than muscle or vein grafts. Epineural conduits augmented with different supportive cells are also currently under development. 12.2.1.2 Synthetic NGCs Synthetic hollow NGCs are being widely used in nerve regeneration due to the prevention of fibroblast invasion, scar formation, and enabling accumulation of neurotrophic factors (NTFs). Generally used fabrication methods of NGCs include injection molding, dip coating, centrifuge casting, film rolling, extrusion, electrospinning, and braiding (Fig. 12.1) (Wang and Cai 2010). The first-generation hollow NGCs are synthesized from biologically inert, non-resorbable silicon or expanded polytetrafluoroethylene (ePTFE) materials. Since silicone is non-porous and biologically inert, its presence as conduit material lead to compaction and diminished transmissivity in axons. Due to this, its clinical usage gradually declined. Similar to silicon, usage of ePTFE also slowly declined. Regeneration events leading to axon regrowth within the hollow NGCs occur through various stages. Approximately 12 h after injury, during the initial fluid phase NGC chamber becomes full with fluid originating from the damaged nerve ends which mostly contains NTFs. Establishment of acellular fibrin polymer matrix occurs during the second phase within the first week and is followed by a third cellular movement phase where cellular components migrate through the matrix. Fibrin matrix offers a scaffold for the regeneration by implantation of migrated cells that include SCs, endothelial cells, fibroblasts, and perineurial cells penetrating the matrix from both nerve ends. Presence of NTFs along with cell migration across fibrin matrix leads to the initiation of axon regrowth towards the distal end in the fourth phase after 2 weeks. In the conduit, growth of fibroblasts and SCs is faster than the blood vessel formation and axon regrowth and axonal elongation solely depends on the prior existence of SCs. In the final myelination stage, nearly 4 weeks after injury, formation of myelin is observed from the surrounding SCs. Further, development of fibrin matrix is critical to initiate axon regrowth in NGCs. Failure of matrix formation has been observed when NGC is used to repair long gaps where the topography, cellular components of the fibrin matrix are completely influenced by the dimensions of the NGC (Fig. 12.2) (Muheremu and Ao 2015).

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Fig. 12.1 Schematic representation of extensively used NGC fabrication methods. Based on the mechanical properties of the polymers, different fabrication techniques are used for the threedimensional construction of NGCs. Reproduced and modified from Wang and Cai (2010)

Due to the fact that axonal fiber growth in hollow NGCs results in dispersion of regenerating axons and poly-innervation of diverse target tissues hollow NGCs are mostly preferred to fill small nerve gaps of approximately 10 mm (Du et al. 2018). Further, the first-generation conduits necessitate second-stage surgeries to remove the conduit due to fibrotic encapsulation of the conduit and nerve compression. Even though hollow NGCs display good cellular attachment and axon elongation they lack biological, physical, and chemical cues necessary for guiding axonal outgrowth. These disadvantages necessitate further improvement of NGCs with more innovative structural strategies.

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Fig. 12.2 Schematic representation of NGC and the nerve regeneration process. (a) Represents the macrograph of an example NGC, whereas (b) depicts its scanning electron microscopy micrograph. Scale bar: 400 μm. (c) illustrates the process of nerve regeneration through the NGC. By recreating a favourable tissue microenvironment and providing trophic support to the proximal as well as distal nerve stumps, NGCs prevent invasion of neighbouring tissue into the damaged site and enhances axonal regeneration. Reproduced from Shen et al. (2013), Muheremu and Ao (2015)

Second-generation hollow conduits are made of resorbable, biodegradable, biocompatible synthetic material with controlled mechanical properties along with specific topography of the tube wall structure. Further, they are simple hollow tubes without having any features of an autograft. Various resorbable materials, for example, collagen type I, gelatin (Gel), fibronectin (FN), keratin, silk fibroin (SF), CS, laminin (LN), hyaluronic acid (HA), poly(lactic acid) (PLA), poly(L-lactic acid) (PLLA), poly(glycolic acid) (PGA), poly(lactic acid-co-glycolic acid) (PLGA), poly(ε-caprolactone) (PCL), poly(L-lactide-co-ɛ-caprolactone) (PLLA-CL), poly (caprolactone-co-lactide) (PCLA), polyurethanes (PUs), tri-methylene carbonateco-ε-caprolactone, poly(D,L-lactide-co-ε-caprolactone) (PLCL), and polyvinyl alcohol (PVA) containing NGCs have been employed to avoid second-stage surgeries. Non-biodegradable polymers such as methacrylate-based hydrogels, polystyrenes were also employed as nerve conduit materials. The biological, mechanical, physiochemical properties of NGCs depend on the type of material and preparation method used. Hollow NGCs are designed and fabricated by different procedures including crosslinking, physical film rolling, injection molding, electro spinning, braiding, and melt extrusion. Novel, third-generation NGCs comprise controlled delivery and release of NTFs, electroconductive biomaterials, stem cells, SCs, ECM proteins, decellularized matrices, luminal fillers or surface micropatterning, hybrid scaffolds as guidance edifices (Fig. 12.3). Inventions in materials science, nanotechnology, and biotechnology

Fig. 12.3 Improvements, various designs of NGCs and their role in enhancing peripheral nerve regeneration. (a) Schematic diagram depicting the advancements in1st, second, and third generation NGCs. (b) Progression in the designs of NGCs. Hollow/non-porous NGCs represent the prototype with simple design and ease of fabrication. Multichannel conduits pose complications during manufacture but accelerate the axonal alignment during nerve regeneration. Porous surface enables the mass transport of nutrients and trophic factors. On the other hand, micro grooved and fiber-filled NGCs with their intricate architecture facilitates the axonal directional alignment. Hydrogel fillers supply a cell-friendly microenvironment and also allow the incorporation of biochemical cues for augmentation of nerve regeneration. Reproduced and modified from Gaudin et al. (2016) and Vijayavenkataraman et al. (2018)

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provide new approaches for the fabrication of new generation NGCs that address current limitations and gaps existing in nerve repair and regeneration (Zhang et al. 2012; Gaudin et al. 2016). Collagen type I is a fundamental structural protein found commonly in the body, basal lamina, in the endoneurium as fibrils and is one of the commonly used biomaterials in NGC fabrication. Maintaining the fibrillar structure of the collagen during fabrication results in enhanced mechanical strength, and absorbency. Collagen conduits are generally useful to link a 5–15 mm nerve gap and also resulted in good functional regeneration after nerve regrowth and deteriorates in 4–8 months. Type I collagen conduits are able to enhance regeneration and nerve repair in patients with up to 2 cm large-diameter nerve gap. Collagen nerve conduits filled with a blend of collagen-glycosaminoglycan (CGA) matrix, which mimics SC basal lamina, result in greater motor functional recovery when compared to the empty hollow collagen conduit (Lee et al. 2012). Multiple collagen-based materials that are FDA approved are available for clinical use in which NeuraGen® was the first collagen type I nerve conduit (Arslantunali et al. 2014). Collagen hydrogels prepared from the collagen fibrillar sheets are also implanted into 3-dimensional nerve conduits. Similarly, electrospun absorbable novel PCL-type I collagen nanofiber conduits exhibit good biocompatibility and have potential for nerve regeneration by preventing inflammatory reactions and recreating the normal morphology of myelin sheaths (Yen et al. 2019). Further, NGCs produced from a novel recombinant spider silk fabricated non-woven mesh and filled with microfluidics-produced collagen fibers enhanced nerve regeneration by forming neural networks and synapses (Pawar et al. 2019). CS-collagen-icariin composite scaffolds, prepared by blending and crosslinking of CS with collagen and icariin, provide superior structural support to SC proliferation, neurite elongation of neurons and thus function as potential cell carriers in neural TE (Yang and Chen 2013). Gelatin (Gel), prepared by thermal denaturation of collagen, has a wide range of regenerative applications in TE following crosslinking with various chemicals. Gelatin-based nerve conduits, prepared by crosslinking with genipin, or proanthocyanidin were also successfully evaluated as a guidance channels for repair of 10 mm nerve gaps and their functional recovery. Further, gelatin-based materials are non-cytotoxic, exhibit good conductivity. Fabrication of a biodegradable NGC with gelatin cross-linked with bisvinyl sulfonemethyl (BVSM) show strong mechanical properties with low toxicity. Gel-BVSM conduits demonstrate microstructures that form solid and closely packed walls, which prevents invasion of external soft tissues into the lumen of the conduits and can preserve the growth ability of SCs. Further results show that Gel-BVSM allowed moderate infiltration of macrophages, expression of calcitonin gene-related peptide (CGRP), other NTFs, such as insulinlike growth factor (IGF)-1, brain-derived neurotrophic factor (BDNF), and glial cellderived neurotrophic factor (GDNF) for nerve regeneration. These findings suggest that novel Gel-BVSM nerve conduit is a treatment option for injured peripheral nerve defects (Ko et al. 2017). This study paved ways further to fabricate novel Gel-graphene (Gr)-based NGCs and scaffolds comprising tailored 3D microstructures using 3D printing. Due to the electrical conductivity of Gr within

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the Gel-based 3D microstructure, application of electrical stimuli profoundly enhanced transdifferentiation of bone marrow-derived mesenchymal stem cells (BMSCs) to SC-like phenotypes along with their paracrine activity and nerve growth factor (NGF) secretion leading to promising nerve regeneration (Uz et al. 2019). Additionally, bio-composite conduits made from nanobioglass (nBG)-Gelnanosilver (nAg) particles by freeze-drying method are also tested for their efficacy and antimicrobial activity during nerve regeneration since nAg particles show greater antimicrobial activity by binding to microbial DNA and preventing their replication. nBG-Gel-nAg conduits potentially controlled scaffold associated bacterial infection during peripheral nerve regeneration (Koudehi et al. 2019). Fibronectin (FN), one of the main proteins of ECM, is also used in NGCs as a substance to release NTFs and its incorporation into alginate hydrogel matrix also contributed to improve peripheral nerve regeneration in tissue engineered poly-3hydroxybutyrate (PHB) conduits by enabling SC viability (Mosahebi et al. 2003). FN promotes opsonization of tissue debris, relocation, proliferation, and shrinkage of cells during the healing process. Later, FN matrix degrades completely to facilitate the progression of regeneration. However, in a diseased microenvironment, FN clearance is frequently disturbed, and its persistence at the wound site eventually leads to failure of regeneration. Thus, complete clearance of FN at the injury site is a prerequisite for recovery in any tissue. Since FN plays an important role in modulating SC function, its adsorption onto electrospun nanofibers facilitates the ability of topographical guidance cues to influence SC migration, neurite outgrowth which is necessary to design novel scaffolds and NGCs. Therefore, FN presence should be controlled and adjusted at the site of injury in order to define highly regenerative NGCs or scaffolds (Stoffels et al. 2013). Keratin is a valuable biomaterial for peripheral nerve regeneration and keratin hydrogel-based conduits, fabricated using the biomaterial derived from human hair keratin functions as neuroinductive by facilitating a strong and rapid nerve regeneration response through activation of SCs and were comparable to sensory nerve autografts. Further, Keratin-PVA scaffolds are also known to promote cell adhesion, proliferation and show enhanced density of SCs, axons, myelin thickness and were able to bridge 15 mm nerve defects (Teixeira et al. 2019). Silk fibroin (SF), due to its biocompatibility with cultured neurons and beneficial effect on the survival of SC, has properties to serve as a protein-based biomaterial for nerve TE. Silk protein-based conduits with high mechanical properties, biodegradability effectively bridge 11–13 mm longer nerve gaps by promoting axonal regeneration and functional recovery. Similarly, NGCs prepared from hybrid scaffolds comprising CS-SF fibers and SC-derived ECM-modified acellular scaffolds were investigated for joining nerve gaps of 10 mm (Gu et al. 2014). Angiogenesis within a nerve conduit plays an important role in nerve regeneration. SF-blended PLLA-CL nanofibers significantly enhance fibroblast proliferation, and VEGF secretion in vitro. Similarly, SF-PLLA-CL NGCs also enhances peripheral nerve regeneration by improving vascularization within the conduit (Wang et al. 2018). Water-insoluble SF NGCs, fabricated by freezing-induced structural transition, shows enhanced elasticity due to reduced β-sheet content, high content of silk I structure and makes

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it soft enough for peripheral nerve tissue regeneration (Li et al. 2019). A novel SF-based hybrid nerve conduit (SilkBridge™), with tubular scaffold comprising two electrospun inner, outer layers along with an intermediate textile layer characterized by a 3D architecture has been described to optimally balance biomechanical and biological properties. This scaffold is capable of sustaining cell proliferation, inducing greater neurite length in vitro and similarly induced cell colonization, blood vessel formation between different layers of the conduit wall in vivo along with regeneration of myelinated nerve fibers with a thin myelin sheath at the proximal end. These results prove that the SilkBridge™ nerve conduit is a biocompatible and biomimetic substrate (Alessandrino et al. 2019). Polysaccharide biopolymers are highly biocompatible and are used extensively in nerve TE as nerve regeneration conduits. Chitin is the main source of CS which represents the deacetylated form of chitin CS-based scaffolds, due to their cationic nature, capability of forming organized permeable structures and flexible mechanical properties, have many TE applications. CS tubes developed by fine-tuning of various acetylation concentrations could provide degradability, suitable microenvironment for renewing nerve tissue and thus allow functional and morphological nerve regeneration similar to autologous nerve grafts. However, CS tubes suffered from limitations such as high degradation rate and low mechanical stability (HaastertTalini et al. 2013). CS-conduits comprising CS flat membranes, fabricated by crosslinking with γ-glycidoxypropyltrimethoxysilan using solvent casting technique boosted nerve fiber regeneration and functional recovery almost mimicking the autograft (Fregnan et al. 2016). A novel approach to improve the performance of a CS-based NGC was developed by the insertion of fresh skeletal muscle fibers. CS-NGC lumen filled with skeletal muscle fibers (muscle-in-tube graft) produces soluble Neuregulin 1, a key growth factor essential for SC survival and dedifferentiation during the initial stage of regeneration and might be a suitable approach to repair lengthier nerve gaps or when SC degeneration is a restrictive factor for nerve regeneration (Ronchi et al. 2018). Further, it was reported that during the event of delayed nerve repair, CS-based NGCs demonstrate regeneration-supporting properties and have been approved for clinical usage (Reaxon® Nerve Guides) in reconstruction of nerve gaps up to 2.6 cm (Boecker et al. 2019). HA is another immunoneutral polysaccharide that has the capability to transform into various physical forms like hydrogels, flexible sheets, viscoelastic solutions, macroporous fibrillar sponges, electrospun fibers, and nanoparticulates. During the process of nerve regeneration, by assembling the ECM into a hydrated open lattice, HA facilitates migration and growth of the regenerating axons. NGCs made with photocrosslinked HA enable the growth of SCs and neurospheres in vitro and aids for axonal regeneration in vivo during peripheral nerve reconstruction (Sakai et al. 2007). Further, CS-NGC combined with HA display improved neural recovery by increase in number of axons, myelin thickness along with nerve fiber diameter (Li et al. 2018a). PGA is a rigid crystalline polyester with a high tensile modulus. PGA-based NGCs allow oxygen diffusion for efficient axon regeneration with an approximate degradation period of 6–12 months and is the first clinically used conduit. PGA

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conduits are useful to reconstruct the nerve gaps up to 3 cm and can produce results superior to those of the classic nerve grafts without donor morbidity. PGA is usually used together with other polymers during fabrication of conduits. The problems associated with these NGCs include their degradation before the total regeneration is accomplished and toxicity of the by-products released from lactic acid degradation (Lin et al. 2013). PLA-based NGCs fabricated using multi-layer microbraiding, immersion precipitation or micropatterning methods are capable of bridging as long as 20 mm nerve gaps with 80% functional recovery (Ni et al. 2013). The presence of macropores on the outer layer, and interconnected micropores in the inner layer on these conduits offer a higher outflow rate than inflow rate. PLLA is a crystalline form of PLA and PLLA-based NGCs are fabricated using extrusion method. They represent highly porous (83.5% porosity) conduits with consistently interconnected pore assembly (12.1 μm mean pore size) to serve as a scaffold for peripheral axonal regeneration (Evans et al. 1999). PLGA is another copolyester successfully tested as a nerve guide substance due to its weak inflammatory response and ease of fabrication. PLGA-based porous conduits with longitudinally aligned channels were fabricated using low-pressure injection molding technique. The topography of these conduits was planned to approximate the architecture of the nerves along with supporting the attachment of SCs. Biodegradable porous PLGA hollow fiber membranes, synthesized to use as a material in NGCs using wet phase inversion method, exhibit a homogeneous degradation pattern which is necessary to eventually accommodate axonal outgrowth and nerve regeneration (Wen and Tresco 2006). Fabrication of asymmetrically, porous PLGA-based NGCs with Pluronic F127 using modified immersion precipitation method results in the selective permeability and hydrophilicity of NGCs with inner surface having nano-size pores ( 50 nm) and outer surface having micro-size pores ( 50 μm) showed enhanced nerve regeneration. Due to the fact that inner surface avoids fibrous tissue infiltration but permeate nutrients and retain NTFs, whereas outer surface permits vascular ingrowth for effective supply of nutrients into the conduit. PLFA-F127 NGCs are able to heal nerve gap of  10 mm (Oh et al. 2008). PCL is one of the easiest biomaterials to process and manipulate into a wide range of forms due to its low melting temperature and greater viscoelastic properties. PCL-based ultrathin microporous biodegradable scaffolds have the capability of retaining the bipolar spindle-shaped phenotype of SCs in vitro and PCL-based NGCs are able to bridge 10 mm nerve gap with regenerated nerve tissue and penetrated SCs in vivo. A cross-linkable derivative of PCL, polycaprolactone fumarate (PCLF), is also a promising material in nerve TE. Asymmetrically porous PCL-Pluronic F127 NGC with dual NGF and low-intensity pulsed ultrasound (US) stimulation system serves as a promising approach towards clinical treatment of delayed, insufficient functional recovery of a peripheral nerve (Kim et al. 2013). PVA, with its hydrophilic nature, ability of swelling and absence of cellular toxicity is a very friendly polymer to living tissue. This water soluble, non-degradable polymer is also suitable for preparation of scaffolds and

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transformation into hydrogels. NGCs made from partially oxidized PVA exhibit increased axon density relative to neat PVA-based conduits and are biodegradable. Combination of PVA with other polymers has also been tested as scaffolds for their applications in nerve TE. Fabrication of CS-PVA nanofibers strengthened by singlewalled carbon nanotube (SWCNT-CS-PVA) nanocomposites using electrospinning illustrates that SWCNTs can augment the morphology, porosity, and structural properties of CS-PVA nanofiber composites and thus benefit cell proliferation rate. In addition, the cells exhibit their normal morphology during their incorporation with surrounding fibers (Shokrgozar et al. 2011). Electrospun PVA-CS nanofibrous biocompatible scaffolds with large pore sizes serve as potential matrices for nervous TE and repair. PVA-CS scaffolds reveal most balanced features to meet the specified basic requirements such as enhancing the viability and proliferation of nerve cells (Alhosseini et al. 2012). Further, polymeric nanofibers surface-modified with keratose (oxidative keratin, KOS) nanoparticles provide greater wettability, mechanical strength, and biocompatibility for nerve TE applications. Modification of PVA nanofibers with KOS nanoparticles by electrospray deposition after electrospinning led to the formation of KOS nanoparticles-coating PVA nanofibers (KNPs-PVA) that displayed improved cyto-biocompatibility on neural cells in terms of cell morphology, adhesion, and proliferation (Guo et al. 2018). PU is a class of polymers comprising organic units connected by urethane linkages. NGCs fabricated using scaffolds of PCL and poly(ethylene glycol) (PEG)-based block polyurethane polymers possess highly defined porous microstructure (pore size 1–5 μm; porosity 88%) and show strong cell compatibility, micro-morphology, hydrophilicity, nutrient permeability as well as mechanical strength to enhance tissue regeneration. PU nerve guide scaffolds exhibit much better nerve regeneration behavior similar to autograft (Niu et al. 2014). PHB macromolecules obtained as resorbable sheets, particles, films have been used in peripheral nerve regeneration. Due to their fast degradability, PHBs are designated as alternatives for manufacture of biodegradable plastics. Wrapping of resorbable PHB sheets around nerve defects support regeneration of axons with no adverse events or complications reported and PHB NGCs are suitable to bridge long nerve gaps up to 4 cm (Young et al. 2002). Fabrication of electrospun scaffolds blended with PHB and poly (3-hydroxy butyrate-co-3- hydroxyvalerate) (PHBV) demonstrate potential in regeneration of the myelinic membrane. Similarly, presence of type I collagen in the aligned PHB-PHBV nanofibrous scaffolds enhance cell differentiation and has many applications in neural TE (Masaeli et al. 2013).

12.2.1.3 Surface Micropatterning of NGCs Surface micropatterning and the insertion of ECM proteins in NGCs offer most appropriate nanostructure topography for adequate nerve cell development and simulate topographical features similar to the native nerve ECM. To facilitate improved functionality of NGCs in terms of regeneration speed and functional retrieval, the NGC must contain physical guidance cues and provide relevant biochemical signals. Electrospinning technique is frequently used to engineer matrices with desired imprinted micropatterns due to its capability to produce a greater

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area-to-volume ratio in the NGC, which leads to greater adsorption of cell adhesion molecules and subsequent enhanced cell attachment. Further, this method was also used to create nanofiber-aligned structures within the conduits to provide adhesion and guiding extension of neurons for nerve regeneration. Controlling the composition of the pore size in conduit wall, with inner microporous layer and outer macroporous layer is another option to attain bidirectional permeability. Engineering of the scaffolds by accommodating ECM proteins such as laminin, fibronectin, and collagen with biodegradable polymers alters the scaffold permeability from hydrophobic to hydrophilic, which is advantageous for controlling the cell function. Alignment of nanofibers in NGCs along with their composition also alters the nerve recovery at the injury site. For example, PCL nanofibers significantly improve motor function whereas aligned laminin blend nanofibers regain best sensory function (Gaudin et al. 2016). Development of novel micro or nanopatterned substrates allows in-depth analysis of contact mediated guidance mechanisms in neural TE. Preparation of multi-walled PLLA-based conduits using solvent casting, physical imprinting, and rolling-fusing methods creates intraluminal walls of the conduit with accurately defined grooves along the longitudinal axis to provide enhanced neurite alignment and regeneration on micropatterned PLLA conduits (Li et al. 2008). Further, effect of the exterior topography and mechanical properties of smooth, pitted, and grooved structures of ultrathin PCL-PLA mixed solvent-cast scaffolds, that are prepared for the inner lumen of NGCs, on cell morphology and alignment was investigated. The conduit cylinders were prepared by rolling the scaffolds around a mandrel using a thermal welding technique. Even though in vitro tests report decent cell attachment, proliferation on the PCL-PLA polymer scaffolds, the angle of the slope and the space in between scaffold grooves specifically affected cellular responses and morphology. Further, the PCL-PLA patterned scaffolds offer excellent mechanical properties stronger than the natural nerve. These results validate that nerve cell responses inside the conduits were susceptible to the microstructural geometry of the scaffolds (Mobasseri et al. 2013). Similarly, insertion of allied nanofibers is a promising strategy to accelerate nerve regeneration and construction of NGCs with highly aligned nanofibers improve their performance and enhance their repairing capacity as well as recovery of distal nerve ultrastructure as they can stimulate and promote SC orientation and axon growth (Quan et al. 2019).

12.2.1.4 NGC Luminal Fillers Luminal fillers are growth-guiding scaffolds in the lumen of the NGCs and play a regenerative role during the healing process. Regeneration with aligned nanofibers was greater than with randomly aligned fibers. Aligned luminal fibers improve nerve recovery through stabilization of the matrix. In addition, the aligned fiber NGCs augment the expression of ATF3 and cleaved caspase-3 at the middle of the regenerative matrix and at the distal nerve segment indicating activation of SCs (Quan et al. 2019). During fabrication of the fibrin gels, an increase in fibrinogen concentration simultaneously increased the number of fibrin strands resulting to decreased pore size and increased stiffness ensuing reduced neurite length. Similarly, higher concentration of collagen gels resulted to the increased stiffness of the

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gels and reduced neurite out growth length. Similarly, stiffness of the gelatin methacryloyl (GelMA) hydrogel substrate also shown to effect neuronal outgrowth along with cell viability, adhesion, and spreading (Wu et al. 2019). NeuraGen 3D® is a first collagen NGC containing a novel 3D luminal filler made of hydrogel matrix of CGA (chondroitin-6-sulfate). Addition of 3D luminal filler to NeuraGen® enhanced its regeneration potential (Lee et al. 2012). Acellular keratin hydrogel scaffolds, derived from human hair keratin, as conduit lumen filler in NeuraGen® collagen conduit enhanced peripheral nerve regeneration and motor recovery in a non-human primate model (Pace et al. 2014).

12.2.1.5 Stem Cell-Based NGCs Cellular components of tissue engineered nerve grafts include SCs, neural stem cells (NSCs), embryonic stem cells (ESCs), induced pluripotent stem cells (iPSCs), BMSCs as well as many other support cell types. Cell-based therapy performed with the help of nerve conduits has evolved as an alternative therapy of nerve repair in recent years. SC seeded NGCs have been successfully used in peripheral nerve regeneration studies (Euler de Souza Lucena et al. 2014). Encapsulation of cells in 3D scaffolds is one of the many neural TE procedures. Collagen, HA-based interpenetrating polymer network hydrogels combined with LN serves as a novel 3D neural regeneration scaffolds where hydrogel encapsulated SCs show increased secretion of NGF and BDNF with parallel alignment forming structures reminiscent of Bunger bands. This type of cell-laden 3D ECM-mimicking hydrogels hold promise in neural TE and is useful to study the fundamental mechanisms of SCneuron-ECM interactions (Suri and Schmidt 2010). The synergistic effect of NTE and drug therapy on peripheral nerve regeneration was demonstrated using PLGASCs-salidroside (SDS) NGCs. SCs cultured on engineered PLGA films in the presence of SDS, which is a phenylpropanoid glycoside extracted from Rhodiola rosea L with neuroprotective and neurogenic properties, showed increased cell proliferation and growth along with upregulated expression of NTFs such as BDNF, GDNF, and ciliary neurotrophic factor (CNTF). This shows that SDS exerts its neurogenic effect on SCs by regulating the expression of NTFs. Strategy of using engineered PLGA and SCs in NGCs in combination with SDS enables the therapeutic effect of SDS by promoting the NTFs expression and regeneration-associated genes, which subsequently leads to the enhanced axon, nerve regeneration, and functional recovery of myelinated nerve fibers (Liu et al. 2017a). Olfactory ensheathing cells (OECs), originated from the olfactory placode, possess many properties in common with SCs and promote regeneration by reducing scar formation. PLA-CS-collagen NGCs as well as PLGA NGCs seeded with OECs reported to enhance nerve regeneration by differentiating into SC-like cells and providing a permissive environment to facilitate axon growth. OECs treatment increased the number of regenerated neurons, improved nerve fiber diameter, morphology, myelinated nerve fibers, and myelin sheath thickness (Li et al. 2018b). Further, extracellular vehicles (EVs) derived from cultured OECs and fabricated into PCL-NGCs along with matrigel (PCL-OEC-EV-matigel conduit) also showed to

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enhance axonal nerve regeneration and promote neurologic functional recovery (Xia et al. 2019). BMSCs also have increasing applications in cell-based therapies including neural injury, disorders, and gained importance as support cells for peripheral nerve TE applications. LN-modified CS multi-walled nerve conduit together with BMSCs presents another novel nerve regeneration conduit matrix to bridge a 10 mm long nerve gap in which strategic embedding of LN onto the CS film improved cell adhesion (Hsu et al. 2013). Further, PGA tube (PGAt) containing BMSCs was also developed and tested for their efficacy in enhancing nerve recovery. Nerve regeneration was enhanced due to the presence of BMSCs in PGAt and BMSCs were also observed to be integrated in neural tissue (Costa et al. 2013). Gel-based 3D conduits with nanofibrous, macroporous, and ladder-like microstructures fabricated via combined molding and thermally induced phase separation technique to allow transdifferentiation of BMSCs into SC-like phenotypes were also examined to help facilitate neuro regeneration. Transdifferentiated BMSCs within 3D-ladderlike conduits secreted significant amounts NGF and GDNF and enhanced neurite outgrowth. The ladder-like conduits provide the most suitable environment for BMSC transdifferentiation to SC-like phenotypes and demonstrate the necessity of regulating the 3D microstructure to enable novel TE approaches involving stem cells for peripheral nerve regeneration (Uz et al. 2017). Generation of iPS-derived neurons in 3D fibrin-based scaffolds has been successfully achieved for spinal cord injury (SCI) repair and this technique can be employed as a platform for iPSC-based neural TE strategies (Montgomery et al. 2015). Further, NGC made of PGA and collagen has been used as a scaffold for grafting iPS cellderived neural cells showing retention of the cells at the target site and also improved motor neuron function and has application in treating PNS as well as CNS injury (Tomochika et al. 2019). A biodegradable PGA-based NGC filled with NSCs overexpressing GDNF showed enhanced nerve regeneration capability (Shi et al. 2009). PLA conduits grafted with CS-gold nanoparticles and immobilized fibroblast growth factor 1 (FGF1) represent a novel type of nerve conduits containing bioactive growth factors and display highest axon regeneration by protecting the FGF1 activity. Further, addition of NSCs on the micropatterned surfaces improved efficacy of the conduits compared with autografts (Ni et al. 2013). Similarly, LN-CS-PLGA NGC co-transplanted with SCs and NSCs, with notable mechanical strength and plasticity, promote migration and regeneration of nerve cells (Li et al. 2018c). PLGA-based NGCs containing collagen gel-embedded dental pulp stem cells (DPSCs), collagen-based NGCs containing DPSCs were reported to be well resorbed and also enhanced nerve regeneration along with their functional recovery highlights the therapeutic potential of DPSCs in peripheral nerve regeneration (Sasaki et al. 2011). Literature reports highlight that various kinds of NGCs loaded with umbilical cord-derived mesenchymal stromal cells (UCMSCs) show enhanced axon regeneration, myelin sheath thickness and promoted recovery in motor or sensory function (Cui et al. 2018). Further, UCMSCs delivered through PLCL membranes demonstrate successful differentiation into neuroglial-like cells and might improve clinical outcome especially after trauma to sensory nerves or nerve

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injuries with significant loss of nervous tissue that require entubulation or grafting (Bojanic et al. 2020). Though NSCs are now employed as a potential source for cell replacement therapy following SCI, reduced survival rate and low neuronal differentiation persist as main drawbacks to their usage. Biomaterials embedded with NTFs are known to promote the proliferation and differentiation of NSCs. Neurotrophin (NT)-3-immobilized SF matrices, developed by incorporating of NT-3 into a SF layer release NT-3 during SF degradation. Fabrication of PCLA scaffolds and conduits with immobilized NT-3-SF, release bioactive NT-3 continuously for up to 8 weeks and increased NSCs survival, neuronal differentiation in the spinal cord along with axonal regeneration and substantial functional recovery during SCI repair (Tang et al. 2014).

12.2.1.6 NGCs with Sustained Release of Growth Factors Delivery of bioactive NTFs to the injury site and their local release is also another TE approach to enhance axon regeneration. Integration of NGF in the nerve conduit and its release at the injury site contributes to improved myelination and functional recovery of the nerves. In order to accomplish these both functions at the same time, a novel composite NGC comprised of PLLA-CL and NGF in the form of coreshell structured nanofibers was developed. PLLA-CL was fabricated by coaxial electrospinning as shell whereas NGF as core of the conduit. This composite PLLA-CL/NGF conduit exhibited beneficial mechanical properties, biocompatibility and promoted effective nerve regeneration (Liu et al. 2011). Further, aligned PLA-CS fibers containing enhanced wet-state tensile strength, degradation tolerance, and lower swelling rate were also developed. Loading of PLA-CS with NGF-containing alginate solutions resulted to the sustained bioactive NGF release that eventually enhance neurite outgrowth during regeneration (Wu et al. 2017). Collagen conduits capable of releasing NTFs such as NGF, GDNF were fabricated using electrospinning approach and are also resulted to successful restoration of peripheral nerve defects (Madduri et al. 2010). Further, addition of PLGA microspheres capable of GDNF secretion immersed in a fibrin gel to the lumen of a conduit augmented the sensory and motor nerve regeneration (Tajdaran et al. 2016). Collagen-PCL-nano-nBG conduits loaded with NGF are also reported to enhance NTF activity, regeneration, and functional recovery at the site of injury (Mohamadi et al. 2018). Cylindrically woven PLGA filaments treated with pulsed oxygen plasma and coated with CNTF as well as CS were also successfully developed and tested for their efficacy in nerve repair. These PLGA-CS-CNTF conduits were capable of guiding the damaged axons through the lesioned area and repaired 25 mm long nerve defect (Shen et al. 2010). In another study, PLGA microspheres comprising VEGF were synthesized by the double emulsion-solvent evaporation technique and these VEGF loaded microspheres were implanted into the lumen of PCLF conduit after suspending in matrigel. VEGF promotes angiogenesis and also enhances neurons survival by stimulating neurogenesis and proliferation of SCs. PCLFVEGF conduits show degradation of PLGA microspheres and release of bioactive VEGF to increase angiogenesis in a controlled manner up to 4 weeks. Presence of

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matrigel as a suspending medium for VEGF loaded microspheres further allowed regenerating nerve fibers to grow robustly across the segmental defect without interruptions (Rui et al. 2012). Since clinically approved NGCs only offer undefined guidance of regenerating axons across nerve gap without any added advantages, development, and assessment of a bioresorbable drug delivery system capable of prolonged local delivery provides topological cues to direct regenerating axon growth via microgrooves. FK506 is a potent immunosuppressive, neurogenic agent with severe side-effects when delivered systemically. Creation of micropatterned-PLGA films embedded with FK506 by photolithography, capable of extended controlled release of FK506, preserved its neural bioactivity and promoted neurite extension. Thus, drug embedded micropatterned-PLGA films have potential applications in the construction of NGCs (Davis et al. 2018).

12.2.1.7 Conductive NGCs Conductivity is another important parameter in neural TE to acquire physiologically active, properly restored nerves. Conducting biomaterials based on carbon nanotubes, carbon nanowires, graphene, and metallic gold nanoparticles have been extensively studied for nerve TE applications (Goenka et al. 2014; Paviolo and Stoddart 2017). However, their non-biodegradability along with long-term in vivo toxicity and heterogenous dispersal of the conducting particles constrained their effective usage. Instead, conductive polymers display excellent conductivity and stimulate living cells or tissue by means of electrical signals. Electro conductive polymers such as polyaniline (PANI), polypyrrole (PPY), polythiophene (PT), and their by-products display good biocompatibility with extensive applications in nerve TE as neural implants and TE scaffolds (Guo and Ma 2018). Further, conducting polymers also can affect various cellular activities that include cell adhesion, migration, proliferation, differentiation, and protein secretion at the polymer–tissue interface in the presence or absence of electrical stimulations. Due to their brittleness, conducting polymers such as PANI and PPY are blended with synthetic or natural polymers such as PLGA, PCL, CS, and SF for their TE applications. PPY demonstrates exceptional electrical properties and has documented as a promising biocompatible scaffold material for neural TE. PLGA-PPY electro conductive scaffolds designed and produced by mounting PPY on aligned electrospun PLGA nanofibers to form electrically conductive nanofiber structures exhibit high neuro regeneration properties due to the combined effect of electrical stimulation and topographical guidance (Lee et al. 2009). In one study, towards enhancing the nerve regeneration along with conductivity, poly(D,L-lactic acid) (PDLLA)-based NGCs were developed. Construction of conducting composite nerve conduit with PPY and PDLLA using emulsion polymerization, dip coating methods and its stimulation resulted in the marked increase in neurite-bearing cells and the neurite length at the injury site (Xu et al. 2014). Electrically conductive PCLF-PPY hybrid scaffolds are stable and possess material characteristics required for application as nerve conduits since they significantly improve and direct neurite outgrowth upon electric stimulation (Runge et al. 2012). Further, biodegradable, conductive block copolymer of PPY (PPY-b) along with PCL is used to fabricate 3D porous NGCs

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using a novel electrohydrodynamic jet 3D printing. This novel 3D-printing method offers greater control over fiber diameter, pore size, porosity, and fiber orientation. 3D-printed PPY-b-PCL scaffolds are softer resembling the native human peripheral nerves and conductive with potential applications in preparation of cell-free or cellladen NGCs (Vijayavenkataraman et al. 2019). Development of a bioactive and biodegradable polyurethane conductive micropatterned scaffolds, to guide cellular orientation during nerve regeneration, using a micro-molding technique induced alignment and elongation of SCs over the micropatterned surface along with enhanced expression of NGF (Wu et al. 2018). Poly(3,4-ethylenedioxythiophene) (PEDOT) is a non-toxic, conductive polymer extensively used in the fabrication of tissue engineered scaffolds. Uniform assembly of PEDOT nanoparticles on a CS-Gel porous scaffold surface, via in situ interfacial polymerization for neural TE, increased electrical conductivity, hydrophilicity, mechanical properties, thermal stability along with cell adhesion, proliferation, and cellular neurite growth of neuronal cells by upregulation of growth associated protein-43 (Gap-43) and synaptophysin (Wang et al. 2017a). Further, cross-linked PEDOT substrate displayed cytocompatibility and improved NSCs differentiation into neurons along with longer neurites in the presence of electric stimulation (Pires et al. 2015). PANI-coated platinum electrode surface results to the aggregation of retinal fragments, reduced inflammation, scar formation and displayed enduring stability in terms of nanoparticle size, intactness, and morphology under electrical stimulation conditions without any corrosion (Di et al. 2011). Development of electrically active NGCs using synthetic conducting polymers is a novel approach to enhance peripheral nerve regeneration. Polyaniline-silk fibroin nanocomposite-based nerve conduits prepared by electrospinning a mixture of SF protein, PANI showed excellent nerve conduction velocity, thick myelination of axons and nerve regeneration (Das et al. 2017). Similarly, synthesis and coating of PANI nanoparticles (20–30 nm size) on the exterior of microtubes by layer-by-layer deposition and their insertion into 3D NGCs resulted to the fabrication of a conduit that can improve the regeneration of the nerve defect (Wang et al. 2020). Tobacco mosaic virus (TMV) solely infects plants and is entirely harmless for mammals. TMV is resilient against environmental changes, polar organic solvents, protonation and its excellent adjustability with functional materials make it advantageous for biotechnological applications. Fabrication of electroactive nanofibers on the surface of TMV by in situ polymerization of aniline using sodium poly (styrenesulfonate) (PSS) results to the TMV-PANi-PSS nanofibers that can support growth and differentiation of neuronal cells and augment neurite length. Further, electroactivity and topographical cues provided by the aligned TMV-PANi-PSS nanofibers in capillaries facilitate the outgrowth direction of neurites and lead to a bipolar cellular morphology (Wu et al. 2015). Blending of fibrous Gr-nanosheets along with sodium alginate (SA) and PVA mimic the ECM of the peripheral nerve by improving scaffold properties where Gr-nanosheets acts as electrical nanobridges to enhance electrical conductivity, robustness, and mechanical properties. Gr-SA-PVA hybrid scaffolds with

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exceptional toughness and electric conductivity show increased interactions and proliferation of nerve cells (Golafshan et al. 2016; Teixeira et al. 2019).

12.2.1.8 Carbon-Based Nanomaterial-Interfaced NGCs Graphene and carbon nanotubes (CNTs) due to their intrinsic properties of conductivity, flexibility, biocompatibility, interact with the nervous system, promote neural development and have many applications in neural TE. CNTs are cylindrically structured allotropes of rolled-up graphene layers with excellent thermal conductivity and ideal mechanical, electrical properties. In addition, they are non-biodegradable, biocompatible with neuronal cell adhesion and primarily function as substrates or scaffolds for neural cell growth and implants during spinal cord or brain injury. Along with neural cells other cell types such as stem cells, glial cells are successfully known to grow on CNT substrates. However, acute toxicity at the target site along with not completely understood biological interaction mechanisms of carbon-based nanomaterials with surrounding tissue microenvironment in vivo presently limits their usage. In nerve TE applications, SWCNTs and multi-walled CNTs (MWCNTs) are widely used. CNTs, by having morphological resemblance to neurites, improve the neuronal reactivity by making tight contacts with the cell membranes, which in turn acts as an exoskeleton that might support electrical shortcuts between the proximal and distal compartments of the neuron. CNT scaffolds deliver a superior material for guiding and promoting axonal growth in next generation NGCs. CNT scaffolds support in vitro survival, growth and development of primary hippocampal, dorsal root ganglion (DRG), cortical as well as cerebellar neurons along with neuron-like PC-12, NG108, Neuro2a, SH-SY5Y cell lines (Oprych et al. 2016). Implementation of anisotropic 3D CNT scaffolds such as CNT yarns, ropes or aligned electrospun CNT composite fibers were also reported that provide directional guidance to regenerating hippocampal, DRG neurites, SCs, as well as NSCs (Fan et al. 2012; Huang et al. 2012). CNT-GelMA hybrid hydrogels for creating 3D cell culture platforms were shown to be biocompatible and cell-responsive where presence of CNTs specifically strengthened GelMA hydrogels without altering their porosity or inhibiting cell growth. These photopatternable CNT-GelMA hybrid gels allow effortless assembly of microscale structures without involving any severe procedures and can be used for synthesis of composite 3D biomimetic tissue-like structures as well as scaffold material for cellular encapsulation (Shin et al. 2012). Further, 3D printed electroconductive nano-MWCNT scaffolds in combination with electrical stimulation showed synergistic effect on promoting NSCs proliferation and differentiation for therapeutic application in nerve regeneration (Lee et al. 2018a). Fabrication of NGCs based on freeze-dried silk, carboxylated SWCNTs along with FN nanofibers within the structure by electrospinning (SF-SWCNT-FN) displayed aligned fibers, appropriate porosity, and diameter. Further, their in vivo implantation showed SC growth, more myelinated axons in conduit as well as higher nerve conduction velocities and enhanced nerve regeneration (Mottaghitalab et al. 2013). Collagen-PCL fibrous scaffold and carboxyl MWCNT-based NGC prepared by electrospinning altered the composite scaffold’s hydrophilicity, mechanical

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properties and degradability and supported SC adhesion and elongation in vitro. In addition, MWCNT-enhanced collagen/PCL conduit successfully enhanced nerve regeneration, prevented muscular atrophy without invoking chronic inflammation (Yu et al. 2014). CNT-interfaced NGCs were synthesized and used for in vivo nerve regeneration studies. During fabrication, CNTs are first chemically secured onto the surface of aligned phosphate glass microfibers (PGFs). CNT-interfaced PGFs (CNT–PGFs) were then effectively positioned into a three-dimensional poly(L/Dlactic acid) (PLDLA) conduit. CNT-PGF-PLDLA conduit, after implantation showed regenerating axons crossing the scaffold and bridged the gap of 10 mm. However, the capability of the CNT-PGF-PLDLA NGC reported to be inferior to autologous nerve grafts (Ahn et al. 2015). Similar to CNT, graphene and its derivatives with flexibility, mechanical strength, and exceptional electrical conductivity has many applications as scaffolds, sheets in neural TE. Even though pristine graphene known to instigate synaptogenesis and promote neuronal growth, pure graphene sheets could not reiterate the basic neural microenvironment in vivo due to its smooth and chemically inert nature along with absence of biological signals. Since nanostructures such as fibers, grooves, ridges enhance neural cell adhesion, growth and alignment, fabrication of graphene derivatives in the form of graphene-based nanocomposites (GNs), for example, graphene oxide (GO), improved the roughness of scaffold exterior surface, electrical properties, enhanced differentiation of stem cells into neurons and thus played an indispensable role in neural TE applications (Chiacchiaretta et al. 2018; Bei et al. 2019). Layer-by-layer casting of 3D multi-layered, graphene-loaded PCL nanoscaffold conduit coated with polydopamine and arginylglycylaspartic acid showed significant electroconductivity, enhanced axonal regrowth and remyelination during peripheral nerve regeneration (Qian et al. 2018). Development of Gel-graphene based scaffolds and NGCs that contain 3D microstructures using 3D printing and application of electric stimuli resulted in the attachment of MSCs and their differentiation into SC-like phenotypes with increased NGF secretion (Uz et al. 2019). Fabrication of carboxylic-GO (C-GO)-composited PPY-PLLA (C-GO-PPY-PLLA) conducting NGC by electro chemical deposition of C-GOcomposited PPY nanoparticles on PLLA fiber films and application of electric stimulation at the injury site significantly improved the neurite length and neurite alignment, myelin sheath thickness and axon diameter of sciatic nerve during regeneration (Liu et al. 2019; Chen et al. 2019). Similarly, electroactive reduced GO (rGO) nanofiber scaffolds also modulate neuronal cell function and GelMA-rGO based conductive NGCs increased SC proliferation through activation of PI3K/Akt signalling pathway in vitro, promoted sensory, motor nerve regeneration and their in vivo functional recovery (Fang et al. 2020). However, GNs caused toxicity results to the DNA fragmentations, chromosomal aberrations in living cells along with increased reactive oxygen species generation. Additionally, biodegradation of graphene-based materials is another critical issue that needs to be addressed and novel strategies to combat these limitations are in presently progress to enhance the usage of GNs in neural TE (Bei et al. 2019).

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12.2.1.9 Ultrasound Treatment Following NGC Implantation Following injury, application of low-intensity US stimulation significantly promotes nerve regeneration and functional recovery. Non-invasive stimulation with US at the PLGA/F127 NGC-implanted site displayed more rapid nerve regeneration (Park et al. 2010). PLGA NGCs seeded with SCs also displayed similar enhanced nerve regeneration following US stimulation (Chang et al. 2005). These studies confirm that application of US following implantation of NGCs synergistically augment peripheral nerve regeneration. 12.2.1.10 Porcine Small Intestine Submucosa Made NGCs Small intestine submucosa (SIS), a derivative of the submucosal layer of porcine intestine, is an ECM that consists more than 90% of the acellular collagen (type I, III collagens) and numerous biological growth factors. The SIS graft is a promising biological material with biocompatible and non-immunogenic properties and acts as a natural conduit by providing favorable growth microenvironment to bridge the peripheral nerve defects. Further, when compared with other polymer-based NGCs such as PCLA, SIS-NGCs displayed higher nerve regeneration rate (Shim et al. 2015). Similarly, NGC containing ECM from the porcine urinary bladder matrix (UBM) also reported to serve as a novel scaffold in enhancing peripheral nerve regeneration (Nguyen et al. 2017).

12.2.2 Scaffolds for Nerve Tissue Engineering TE largely depends on scaffolds due to their nature of promoting cell differentiation and development. Collagen as a scaffold material has been explored extensively in the form of hydrogel, membranes, and films. Oriented collagen fiber scaffolds proved that aligned collagen fibers help in orienting the growing neurons in the right direction towards the distal end (Ceballos et al. 1999). Aponeurosis is a flat collagen rich extensive fibrous connective tissue that binds muscles together and nerve guidance material prepared from aponeurosis that is consisting of ordered collagen fibers showed to promote neurite outgrowth while reducing inflammatory response at site of implantation. Later the same aponeurosis was further adapted to be used in the form of linear ordered collagen fibrous scaffold (LOCS) filled in silicon conduit. These linear collagen fibers, chemically cross-linked and modified with laminin as well as growth promoting factors, has shown to support neurite outgrowth and aid in neurotransmitter synthesis (Cao et al. 2011). In another study, LOCS was also used in combination with PPY scaffold containing parallel-aligned multiple channels along with modified NT-3 with collagen binding domain. The individual advantages of collagen, PPY along with the trophic properties of NT3 were put into use in combination to bridge the gap in a SCI where it has helped to regain electrophysical and locomotive functions and also promoted myelination of the axons (Chen et al. 2019). Testing of a CS-based NGC filled with fibronectin stabilized, aligned, longitudinally arranged collagen hydrogels to heal a 15 mm nerve injury gap showed to significantly improve the gap compared to the hydrated

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control hydrogels. The fibronectin stabilized oriented collagen hydrogel showed better results as a promising internal filler for neural conduits in healing the gap, promoting axon growth and migration of SCs that help in myelination (GonzalezPerez et al. 2017). Electrospinning process can be adjusted to produce fibers of different thicknesses depending on the target tissue they need to mimic. Gel serves as an appropriate substrate for cell adhesion and differentiation and Gel scaffolds have best worked in their electrospun form in neural TE applications. Polyelectrolyte-based scaffolds and films prepared from CS-Gel and poly-l-lysine mixture promoted differentiation of PC12 cells, DRG neurons and show promising applications in neural TE (MartinLopez et al. 2012). Zhu et al. (2016) have used nano-bio-ink composed of modified GelMA, bioactive graphene nanoplatelets, and NSCs to form a 3D scaffold using 3D bioprinting technology, which has shown to support survival and differentiation of NSCs into neurons in culture. Fabrication of porous, electrically conductive scaffolds show a lot of promise for regeneration in nerve TE. HA doped-poly(3,4-ethylenedioxythiophene) (PEDOTHA) nanoparticles into a CS-Gel matrix (PEDOT-HA-CS-Gel) augmented the electrical and mechanical properties of the scaffold by reducing its porosity, water absorption and serves as an appealing conductive substrate for cell culture by providing cell adhesion, promote proliferation, differentiation, and support synapse formation (Wang et al. 2019). Similarly, fabrication of conductive PEDOT-CS-Gel porous scaffold without HA by an in situ interfacial polymerization method also reported to enhance PC12 cell adhesion and proliferation (Wang et al. 2017b). Ansari et al. (2017) tested the efficacy of the alginate-HA 3D hydrogel scaffold elasticity on various types of stem cells to differentiate into neural fate. They worked on periodontal ligament stem cells (PDLSCs), gingival mesenchymal stem cells (GMSCs), and human BMSCs using different concentrations of alginate and HA with varying elastic properties in the presence of NGF. Increasing the concentration of HA decreased the elastic modulus and this favored the proliferation of GMSCs, but the elasticity of the scaffold did not show any impact on PDLSCs and hBMSCs. This blended scaffold was more favorable for growth and proliferation of the cells rather than the only alginate scaffold. PDLSCs and GMSCs showed higher proliferation rate compared to hBMSCs indicating that they are easy to use compared to the stem cells from the bone marrow. Golafshan et al. (2016) worked on to develop a graphene, alginate, PVA fibrous eletrospun scaffold (Gr-AP) that promoted cell adhesion, migration, and proliferation of PC12 cell line compared to scaffold made of graphene alone. This hybrid Gr-AP scaffold showed sufficient mechanical strength, electric conductivity with enhanced cell interactions and has great potential in developing novel strategies for peripheral nerve regeneration. Fabrication of nanofiber mats containing CS blended with PVA and SWCNT (SWCNT-CS-PVA) by electrospinning showed increased bio-adhesive properties, porosity to allow nutrient exchange, cell attachment, migration as well as non-toxic, strong, longer shelf-life in supporting the cell growth in vitro (Shokrgozar et al. 2011). Direct-write 3D printing of stem cells inside biomaterials offers novel possibilities in TE applications. CS mixed with alginate

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and agarose in the right proportion was used along with hNSCs as polysaccharide-based bio-ink to produce a mini neural tissue construct by 3D printing technology, which subsequently supported the differentiation of hNSCs into GABAergic neurons and glial cells expressing astrocyte markers (Gu et al. 2016). Usage of human hair extracted keratin hydrogel scaffold as an acellular filler in the biodegradable conduit has shown significant increase in the axon density in non-human primates (Pace et al. 2014). Similarly, electrospun nanofibrous scaffold of keratin blended with PVA also showed to sustain neuronal survival and growth with low cytotoxicity and high cell adhesion properties (Mahanta et al. 2014). Zhang et al. (2012) experimented with scaffold containing not only oriented SF fibers but also that contains micropatterns, which helps as guidance cues, to support cell differentiation and migration. They worked with hippocampal neurons, which were shown to adhere, migrate, proliferate, and elongate in the uniaxial multichannel scaffold that assures its application as a biomaterial in nerve repair.

12.2.2.1 Synthetic Scaffolds Synthetic polymers are easy to produce into various forms and are used in different scaffold forms like gels, membranes, films with high mechanical strength. PEG has been explored to be used as a scaffold modified with hydrolytic lactide units from PLA and chemically cross-linked with methylacrylate units to encapsulate the neural precursor cells. Mahoney and Anseth (2006) have used this scaffold to monitor the growth and differentiation of isolated neural precursor cells and this scaffold has shown to promote differentiation of these precursor cells into glia and neurons without any additional supplements of NTFs or ECM apart from decreased rate of anoikis indicating that it can be further used in vivo. Liu et al. (2015) experimented with PLGA and PEG electrospun scaffolds to check their efficiency in treating SCI along with Sox-2 modified mouse embryonic fibroblasts (MEF) to take up a neural fate. The PLGA/PEG nanofiber membranes were rolled into 3D scaffolds along with Gel sheets and used as a support for the seeding of MEFs. The MEFs slowly differentiated into glia and neurons and when transplanted at the injury site, it has helped to recover into a normal functioning tissue. The combined effects of PLGA and PEG proved to be efficient along with natural Gel polymer that aided to form a 3D architecture in repairing the damage. 12.2.2.2 Piezoelectric Scaffolds Piezoelectric scaffolds are considered as smart scaffolds as they show stable variations and are responsive to the external stimuli. Since they are able to produce electrical charges or develop voltage during mechanical stimulation without any additional energy sources, electroactive and responsive piezoelectric materials show strong potential in axon regeneration. Further, as neurons are sensitive and responsive to the electrical signals, piezoelectric polymers permit uninterrupted transfer of electrical stimulation with their electrical activity during nerve tissue regeneration. Polyvinylidene fluoride (PVDF), polyvinylidene fluoride-trifluoroethylene (PVDFTrFE) are the widely used polymers due to their high piezoelectric properties and can

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regulate differentiation and outgrowth of neurons in vitro. In nerve TE, to reduce its hydrophobicity, PVDF has been modified by addition of various metallic nanoparticles, nanotubes as well as PEG and PVA polymers. These modifications are also known to influence piezoelectric scaffold nanostructure and positively affect neuron length and nerve tissue regeneration. Piezoelectric scaffolds promote differentiation of NSCs into neurons, enhance neurite outgrowth and also are used in NGCs for neural repair applications (Zaszczynska et al. 2020). PHBV is also used as piezoelectric scaffold along with collagen nanofibers and has been used as substrate to support neuronal cell growth as well as to promote axon-dendrite segregation (Prabhakaran et al. 2013).

12.2.2.3 Electroconductive Scaffolds Electrical stimulation is known to quicken the regeneration process especially in the neural fated cells as they are known to communicate majorly through electrical impulses. Therefore, incorporating conducting materials in the 3D scaffolds used to regenerate the damaged nervous tissue, assures faster and reliable repair process by reviving the conducting cells. PPY, PANI, SWCNT, MWCNT, PEDOT, and graphene are few of the conducting materials being used in TE applications. Shin et al. (2017) developed a 3D scaffold using catechol functionalized HA (HA-CA) hydrogel incorporated with CNT-PPY particles which showed moderate conductivity required for the regeneration process. Human fetal NSCs and human induced pluripotent stem cell-derived neural progenitor cells were used to test for the performance of this 3D scaffold. Improved differentiation along with upregulated calcium channels, activation of depolarization and an increase in the influx of calcium ions was observed in cells stimulated with electrical signals compared to the cells in the scaffold devoid of conducting materials implying that they can be further used for in vivo repair process using stem cells. PANI is another conductive material which is used in TE for its highly environmentally stable, biocompatible nature along with its ability to switch between conductive and resistive states. For example, Xu et al. (2014) successfully used PANI to coat diaminotriazine-based, hydrogen-bonding-reinforced hydrogel for the differentiation of NSCs. Other conducting materials also have found their use in the field of TE as nerve conduits, electrodes, etc. along with being drug delivering materials. Most of the studies have shown that the combination scaffolds showed much promising results in regenerating and repairing the damaged tissue when compared to the scaffolds fabricated using either only natural polymers or only synthetic polymers. 12.2.2.4 Conductive Hydrogels Incorporation of a soft hydrogel with a conductive polymer results in the formation of a conductive hydrogel. Conductive hydrogels, due to their tolerable electrical properties in providing electrical signals to the target cells, are the most effective biomaterials and gaining popularity as they are able to mimic the biological and electrical properties of tissues in the human body. Conductive materials such as PPY, PANI, PT, PEDOT, and poly (p-phenylene vinylene) (PPV) are widely used biomaterials in making conductive hydrogels. They are fabricated by using several

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conductive materials such as metal nanoparticles, carbons, and conductive polymers with the help of methods including blending, coating, and in situ polymerization (Min et al. 2018). In situ polymerization of electrically conductive PPY nanoparticles with nanoporous cellulose gels (NCG) resulted in the PPY-NCG aerogel that show enhanced biocompatibility and neurite outgrowth (Shi et al. 2014). Similarly, conductive polymer hydrogels prepared by combining SA, carboxymethyl chitosan (CMCS) and PPY (SA-CMCS-PPY) provide sufficient help for peripheral nerve regeneration and are also used as a filling material in NGCs (Bu et al. 2018). Chemical polymerization of PPY within ionically crosslinked alginate hydrogel networks resulted to the formation of alginate-PPY hydrogels that show enhanced cell adhesion and growth of BMSCs by increasing the expression of neural differentiation markers (Yang et al. 2016). MWCNT-PEG nanocomposite hydrogels prepared with 20% w/v PEG along with 0.1% w/v MWCNTs and stimulated with exogenous electrical fields resulted to the robust neurite outgrowth (Imaninezhad et al. 2018). Hydrogels made with the combination of CS and nano-GO promoted nerve cell growth up to 20% and presence of nano-GO in the hydrogel alters the pore size and increases mechanical strength (Jafarkhani et al. 2018). Synthesis of a composite hydrogel with the combination of polyacrylamide (PAM), GO, Gel, and SA (PAM-GO-Gel-SA), using the method of in situ free radical polymerization, has shown to provide support to the attachment and proliferation of SCs along with increased expressions of Sox10, GAP-43, and myelin basic protein (Zhao et al. 2018). Further, fabrication of conductive polysaccharide hydrogels based on agarose, CS, and alginate for the therapy of neural disorders is also reported (Alizadeh et al. 2019). Additionally, for the development of a 3D printable conductive hydrogel with enhanced electrical conductivity, a mixture of PEDOT and PSS aqueous solutions were freeze-dried and combined with polyethylene glycol diacrylate as the photocurable polymer base. The resulting conductive hydrogel was patterned on the substrate by stereolithography 3D printer. This study shows that the conductive hydrogels can also be used as a 3D printing base biomaterial for neuro TE applications (Heo et al. 2019).

12.2.2.5 Magnetic Scaffolds and Nanoparticles By merging various types of biomaterials like hydrogels, nanoparticles, and electrospun fibers with magnetic elements, magnetic composite biomaterials can be fabricated that have multiple physical and chemical signs to enhance the process of nerve regeneration (Funnell et al. 2019). Superparamagnetic iron oxide nanoparticles (SPIONs) are metallic-nanoparticles with exceptional magnetic properties and fabrication of Anisogel, a new type of anisotropic, injectable tailored microgel, with incorporated SPIONs arrange them unidirectionally in the presence of external low magnetic field and nerve cells seeded into these micro hydrogels also demonstrate alignment and directed growth (Rose et al. 2017). Conjugation of NGF to iron oxide magnetic nanoparticles (NGF-MNPs) and controlled use of magnetic fields helped to extend half-life of NGF resulting in the enhanced neurite outgrowth in PC12 cells and deliver NGF-MNPs to particular sites in vitro as well as in vivo including peripheral sciatic nerve and retina (Marcus et al. 2018). Liu et al. (2017b)

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fabricated a magnetic nanocomposite scaffold using MNPs and chitosan– glycerophosphate polymer and studied the effect of magnetic scaffold (MS) along with applied magnetic field (MF) on the survival of SCs and peripheral nerve injury repair. Collective application of MS and MF synergistically increased the viability of transplanted SCs and also augmented nerve regeneration and functional recovery (Liu et al. 2017b). Giannaccini et al. (2017) injected growth factor conjugated MNPs into NGCs to efficiently deliver NGF and VEGF to the injury site. This study displayed that injecting MNPs into NGCs is safe approach and MNPs remain long time inside NGC. Further, NGF and VEGF-conjugated MNPs robustly enhanced the peripheral nerve regeneration and the retrieval of its motor ability (Giannaccini et al. 2017). Further, Johnson et al. (2019) showed that electrospun fabricated magnetically responsive aligned PLLA scaffolds can be injected as a collagen or fibrinogen hydrogel solution into NGCs. Later, after magnetically focused them in situ using an external MF, the oriented magnetic fiber scaffolds offer reliable directional guidance to neurites within the hydrogel and significantly augmented neurite length as well as alignment within the NGC (Johnson et al. 2019).

12.2.2.6 ECM-Derived Scaffolds ECM by functioning as an attachment for cells also influences their differentiation and proliferation. Decellularized brain matrix contains complex array of functional proteins such as various types of collagens (collagen I, III, IV, V, and VI) along with laminin, perlecan and have many TE applications. Processed brain matrix serves as a scaffold for neuronal cell growth and seeding of neurons derived from human iPSCs on the brain matrix scaffolds show neuronal morphology and enhanced expression of neuronal markers (DeQuach et al. 2011). Similarly, ECM derived from the UBM also serves as a novel scaffold for augmenting sensory nerve growth and also reduces scar formation by inhibiting the fibroblast invasion to the injury site due to the presence of intact basement membrane on its surface (Nguyen et al. 2017).

12.3

Biomaterials for Pancreatic Tissue Engineering

The most important need for TE, especially for endocrine organs, is seen with the varying complications that arise in the patients of autoimmune disorder, type 1 diabetes mellites, where the beta cells of the pancreas are attacked by the individual’s own immune system, that subsequently leads to less or no insulin production resulting in hyperglycemic conditions. Though insulin therapy is a commonly used treatment option, pancreatic islet transplantation proved as an efficient therapeutic option and gained attention recently due to its long-term effects in treating the disease for those who are unresponsive to the insulin treatment. However, application of this therapy is limited due to the fact that the ideal islet encapsulation material and site of implantation are largely speculative and yet to be determined (Zhu et al. 2018). Encapsulation of islets could be a possible solution to safeguard the transplanted islets since encapsulation protects islets from the host immune response and also allows efficient diffusion of essential molecules such as

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oxygen, nutrients, glucose, insulin, and waste metabolites. The insulin-secreting islet transplantation follows the Edmonton protocol, where the human islet cells are isolated from cadaveric donors, purified and used for clinical transplantation into the patient (Salg et al. 2019). Furthermore, due to the global scarcity of the donor tissue material, there is an increased demand for novel TE approaches to evaluate and improve different biomaterials to host the islet cells without causing adverse immune, non-immune-related reactions. A lot of biomaterials have been tested for their efficiency in maintaining the islet cells in a healthy state for a long time and with minimal trigger to the immune system.

12.3.1 Biomaterials in Restoring Pancreatic Function Biomaterial scaffolds housing the pancreatic insulin-secreting cells for transplantation can be grouped into categories based on the nature of the material used that include (1) synthetic polymers such as hydrogels, (2) biopolymers such as ECM proteins or SF, and (3) recently established decellularized pancreatic ECM scaffolds. The ideal biomaterial for pancreatic islet transplantation should be with a pore size required to be able to encapsulate the islet cells with required trophic factors, immunosuppressing drugs (when required), compatible to be implanted, support survival of the graft and must be able to degrade when the tissue microenvironment for the graft is attained. The biomaterial used in pancreatic TE should also be capable of restoring vascularization following transplantation to provide sufficient oxygen and nutrient supply to the islets.

12.3.1.1 Biological Polymer Scaffolds In TE of pancreatic islets, biologically derived collagen, fibrin, alginate, Gel, and agarose were tested as scaffold materials. Due to their biodegradability, biocompatibility and non-immunogenicity, Gel incorporated porus interpenetrating polymer network scaffolds with polyvinylpyrrolidone are proven to be cytocompatible for encapsulation and prolonged maintenance of islets by providing adequate mechanical strength (Muthyala et al. 2010). Collagen has many applications in pancreatic TE due to its presence in the ECM. Modified collagen IV-scaffolds, collagen crosslinking with other polymers like CS, chondroitin-6-sulfate, LN supports islet survival, metabolism, glucose-induced insulin production and vascularization posttransplantation (Yap et al. 2013; Ellis et al. 2013). Fibrin usage in pancreatic TE is mainly based on its properties of self-assembly and elasticity. Fibrin scaffolds provide an excellent 3D extracellular support and encapsulation of endometrial mesenchymal stem cells in fibrin hydrogels differentiate them into pancreatic beta cells (Niknamasl et al. 2014). Additionally, fibrin is known to enhance insulin secretion and this capability is reported to be further increased in 3D fibrin scaffold gels (Kim et al. 2012; Niknamasl et al. 2014). The rapid and strong gelling capability of agarose along with its biodegradability has many advantages in islet TE. Arrangement of primary islets into pre-defined 3D aggregates having similar dimensions and equal cell number with the help of agarose microwell chips prepared on

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polydimethylsiloxane molds creates a suitable environment for optimal beta cell function and increase islet cell viability, function (Chowdhury et al. 2013; Hilderink et al. 2015). Agarose-based microwells, due to their precise fabrication, increase cell–cell interactions, islet function, and insulin secretion (Ichihara et al. 2016). In pancreatic TE, alginate has several applications in the form of beads, porous as well as 3D printed scaffolds due to its rapid crosslinking property. Alginate is a generally used polymer for microencapsulation and islets encapsulated with alginate beads function as bioartificial endocrine pancreas (Lim and Sun 1980). Later, multi-layer bioartificial pancreas containing islets encapsulated with two layers of alginate separated with poly-L ornithine allowed specific delivery of growth factors and blood vessel formation that eventually improved graft functional viability. The inner alginate layer encapsulates the islets whereas the outer alginate layer encapsulates angiogenic protein to initiate vascularization around the graft (Opara et al. 2010). Further modification of this two-layer encapsulation with redesigned thick outer alginate layer cross-linked with FGF1 resulted to the prolonged stability of the islets (Pareta et al. 2014). Furthermore, alginate encapsulated hESCs can be differentiated to pancreatic-islet-like cells and this method provides a translatable treatment option for type I diabetes (Richardson and Banerjee 2014). Similarly, clinical islet transplantation studies were also conducted to treat patients with type I diabetes with alginate encapsulated islets due to their long-term stability (Zhu et al. 2015a).

12.3.1.2 Synthetic Polymer Scaffolds PEG and its derivatives including poly ethylene glycol diacrylate and poly ethylene glycol dimethacrylate are being used to create biomimetic hydrogels that mimic cell– cell communication for enhanced survival and function of the isolated, dispersed pancreatic islets. Fabrication of hydrogels with desired cell-packing density using these materials can be adjusted based on their molecular weight and can also be quickly cross-linked using ultraviolet irradiation (Lin and Anseth 2011). A composite hydrogel of PEG and alginate, cross-linked by Staudinger ligation method was also employed for microencapsulation of islets to improve stability and cellular compatibility (Hall et al. 2011). PLA, PGA, PVA, PCL are among other scaffold materials that are being used to provide mechanical strength to the cells and when hybridized with a natural hydrogel these provide additional benefits of ease in attachment of the cells to the material and protection from immune system (Marchioli et al. 2016). The function and cellular homeostasis of islet cells is hugely altered by the hydrophobic nature of the scaffold material used and hence these synthetic polymers further need to be modified to increase hydrophilicity and their compatibility with the islet cells. Therefore, natural polymers that can form hydrogels including alginate, collagen, fibrin are individually used and also hybridized with these synthetic polymers for increased tissue functionality (Kumar et al. 2018). Even though synthetic polymers provide an advantage in terms of manufacturing, fabrication along with mechanical strength to the cells, they need to be improvised for functionality to use along with growth factors, ECM components and also pose risk during degradation as the released by-products cause substantial pro-inflammatory

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responses in vivo (Sheikh et al. 2016; Salg et al. 2019). Further, polymer-based scaffold material lacks porosity and hence can be detrimental to the formation of suitable tissue microenvironment for the transplanted islet cells and also inhibits cell infiltration (Mao et al. 2017). While the hybrid gels made of fibrin and PEG are also used in pancreatic islet transplantation, these scaffolds have difficulty in being reproduced as well as in the retrieval of the graft and hence have very limited clinical relevance (Mason and Mahoney 2010; Mao et al. 2017).

12.3.1.3 Silk Fibroin Fabrication of SF scaffolds along with other ECM materials like heparin serves as delivery vehicles for islet cells and also induces islet revascularization posttransplantation. Further, secretion of VEGF and an increase in angiogenesis indicate formation of organ microenvironment by eliminating conditions like cell death due to hypoxia (Mao et al. 2017). Encapsulation of pancreatic islets in SF hydrogels is known to enhance their survival and function (Davis et al. 2012). Likewise, during pancreatic islet transplantation, SF-based encapsulation, and heparin-releasing SF scaffolds (H-SF) displayed promising results in maintaining the health of islets by increasing collagen IV secretion which reduces cell apoptosis, increasing VEGF secretion, and relatively reducing mRNA expression of inflammatory cytokines. Together, SF-based scaffolds contribute to the long-term survival of islets resulting in reversal of hyperglycemia to euglycemia (Hamilton et al. 2017; Mao et al. 2017). Therefore, islet transplantation with SF-based scaffolds will be an effective therapeutic option in treatment of Type 1 Diabetes and can be applied clinically. However, the site of implantation and other fabrication conditions needs further standardization.

12.3.2 Decellularized Pancreas as Native ECM Scaffold ECM-derived non-cellular biological scaffolds that contain various macromolecules and different fibrous proteins including collagens, fibronectin, and laminins play pivotal role in TE and regenerative medicine (Chen and Liu 2016; Sackett et al. 2018). The pancreas also can be decellularized retaining its native ECM with internal elastic basal lamina, the ducts with their basal membrane as well as the glycosaminoglycan and collagen structures along with its vasculature by perfusion using 0.5% sodium dodecyl sulfate (Guruswamy Damodaran and Vermette 2018) which can serve as a biological scaffold material. The process of decellularization can be achieved by (1) mechanical vigorous shaking in suspension, freeze-thaw cycles, hydrostatic pressure, vascular perfusion, disruption (ultrasonic/manual) (2) chemically through the use of detergents, solvents as well as acid, alkali, and ionic solutions, and (3) through biological approaches using enzymes (for example, benzonase endonuclease). Although these techniques can be appropriately used to achieve maximum removal of cell debris and to retain the ECM components with minimal damage, these methods still need further standardization (Fig. 12.4) (Gilpin and Yang 2017).

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Fig. 12.4 Preparation of decellularized pancreatic scaffold. (a) Separation and collection of intact pancreases from the adjacent tissue. (b, c) Perfusion of the pancreas by various methods using detergents and washings with phosphate buffered saline gradually removes blood and cellular material makes it transparent by retaining its intact shape and structure. (d) Infusion with trypan blue confirms the intact vasculature branches of the decellularized pancreas. (e, f) Scanning electron micrographs showing the microarchitecture of native and decellularized pancreatic scaffolds. Absence of any residual cells and preserved ECM microstructure was evident in decellularized pancreatic scaffolds. Reproduced and modified from Wan et al. (2017)

Present method of islet transplantation into the intraportal vein of liver has limitations such as bleeding, thrombosis along with anoikis, hypoxia, and inflammation mediated immune response which subsequently results to loss of function of the islets and graft failure. During transplantation, the non-toxic, biocompatible decellularized pancreas can be infused with isolated islets gradually and the injected islets were retained in the ductal system of decellularized pancreas and the transplanted beta cells were also found to be biologically active in their function. Further reports from different studies also confirm that most of the islets are mainly located and attached to the decellularized pancreatic duct. Presence of LN-1, FN, and collagen IV in the ductal ECM improves transplanted β-cells survival and islet morphogenesis. Furthermore, glucose-stimulated insulin secretion response of islets in the decellularized pancreas 48 hours post-transplantation was also phenomenal (Guruswamy Damodaran and Vermette 2018). In conclusion, decellularized pancreatic ECM bioscaffold could serve as a “carrier” to transfer islets during transplantation and also function as a 3D framework for pancreatic TE in the generation of bioartificial pancreas. Further, due to their minimal complications, decellularized pancreatic scaffolds could provide an innovative therapeutic opportunity in TE of bioartificial pancreas for transplantation and could be one step closer to in vivo

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applications in regenerative medicine (Guruswamy Damodaran and Vermette 2018; Hashemi et al. 2018).

12.3.3 Surface Engineering of the Pancreatic Islets Cell surface engineering with natural or synthetic molecules is a renowned strategy to provide required cell surface modifications to decrease the size of the islets, protect implanted cells from host immune system and also improve graft function by enhancing its survival post-transplantation. Islet-initiated instantaneous blood-mediated inflammatory reaction (IBMIR) results to the thrombin generation that activates endothelial cells, macrophage infiltration along with generation of pro-inflammatory cytokines resulting to the significant loss of islets due to apoptosis in the peri-transplant period (Contreras et al. 2004). Heparinization of the islet surface inhibits the onset of IBMIR and increases the islet viability (Cabric et al. 2007). Similarly, since thrombomodulin (TM) is known to exert its anti-thrombosis, anti-inflammatory effect by activated protein C-dependent (APC) mechanisms, a novel strategy to re-engineer the surface of islets with recombinant TM (rTM) was also developed. Surface modification of isolated islets through TE with immobilized rTM, recombinant azido-functionalized TM increases generation and activation of APC and islet survival rate (Stabler et al. 2007; Wilson et al. 2010). Islet surface modification with the help of ultrathin heparin-incorporated starPEG (Hep-PEG) nanocoating was also reported to inhibit IBMIR. Further, due to the binding affinity of heparin to various cytokines, plasma proteins and growth factors Hep-PEG functions as a staging to incorporate novel biologically active mediators for living cell surface alterations (Yang et al. 2018). Immobilization of urokinase, a serin protease that dissolves fibrin blood clots, on to the surface of islets along with heparin and TM also improved the islets survival following transplantation (Chen et al. 2011). Immobilization of the synthetic heparin-binding peptide amphiphile on the islet surface through electrostatic interactions was successfully shown to hold growth factors to increase angiogenic response and revascularization without disrupting viability and function of the islets (Chow et al. 2010). Furthermore, islet encapsulation with living cells is another strategy to protect islets, improve biocompatibility and inhibit IBMIR. Immobilization of human embryonic kidney cells with amphiphilic PEG-conjugated phospholid derivatives or polyDNA-PEG-lipid conjugates onto the islet surface retained the insulin secretion ability (Teramura et al. 2010). Similarly, immobilization of testicular Sertoli cells also known to protect islet grafts from immune destruction due to their immunomodulating activity (Mital et al. 2010; Zhu et al. 2015b). Further, chondroitin sulfate-incorporated starPEG nanocoating (CS-PEG) approach has been developed for surface engineering of the islets where CS-PEG conjugates to the amino groups of the islet cell surface without altering islet volume. CS-PEG surface engineered islets mitigate IBMIR and show enhanced viability, revascularization, unaltered insulin secretion and also are believed to better the therapeutic potent of the islet transplantation (Yang et al. 2019).

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Cell surface engineering through layer-by-layer (LbL) deposition of polymers to modify islet surfaces with preferred permeability, thickness, mechanical stability, and surface chemistry is a widely used surface alteration method to prepare molecularly even ultrathin polyelectrolyte multi-layer films as well as nanothin PEG coatings comprising diverse biomolecules. These tailored surface modifications offer diverse multipurpose scaffolds for the confined arrangement and release of various bioactive agents that are aimed to improve islet function, viability, and engraftment (Wilson et al. 2011; Fakhrullin et al. 2012; Mets et al. 2013).

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Future Perspectives

In the recent years, 3D bioprinting has acquired great attention from the researchers in the TE domain for its ease of fabrication process, automation and its adaptability in creating natural environment for the cells in the culture. It is similar to the 3D printing process used to make various objects LbL but here only material required to print the biological tissues/organs is used and must be maintained at high sterility as the printed material is used in studies pertaining to regenerating damaged tissues. It is divided into three steps. During the first pre-bioprinting stage, the tissue is scanned for its 3D image using MRI or CT scan and this image is used in the next stage as a template to generate a 3D structure of the tissue or organ. The 3D image generated can be edited using software (for example, AutoCAD) to customize according to length, width, and depth based on the architecture and location of the damaged tissue/organ. In the second bioprinting stage, the printer is loaded with bio-ink. Bio-ink is a mixture of natural/synthetic polymer or a combination gel with/without conductive materials (for nervous system defects) incorporated with cells (specific to the tissue). Using the 3D image generated as a model, the tissue or organ is printed layer-by-layer. In the last post-printing step, printed tissue is allowed to cross-link and stabilize using physical (light, heat etc)/chemical agents. The generated scaffold is incubated further for the cells to grow, multiply, and differentiate. The cells growing are monitored at regular intervals for their function and are evaluated by comparing with the wild type cells in the corresponding tissue. If the scaffold has succeeded in mimicking the natural environment and supported the survival of cells along with retaining their function, it can be advanced to use in vivo to repair the damaged tissue. The generated scaffolds also can be used as disease models for various kinds of pharmacological studies. Many fabrication procedures are employed to construct 3D biomimetic scaffolds used in nerve and pancreatic TE including droplet microencapsulation, lithography, electro spinning, 3D printing, freeze-drying, bioprinting, decellularization, solvent casting, gas-foaming, cryogelation, porogen-leaching, rapid prototyping, particulate leaching, selfassembly, and phase separation. Based on the tissue specificity and nanostructured scaffold requirements suitable combinations or individual scaffolding approaches will be implemented. These fabrication procedures need sophisticated instruments and provide exceptional control over scaffold design, construction, and porosity (Lu et al. 2013; Kumar et al. 2018; Lee et al. 2018b; Salg et al. 2019).

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In neural TE, 3D printed conductive CNT scaffolds, hydrogels, NGCs, and neural mini-tissues (Lee et al. 2018a; Vijayavenkataraman et al., 2018; Vijayavenkataraman et al., 2019; Gu et al., 2016; Qian et al., 2018; Heo et al., 2019) have been reported to enhance nerve regeneration. Similarly, 3D bioprinting technology also has potential to develop artificial pancreas organ (Lee et al. 2018a). A lot of research has been going on in producing actual pancreas-mimic artificial organ for clinical applications in which 3D bioprinting of islets for generation of an artificial pancreas is gaining instigation (Ravnic et al. 2017; Kim et al. 2019a). Combination of islet encapsulation in alginate-methylcellulose hydrogel blends with 3D extrusion bioprinting allows the assembly of 3D structures with accurate arrangements to produce macroporous hydrogel constructs with embedded islets. In this study, 3D printed encapsulated islets showed insulin and glucagon production and also reacted to glucose stimulation (Duin et al. 2019). Further, ECM-based bio-inks play a critical role by providing suitable tissue microenvironment along with functioning as a native substrate during 3D cell printing. Kim et al. (2019b) showed that human iPSC-derived insulin-producing cells show enhanced insulin secretion, maturation in the presence of pancreatic tissue-derived ECM bio-ink which has potential applications in fabrication of 3D pancreatic tissue constructs (Kim et al. 2019b) whereas fibrin-based specialized bio-ink enhance iPSC-derived neural tissue 3D printing (Abelseth et al. 2019). Taken together, tissue modelling combining the stem cell technology, biomaterials, and 3D bioprinting can help in solving the existing encounters towards generation of 3D artificial organs for regenerative and therapeutic applications (Lee et al. 2018b; Kim et al. 2020). Acknowledgements The authors thank the Hyderabad Eye Research Foundation (HERF) for supporting this work.

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Biomaterials and Stem Cells in Tissue Engineering and Regenerative Medicine: Concepts, Methods, and Applications

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Vasagiri Nagarjuna

Abstract

Regenerative medicine (RM) and tissue engineering (TE) are the two novel approaches that mankind can bank upon to explore new avenues in understanding the complexities of living systems and to lead the future bio-medical research. Nevertheless, the concepts of TE and RM are based on growth, regeneration, proliferation, and differentiation of cells. As such, stem cells make the core with properties to perform all the functionalities conceptualizing TE and RM. Stem cells along with biomaterials form an integral part of these approaches which underlie the healthcare research and its eventual ensuing advances. Properties, structure, composition, and functionality of biomaterials used play a key role in development of cells and tissues in TE and RM. Since, supporting material is vital for cells for regeneration/repair of tissues or in organ development, different biomaterials as scaffolds are used for the purpose. Understanding the intricate developmental process of cells to tissues to organs is vital in TE and RM. Use of different biomaterials along with various lineages of stem cells supplemented with nutrients, growth factors, etc. will help to decipher the underlying entangled mechanisms. A good understanding of the basic and advanced concepts of biomaterials and stem cells, the different methods involved, and their applications in TE and RM are of at most prerequisite to further the goal of developing functional tissues and organs. Hence, here below in this manuscript, we provided detailed insights of the concepts, methods, and applications of biomaterials and stem cells in TE and RM. Keywords

Stem cells · iPSC · MSC · ESC · TE · RM · Biomaterials V. Nagarjuna (*) Society for Biological Chemists India, Bangalore, India # The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 B. Bhaskar et al. (eds.), Biomaterials in Tissue Engineering and Regenerative Medicine, https://doi.org/10.1007/978-981-16-0002-9_13

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Abbreviations ESC iPSC MSC PCL PLGA PLLA

Embryonic stem cells Induced pluripotent stem cells Mesenchymal stem cells Poly(ε-caprolactone) Poly(lactic-co-glycolic acid) Poly(L-lactic acid)

13.1

Introduction

Regeneration of lost/dysfunctional tissues/organs has been a topic of much interest and has been discussed even in ancient mythology. Archeological research has presented great evidences of transplantation/replacement of teeth, etc., during ancient civilizations like Roman, Egyptian, Indian, South American, etc. Later, the concept of regeneration was presented by Aristotle in his work on natural history. This was furthered by the evidence of regenerative capacity of freshwater polyps by Trembly. Research leading to deeper insights into cellular functioning and with the establishment of cell therapy, the fascinating studies on regenerative medicine took a new turn and thus started the journey to the modern era of RM and TE. Introduction of stem cells for use in TE and RM has led to several breakthroughs in cell/tissue regeneration. Use of different stem cells like MSCs, ESC, and iPSCs in tissue regeneration has been reported in various studies. However, problems such as immuno-compatibility, low differentiation potential, poor survival, and risks pertaining to the formation of teratoma came in the way of using stem cells alone for clinical purpose (Drukker 2008). Moreover, for use in TE, the need for proper surface which mimics the natural microenvironment is of at most importance for tissue regeneration either in vivo or in vitro. The use of biomaterials as matrix and scaffold materials proved to be vital in providing this environment. Advances in biomaterials helped in fabricating materials which mimicked the natural ECM, provided the structural architecture and porosity for cell adhesion, migration, and regeneration in the process of tissue formation, thus paving the way for use of stem cells in therapeutic approaches (Burdick and Vunjak-Novakovic 2009).

13.1.1 Biomaterials Biomaterials are the materials which are biocompatible and are used for various biological applications to interact with the cellular systems without causing any adverse effects. Biomaterials possess various properties such as biocompatibility, biodegradability, bioresorbability, etc. or engineered to have those properties for their use in biological systems. Biomaterials may be natural or synthetic.

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13.1.2 Stem Cells Stems cells are broadly defined as the cells which possess the capacity of regeneration and to differentiate into other progenitor cells. These cells can be totipotent, pluripotent, multipotent, or unipotent depending on their capacity to regenerate to different cell types. Depending on the derivation of cells from developmental stages these are classified mainly into ESCs, iPSCs, and adult stem cells. ESCs are pluripotent stem cells derived from embryonic blastocysts inner cell mass. Though they are pluripotent and can be differentiated into several cell lineages, their use is mostly restricted due to ethical issues (Vazin and Freed 2010). Also they may be subjected to immune rejection due to their allogenicity with the patient and 100% differentiation of these may not be possible leading to retaining of a small fraction of undifferentiated cell which can form teratomas (Rong et al. 2014). As such is the case with ESCs, Yamanaka et al. from RIKENs institute, Japan introduced the method of reprogramming somatic cells into primordial cell state known as induced pluripotent stem cells (iPSCs). iPSCs have been shown to possess the properties which successfully overcome the drawbacks faced by ESCs (Li et al. 2017). The only risk with iPSCs is the chance of developing neoplastic cells from the differentiated cells derived from iPSCs since the development of tumors is associated with reprogramming factors. Adult stem cells/Mesenchymal stem cells (MSCs) address most of the problems faced by the other two and are being increasingly used for clinical applications nowadays. MSCs are the stem cells isolated from adult tissues such as bone marrow, tonsils, adipose tissue, etc. They express various cell surface markers such as CD90, 105, etc. and can be differentiated into various mesenchymal lineages such as muscle cells, osteoblasts, etc. in addition to their self-renewal capacity. As such as they are isolated from the patient they do not have immune-rejection problem. They possess immuno- modulatory properties, and pro-angiogenic properties which make them apt for use in various clinical applications (Murphy et al. 2013).

13.1.3 Concept of Stem Cell The concept of stem cells derives from the fact of their differentiation and regeneration capacity. Stem cells possess the unique properties of self-renewal, proliferate and differentiate into other cell types. Depending on their ability to differentiate they may be classified as follows. Totipotent stem cells: These are the most potent stem cells which have the capability to differentiate into any cell type giving rise to all the three germinal layers and to organism as a whole. Further these can give rise to extra embryonic tissues such as placenta. E.g., Zygote, Initial blastomeres. Pluripotent stem cells: These have the capacity to regenerate into all the three germinal layers, i.e., ectoderm, mesoderm, and endoderm but cannot give rise to placental tissue. E.g., Inner cell mass of blastocyst, cells from morula stage of embryo etc.

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Multipotent stem cells: These are the cells which have the capability to give rise to cells of their germ line origin only. E.g., MSC. Unipotent stem cells: These are also known as progenitor cells which are capable of giving rise to cells of specific lineage. E.g., Cardiac stem cells, neural progenitor cells, etc.

13.1.4 Different Types of Stem Cells Embryonic stem cells (ESCs): These are derived from the inner cell mass of blastocyst and have the capability to differentiate into any cell type. Mesenchymal stem cells (MSCs): These are derived from different sources such as bone marrow, peripheral blood, adipose tissue, etc. and they have the capacity to differentiate into respective tissues. Induced pluripotent stem cells (iPSCs): These are the adult cells which are reprogrammed into pluripotent cells and they can differentiate to any cell type of the organism form which they are derived. Adult stem cells (ASC): These are the stem cells derived from adult tissues such as nerve, bone, blood, etc. and they have the capacity to differentiate into the particular cell type from which they are derived. Fetal stem cells: These are the stem cells derived from the tissues of fetus and they can differentiate to the tissues they are derived of. Uses: Stem cells are used in a wide range of diversified applications in biology and medicine. They are used in various therapeutic and clinical applications. In the present context of this chapter their use in TE and RM has been discussed below in detail.

13.1.5 Tissue Engineering and Regenerative Medicine TE and RM are the two approaches that have shown promising results in human health care. Regenerative medicine uses the concepts of multidisciplinary areas such as biology, material and engineering sciences to regenerate/restore/replace damaged or diseased tissue. TE is the approach of regenerating/engineering tissues/organs for replacement of damage ones. These have revolutionized medicine by their applications in varied spectrum of treatments where conventional treatment has no cues. Serious research in TE and RM started during the past few decades and it has taken an enormous leap with the introduction of biomaterials and stem cells for the generation of various tissues and organs or in the generation of different cell types and their delivery at the targeted site.

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13.2

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Biomaterials and Stem Cells in TE and RM

Advances in the development of biomaterials in the past few decades have breathed new life in the establishment of TE and RM as promising areas wherein cues for various clinical and therapeutic problems which were beyond the scope of understanding and curing in the past were unfolded. The fundamental aspect that needs to be emphasized in the application of biomaterials in TE and RM is that to mimic the natural microenvironment of the targeted cells, tissues/organs. Development and differentiation of stem cell need specific micro-environmental conditions. Different properties of biomaterials like mechanical, chemical, surface morphology need to be considered for use in stem cell cultures since they regulate the cell matrix interactions and cell adhesion properties which in turn regulates the differentiation lineage specificity of stem cells. Cell signaling, adhesion, migration, differentiation, and matrix organization can be modulated by the use of different biomaterials which directs the differentiation process of stem cells. Therefore use of appropriate biomaterials and fabrication techniques in the preparation of scaffolds is of utmost importance in TE and RM. Both natural and synthetic biomaterials are used for the purpose of stem cell culture in TE and RM. The natural biomaterials used for the purpose include collagen, gelatin, fibrin, hyaluronic acid hydrogels, hydroxyapatite, etc. Use of natural biomaterials has the advantage of mimicking the natural ECM in adhesion, biodegradability, and natural biosignaling, but fast degradation rates and complications in purification make to their disadvantage. Moreover the difficulty in modification of these natural biomaterials along with their poor mechanical strength when compared to synthetic biomaterials adds to the problem. Different synthetic biomaterials like synthetic polymers, e.g., PLA, PEG, PLGA, PHEMA, synthetic ceramics, metals are used for the purpose of stem cell cultures in TE. Synthetic biomaterials have their advantage in that the structure and mechanical properties along with the composition can be decided as per the need of the application. But more hurdles exist in use of synthetic biomaterials in the form of biocompatibility, cell adhesive properties, biosignaling, etc. Of late, to overcome these problems, composite biomaterials are being used in the fabrication of scaffolds for the purpose of stem cell cultures in TE and RM. A brief detailing of different biomaterials being used for culturing stem cells in TE and RM is been discussed below. Collagen which is the main component of ECM possesses higher biocompatibility and does not cause adverse immune effects. Due to its higher biodegradability, it is mostly used in skin grafts (Trottier et al. 2008). Also it is being used in cartilage regeneration and corneal TE. Gelatin, a derivative of collagen is being used in cartilage TE due to its biodegradability, biocompatibility, and its property to form hydrogels (Schuurman et al. 2013). Stem cells encapsulation by hydrogels formed by gelatin activated with unsaturated methacrylamide was found to be better in performance than hydrogels formed from other biomaterials (Van Den Bulcke et al. 2000). Matrigel forms an active 3D matrix by polymerizing at room temperature, which mimics the natural in vivo conditions and helps in the differentiation process

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of stem cells (Rowland et al. 2010). Hydrogels of hyaluronic acid upon photo cross linking are used for the purpose of encapsulating MSCs with great survival rates in chondrogenesis (Chung and Burdick 2009). Fibrin is used for the purpose of nerve TE from stem cells (Willerth et al. 2007). Synthetic polymers such as PLA when surface coated with polydopamine have been shown to regulate the adhesion, proliferation, and differentiation of hAD-MSCs (Kao et al. 2015). PEG in the form of gel was used in human MSC differentiation for osteogenesis. It provides enough extra cellular space for diffusion of wastes and nutrients between stem cells and ECM (Nuttelman et al. 2004). Synthetic ceramics are the composites of hydroxyapatite with calcium phosphate, bioglass, etc. These composites optimize the differentiation of osteocytes from stem cells in bone and dental TE (Hutmacher et al. 2007). Metals such as titanium, stainless steel, cobalt alloys are widely used in TE applications. Titanium has been shown to enhance the multi-lineage differentiation capacity of MSCs in vitro. Many studies have shown the biocompatibility, enhanced efficiency, etc. of titanium and tantalum alloys in the culture of MSC for bone TE making them suitable in the preparation of biomaterial scaffolds (Hee et al. 2018). Graphene is another biomaterial which has been studied for its use in stem cell cultures since it can act as biosensors for Nanog protein which helps in quantification of pluripotency of stem cells (Chikkaveeraiah et al. 2012). Nevertheless, the above examples are glimpses of the use of different biomaterials for culturing stem cells for use in TE and RM. Different applications of stem cells in TE and RM and the biomaterials used for the purpose have been detailed further in this chapter.

13.3

Applications of Biomaterials and Stem Cells in TE and RM

Present day concept of TE cannot be envisioned without the application of stem cells. TE and RM relay on the principle of replacement/regeneration of cells/tissues which are damaged or lost. Since stem cells possess the properties of regeneration and differentiation, the vital aspects in TE and RM, use are of stem cells in various TE related applications will not only provide deep insights into the developmental aspects but also results in promising outcome. Also, the use of appropriate biomaterials is of utmost importance for this purpose. The detailing of different biomaterials and their use in TE and RM in the context of stem cells has been briefed above. In this section we try to bring in the various applications of biomaterials with respect to stem cells for use in TE and RM. Our main focus will be on discussing how different biomaterials play their role in directing the differentiation of stem cells into different lineages and their development in TE and RM.

13.3.1 Stem Cells and Biomaterials in Bone Tissue Engineering One of the major applications of TE is regeneration of bone tissue. Different stem cells such as ESCs, MSCs, iPSCs, and ADSC are cultivated on 3D scaffolds which are osteoconductive to evaluate their osteogenic potential (Marolt et al. 2012).

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Results showed that iPSCs and ADSC are more promising in bone regeneration with better angiogenic, osteogenic, and differentiation capabilities (Bastami et al. 2017). Also, ADSCs have shown good results in pre-clinical allograft studies (Dufrane 2017). In other studies on hESC cultured on 3D nano fibrous scaffolds, it has been shown that there is enhanced osteogenic differentiation (Smith et al. 2009). Another promising source for bone regeneration is that of human embryonic stem cellderived mesenchymal stem cells (hESC-MSC) which when grown on NF scaffolds resulted in increased chondrogenic and osteogenic differentiation (Hwang et al. 2006). Biomaterials used for the purpose of bone TE need to have differential pore morphology to effectively support osteogenesis. It was shown that bimodal porous PCL scaffolds characterized by macropores resulted in enhanced colonization of MG63 osteoblast cells (Salerno et al. 2011). In particular, growth of chondrocytes and osteoblasts was favored with pore size ranging from 380 to 450 nm and bone differentiation was accelerated at pore sizes of 290–310 nm. Different natural and synthetic biomaterials were used for the purpose of bone regeneration. Enhanced efficiency was achieved by the fabrication of composite biomaterials which possess the properties of osteoconductivity, osteoinductivity, and mechanical stability (Motamedian et al. 2015). It was shown that starch/PCL scaffolds with an osteogenic layer improved osteoarthritic treatment (Rodrigues et al. 2012). PCL has been shown to have wider acceptance of the biomaterial of choice in bone regeneration due to its biocompatibility, non-toxic, and non-inflammatory products and its degradation rate, which matches the regrowth of natural bone tissue.

13.3.2 Stem Cells and Biomaterials in Cardiovascular TE and RM Cardiovascular tissue engineering is one of the most promising and sought after applications of TE since one of the leading causes of mortality worldwide is due to cardiovascular diseases. Various stem cells that are used for the purpose include iPSCs, parthenogenetic stem cells, and ESCs. To achieve optimal efficiency in integration and survival of cardiac stem cells, use of biomaterials was necessitated. Various biomaterials like collagen, alginate, peptide amphiphile nanofibers, PEG, hydrogels, chitosan, fibrin, etc. were used for the purpose of heart TE (Hirt et al. 2014). Studies showed that biomaterials such as alginate, methyl cellulose, hyaluronic acid, etc., devoid of cells, when applied, resulted in restoration of heart (Tang et al. 2017). It was also shown that collagen when incorporated with follistatin-like 1 after MI, stimulated cardiomyogenesis in vivo (Wei et al. 2015). Due to the complexity in structure and regeneration of heart tissue, stem cell based approaches had gained prominence over other methods. Human pluripotent stem cell-derived cardiomyocytes were used in combination with different biomaterials for this purpose. Biomaterials such as hydrogels were used in the repair of cardiac tissue. Encapsulation of CMs for enhanced survival, differentiation, and attachment can be achieved by using alginate hydrogels (Yu et al. 2010). Use of 3D PEG hydrogels with thiosulfate cyanide sulfur transferase (TST) was shown to minimize

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the ischemic effects and damage due to reperfusion and stimulated angiogenesis at the implanted site (Mauretti et al. 2016). Further, nanofiber based hydrogels were shown to possess enhanced elasticity and induce the differentiation of MSC into smooth muscle cells (Wingate et al. 2012).

13.3.3 Stem Cells and Biomaterials in Pancreatic Tissue Engineering Damage or loss of function of pancreatic β cells leads to type 1 diabetes mellitus, characterized by depleted levels of production of insulin. Pancreatic β cells derived from stem cells are a major source in treatment of the disease. iPSCs derived from the patients are used for the purpose. For the generation of tissue porous scaffold material which mimics natural microenvironment is needed. Encapsulation is one of the major methods applied in pancreatic TE wherein different biomaterials such as alginate, agarose, PEG, polyacrylates, etc. are used for coating and to produce micro/ nano devices in encapsulation. It was shown that PLGA NP scaffold in the presence of 2-(4-morpholinyl)-8-phenyl- 4H-1-benzopyran-4-one induces endoderm formation prior to pancreatic differentiation from iPSCs (Kuo et al. 2017). Another study has shown the increased efficiency in insulin production and enhanced survival of islets upon use of PLLA/PLGA scaffold seeded with human fibroblast cells, HUVECs, and pancreatic islets (Kaufman-Francis et al. 2012).

13.3.4 Stem Cells and Biomaterials in Nerve TE Nervous system is one of the most complicated systems of human body. Various neurological diseases such as Alzheimer’s disease, PD, etc. have severe impact on human health. Regeneration/repair of effected/damaged nervous tissue will address the problems posed by these disease conditions. Several biomaterials were used for the purpose of nerve TE. Scaffold compatibility is one which should be taken care of while using stem cells for differentiation into glial and neural cells. It was shown that nerve progenitor cells seeded on collagen 3D matrix are used in the regeneration of central nervous system (Winter et al. 2016). Studies have shown the enhanced efficiency in differentiation of MSCs into neurons upon use of Poly (L-lactic acid)-co-poly-(3 caprolactone) scaffold biomaterials (Prabhakaran et al. 2009). It was also reported that increased rate of proliferation was found to be achieved with the use of PLGA scaffolds seeded with KT98 stem cells (Peng et al. 2017). Further use of PC12 stem cell type with PEDOT-HA scaffolds results in enhancement of growth of synapse (Wang et al. 2017). Several other applications of biomaterials and stem cells in TE of various tissues have been reported in literature. Since detailed description of each of the applications is beyond the scope of this chapter we have tried to put in some of the most important applications above.

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13.3.5 3D Bioprinting and Stem Cells in TE Use of 3D bioprinting technology has been the latest and most sought after process in the regeneration of tissues nowadays. Since most of the tissues are made of multiple cells and layers, the need for the development of multi-layered 3D scaffolds for constructing tissues using stem cells became necessary. Advances in this technology had made it possible to place different stem cells in multilayered fashion on the scaffolds for the regeneration of tissues (Mandrycky et al. 2016). Inkjet and laser deposition 3D bioprinting technologies are majorly used for the purpose. When compared to inkjet, laser printing has the advantage that they do not require additives to maintain liquid form and do not include physical stress in the process as do inkjets have. Also, cell viability studies have shown more than 95% survival rate without effecting either cell proliferation or apoptosis (Koch et al. 2013). Different biomaterials both natural such as gelatin, fibrin, alginate, collagen, etc. and synthetic (PEG, Pluronic gels) are used as bioinks for the development of 3D scaffold material (Zorlutuna et al. 2013). Since these should mimic natural ECM, for in vivo cultures use of decellularized ECM for use with stem cells for 3D bioprinting has been developed (Chen et al. 2018). Another factor that plays a vital role in the development of tissues is in the proper supply of blood and nutrients, i.e., development of blood vessels. To achieve this, addition of pro-angiogenic factors during 3D bioprinting can be considered as the best way possible in development of 3D bioprinted tissues (Jeon et al. 2018). One of the applications of 3D bioprinting in TE using stem cells is in the development of skin grafts. Studies have shown that use of skin derived dECM instead of collagen for bioink preparation is more advantageous since it naturalizes the degradation and concentration quotients when compared to collagen. Use of AD-MSCs and endothelial progenitor cells with dECM as bioink for the preparation of pre-vascularized skin grafts was shown to have accelerated the process of wound healing (Kim et al. 2018). Another application is in bone regeneration wherein 3D printed calcium phosphate scaffolds have been developed along with the use of polydopamine for increasing the osteogenic capacity of stem cells (Inzana et al. 2014; Teixeira et al. 2019). Also, use of nano fibrillated cellulose and alginate has been reported in the development of cartilage tissue for fabrication of 3D bioprinted ear (Markstedt et al. 2015). Other applications include use in development of various tissues including vascular grafts for cardio vascular defects, liver tissue, nerve tissues, skeletal tissue, etc. More and more focus is being projected in this area of research as this is making promising inroads into the development of functionalized tissue, thus enhancing the future prospects of TE.

13.4

Conclusion

With the advent of stem cell research and their applications, major breakthroughs in science have been achieved. Some of the most benefiting aspects of stem cells came along with their introduction in TE and RM. Different stem cells such as pluripotent,

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multipotent, etc. have different applications in their use in TE. Since selection of appropriate biomaterial scaffold materials is one of the prerequisite along with the stem cell type seeded for the regeneration of targeted tissue, research to this end in the development of advanced biomaterials has opened up new vistas in the fields of TE and RM. The properties of biomaterials along with their spatial structural arrangement provided the necessary microenvironment for the adhesion, proliferation, and differentiation of stem cells. Differentiation of stem cells into different cell specific lineages can be achieved by using novel biomaterials. Further development of different fabrication techniques has provided an added edge in the development of various tissues such as nervous tissue, muscular tissue, bone, heart, etc. In together, stem cells along with biomaterials have paved the way for enormous scope of applications in TE and RM leading to transformation in human health care research.

13.5

Future Prospects

Future prospects lie in the development of novel composite biomaterials and their fabrication into scaffolds for use in specific applications for the differentiation of stem cells. Further development of 3D bioprinting technologies along with fabricating scaffolds mimicking the natural microenvironment can lead to great achievements in TE and RM.

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Part IV Advances in Biomaterials

Biomaterials in Tissue Engineering and Regenerative Medicine: In Vitro Disease Models and Advances in Gene-Based Therapies

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Swathi Dahariya and Vasagiri Nagarjuna

Abstract

The two interrelated and interdisciplinary areas of research which offer promising health care solutions and have the potential to lead the future are tissue engineering (TE) and regenerative medicine (RM). Biomaterials form an indispensable part of TE and RM and recent innovations in biomaterials development have laid new vistas in TE and RM research. Advances in biological/medical research have its strong foundations in the evaluation of in vitro disease models and gene-based therapies have become the cornerstone for disease treatments in modern era. As this is the scenario, a deep insight into how in vitro model systems and genebased therapies made their way in the advancement of TE and RM together with their applications makes an interesting point of study along with the biomaterials being used for the purpose. Here in, we have discussed in detail the different in vitro disease models, their importance and use with respect to TE and RM. Also, various gene-based therapies and their application in TE & RM were also discussed along with the different biomaterials used for the purpose. This will present a complete picture of how research in TE and RM is reshaping with the inclusion of gene-based therapies and in vitro disease models leading the way of future human health care treatments. Keywords

Gene therapy · Stem cells · Invitro disease models · Adeno-associated viruses · Biomaterials S. Dahariya Department of Biochemistry, School of Life Sciences, University of Hyderabad, Hyderabad, Telangana, India V. Nagarjuna (*) Society for Biological Chemists India, Bangalore, India # The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 B. Bhaskar et al. (eds.), Biomaterials in Tissue Engineering and Regenerative Medicine, https://doi.org/10.1007/978-981-16-0002-9_14

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Introduction

TE is the amalgamation of cell therapy, biomaterials, biomedical engineering and genetics all with the aim “to stimulate regeneration of tissues and organs by either implanting biomaterials for in vivo regeneration or by constructing substitutes in vitro”. Mason and Dunnill (2008) described regenerative medicine as a therapy accorded by replacing or regenerating human cells, tissues or organs to restore and establish their normal function. For example, Professor Thomas Brown reported through his heart and lung research in Knut that Knut heart can recover during heart attack. He used the unique property of Knut that can regenerate various types of tissues very efficiently and tried to apply similar mechanism in human to repair damaged heart. Heart attack is a condition in which a part of heart is not properly supplied with blood, resulting in depreciation in supply of oxygen to heart muscles cells and formation of scar tissue which cannot contract in the same way as a normal muscle of heart in human. But he noticed in Knut the scar tissue turns back into functioning muscle cells and beats at full strength again after few days of heart attack. Professor Thomas Brown and his team observed rejuvenation process carried out by group of baby cells or younger cells. Similarly, in mice they conducted the rejuvenation process by inserting embryonic cells to capitulate heart cells to dedifferentiate and observed that heart cells transformed to precursor cells once more leading to proper functioning of heart. In this way gene therapy applied in combination of stem-cell-based tissue engineering was found to have enormous potential in stimulating the regeneration of heart cells in the ambience of optimal expression of regulatory proteins and would pave the way to treat various cardiovascular diseases (Goker et al. 2019). The coordinate interaction of cells with soluble factors and extra cellular matrix (ECM) determines the exact proliferation, differentiation migration, and maturation process of tissue cells. The simple principle underlying the basics for TE/RM is to mimic the natural process of healing in the reconstruction or regeneration of damaged tissue. Tissues or organs that are acutely diseased or lost either by congenital anomaly or trauma and which cannot be treated conventionally, restoration of damaged/lost tissues or organs can be achieved through transplantation. The main drawbacks in surgical treatment for organ transplantation include shortage in donor tissues/organs and immunocompatibility/rejection. Although immunosuppressive therapies have recently been much advanced, a number of challenges are yet to be addressed. Nearly three decades back, with the advent of “Tissue engineering” (TE), a new paradigm in the field of biomedical sciences ushered a new era to modern biotechnology and health science research. In the year 1993 Langer and Vacanti, known to be founders of TE, pioneered the concept of TE. According to them TE is a highly multidisciplinary area of research which focuses on developing alternative therapies for tissue/organ repair and replacement by applying the principles of engineering technology and life sciences in developing therapeutic strategies with high success rates. The concept of TE is not completely new by itself, since the ancient Rasayana concept of Ayurvedic physiology explains its theory and the current understanding of tissue engineering as an alternative approach to develop new cell/tissue and

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artificial organs emerges from it. Development of biomaterials is one of the corner stone for success in TE since the development of functional tissues/organs is due to combination of cells scaffolds and biologically active molecules wherein scaffolds form the backbone to achieve the desired outcome. Tissue engineering, the outcome of interface of biology with other sciences and modern technology, aims at achieving major medical breakthroughs through development of regeneration platforms for the treatment of diseases and trauma by providing functional tissues/organs as in vivo transplants to ease the problem of organ shortages or as in vitro models to study disease mechanisms and in drug discovery. The field of tissue engineering relays on four basic fundamental parameters: selection of right cells for required outcome, providing the right environment such as scaffold to support the cells; supplementing requisite molecules like growth factors, cytokines, etc. to make those cells healthy and productive; and physical and mechanical factors to influence the development of the cells (Heyde et al. 2006; Eggert and Hutmacher 2019). The cells can be harvested directly from the target organ or can be developed from precursor/stem cells or else taken from cell lines cultivated in the labs, ideally isolated from the patient, as that limits problems with rejection. The supporting structures can be obtained either from donor tissue or from natural and synthetic polymers or biomaterials to provide them with strength and endurance. The required biomolecules can be supplemented either directly or coaxed from the cells that are embedded along with the scaffold. Depending upon the materials used, scaffolds may dissolve overtime but the remaining components provide support to the organ. A few examples of successfully tissue engineered tissues and organs that have already been implanted into humans are cartilage, bladders, small arteries, trachea, and skin grafts (Gagandeep et al. 1995; Gonçalves 2017; Kebriaei et al. 2017). Biomaterials form an integral part of tissue engineering. They are designed to provide the architectural framework which can mimic the native extracellular matrix in order to support the cell growth and eventual regeneration of tissue. Biomaterial development is an important aspect of tissue engineering that helps not only in providing the structural scaffolds necessary for the development of tissue but also promotes regenerative process by proper transportation of therapeutic agents and cells (Lee et al. 2014). Improvement in the quality of tissue engineering constructs has been achieved over the years due to advances in development, design, and fabrication techniques of biomaterial scaffolds. These scaffolds provide the microenvironment for cells that are seeded to adhere, proliferate, differentiate, and to eventually form into tissue. An ideal scaffold is one that possesses biomimetic properties and has good physicochemical properties which aids in tissue regeneration (Kazemnejad et al. 2016). Various types of scaffolds have been developed either from synthetic (polymers and composites) or naturally (polysaccharides and proteins) biomaterials. Natural biomaterials possess the properties to mimic native tissue, conferring biocompatibility and can be recognized by the body (Gonçalves 2017). The porous nature of scaffold allows for high mass transfer and aids in waste removal (Lodish et al. 2000). Natural polymers have wide range of material properties similar to the native ECM and also possess surface modifying properties

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to modulate themselves according to the need. Some examples of natural polymeric biomaterials are collagen, gelatine, elastin and silk which are protein-based biomaterials and alginate, hyaluronic acid, chitosan and cellulose which are polysaccharide-based biomaterials. Cellulose has been in focus mainly because of its biocompatibility and relative readiness for chemical modifications along with adjustments in mechanical properties which drives at various surface modifications in various TE applications (Modulevsky et al. 2014, 2016). A number of cellulose derivatives have been developed including porous hydrogels, membranes, and porous scaffolds (Gershlak et al. 2017). Gene therapy can be defined as direct delivery of genetic material (nucleic acid) into patient’s cells for the replacement of abnormal genes or to make a beneficial protein to treat disease. This has become a promising treatment approach in medical field designed to treat a number of diseases (e.g. inherited disorders, some types of cancers and viral infections and diseases that have no other cures). The aim of gene therapy is to focus on restoring the normal function of the gene/protein, against abnormal gene or mutated gene which results in faulty coding a necessary protein or missing of protein in disease condition by introducing a normal copy of the gene. Gene therapy is used in tissue engineering and regenerative medicine for disease treatment, to create functional human tissues for enhanced intrinsic production of hormones, human growth factors, monoclonal antibodies, anti-haemophilic factors and many other proteins from the target cells of the tissue. For over a decade, the in vitro approach of introducing a gene into a cell line with the help of postie and evaluating the gene for its expression and functional activity has been in studies for the evaluation of small molecule drug–drug interaction properties. The hybrid approach of using gene therapy with tissue engineering has opened up an new era with excited applications in medical field in not only treating diseases but also leading to ultimate cure and holds a great promise in repairing, replacing tissues and organs of damaged tissues/organs and also regenerating new tissues/organs that fail due to congenital abnormalities, genetic errors, disease or traumatic injuries (Eugene 2016).

14.2

In Vitro Disease Models

Any tissue engineering/regenerative medicine process requires specific cells of interest which make up the principal engineering materials of the process. Since cellular systems are extremely complex functional systems with intricate structure and functions, it is a quite complicated task to explore and to understand the interactions between individual components and to decipher their basic biological functions. Studies performed outside the living organisms to elucidate the cellular systems are termed as “in vitro models”. In vitro models are very important in medicine and biology, because they simplify the study of living systems and provide deep insights into cells’ structural, functional as well as behaviour aspects. In vitro model not only simplifies the living system to study but also complement animal models by offering a high degree of controllability, repeatability and reproducibility

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(Healy 2018). Since these studies are set outside the natural environment of the cells, these models cannot precise in predicting the effects on the entire organism. To overcome this, studies are performed inside the living system using the whole organism as an experimental model termed as “In vivo model”. In vivo models enable a more convenient species-specific, simpler and more detailed evaluation of the system. However, corresponding in vitro model systems are necessary to understand the mode of pathogenesis at the cellular level and to fast-track the discovery of therapeutic compounds. At present, for the study of many human diseases and to develop therapeutic approaches in medical research dependence on appropriate model systems is the main stay. An in vitro disease model can be established ideally from human diseased tissue/cells or disease induced cells which exhibits relevant disease/degenerative mechanisms. Over the last decade rapid advances in the field of stem cell biology paved the way for the development of in vitro models to study the cellular functions and to explore mechanisms of cellular interactions and disease development for certain diseases which were never thought that they can be studied previously. There are quite few examples of in vitro engineered diseases models of the skin and vascular tissues, heart, liver, lung, intestine, liver, kidney, cartilage, musculoskeletal, endocrine and nervous systems, and also the models of infectious diseases and cancer have been detailed below.

14.2.1 Different In Vitro Disease Models Used in TE &RM 14.2.1.1 Primary Skin Fibroblasts as a Model of Parkinson’s Disease Parkinson’s disease is neurodegenerative disorder with second highest occurrence of cases observed worldwide. Studies have revealed that most of the cases occur due to sporadic mutations in number of genes which includes Parkin (PARK2) and PINK1 (PARK6) which have been associated with the disease. Different in vivo animal models and in vitro models such as skin fibroblasts from patients and recombinant cell lines are used as model systems in the study of Parkinson’s disease. Since PINK1 and Parkin genes are involved in stress response pathways and that human fibroblasts are shown to express relevant levels of PINK1 and Parkin genes, working on them would be an apt in vitro model for Parkinson’s disease. Also, using skin fibroblasts helps in presenting the system with defined mutations and in subsequent evaluation of the cumulative cellular damage caused (Auburger and Alexander 2011).

14.2.1.1.1 Advantage of Skin Fibroblasts as an In Vitro Model of PD Skin fibroblasts are readily available and are robust in nature which makes them the most favoured pick for use in in vitro model of PD. Also, they represent human primary cell model which ideally makes up to the patients biological aging and chronological order with respect to the patients polygenic predisposition and environmental etiopathology further adding to their advantage.

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14.2.2 In Vitro Model Study of Fibroblast Activation Using Hydrogel Scaffolds The mechanism of wound healing is a result of complex signalling in response to injury between cells and their environment. Fibroblasts present in the extracellular matrix (ECM) of different connective tissues are vital in the synthesis of matrix and repair mechanisms. Fibroblasts are activated when subjected to injury/chronic insult into the wound healing cells, myofibroblasts, which helps in repair of damaged tissue by secreting the necessary enzymes and proteins. When there is misregulation or persistence presence of myofibroblasts it can lead to tissue stiffening, uncontrolled accumulation of matrix proteins and ultimately disease. In order to study the mechanism of wound healing in vitro models are developed for evaluation. For this purpose, hydrogels which are water-swollen polymeric cross-linked networks are used. These can mimic the native ECM for use as in vitro models. In some hydrogel materials natural biological signals which implicit the native ECM are present while in other materials amalgamation is done to mimic the biochemical properties and mechanical properties of native ECM. Native ECM components which are naturally derived such as fibrin, collagen and hyaluronic acid and synthetic materials such as polyethylene glycol (PEG) and polyacrylamide which are functionalized with bioactive moieties such as modification with protein mimetic peptides or with whole ECM proteins, respectively, are used for the development of hydrogels to culture fibroblasts in vitro. Synthetic bio inert hydrogels are the preferred materials for hydrogel preparations in fibroblast culture in vitro models since studies can be performed by regulated addition of different biochemical ECM cues. This helps in evaluating the effects of different components in the proliferation, migration and activation of fibroblasts along with their cell adhesion properties both in 2D and 3D cultures in vitro (Smithmyer et al. 2014).

14.2.3 Induced Pluripotent Stem Cells as In Vitro Disease Models One of the most important and powerful approaches that holds great promise for the future of regenerative medicine was the adult somatic cells reprogramming into induced pluripotent stem cells (iPSCs). Takahashi and Yamanaka (2006) have successfully reprogrammed somatic cells into induced pluripotent cell state for the first time. Oct4, Sox2, c-Myc and Klf4, the four transcription factors are the corner stones for this pioneering work as they are found to be vital in the transformation of terminally differentiated cells into pluripotent state. The resulting iPSCs are found to be capable of generating the cells and tissues of all three germinal layers as well as whole organism, a property identical to the human embryonic stem cells (hESCs). Their work was further reconfirmed by using different donor cell types such as haematopoietic cells, neuronal cells, adipose stromal cells, skin cell, etc., by many researchers. Since then, iPSCs were used in various applications including their potential role in modelling diseases and drug screening. Use of iPSCs as in vitro disease models helps in overcoming various limitations faced by normal diseased

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cells. Thus they are proved to be versatile and provide a powerful tool basic research and in understanding the disease and its therapy.

14.2.4 Human Mesenchymal Stem Cells as In Vitro Disease Models Mesenchymal stem cells (MSCs) are the adult stem cells which are isolated either from human or animal sources. Human MSCs (hMSCs) are non-haematopoietic, multipotent stem cells which possess the capacity to get differentiated into ectodermal, mesodermal and endodermal lineages. Cell surface markers such as cluster of differentiation (CD) CD29, CD44, CD73, CD90, CD105 are expressed by MSCs and they lack the expression of human leucocyte antigen (HLA), CD14, CD34, CD45. Due to their homing ability, multilineage potential, secretion of antiinflammatory molecules and immune regulatory effects, they are considered to be a promising source in the treatment of inflammatory, autoimmune and degenerative diseases. The role of MSCs as disease models in treating various chronic diseases is discussed below. Rheumatoid arthritis: Loss of immunological self-tolerance leads to rheumatoid arthritis (RA) which is a joint inflammatory disease. MSCs were found to play an important role in the disease recovery and in reducing the progression of disease in the preclinical animal model studies (Van Velthoven et al. 2011). Mesenchymal stem cells in the treatment of neonatal ischemic brain damage: Transplantation of MSCs into neonatal animal models for the treatment of ischemic brain injury was shown to effectively reduce the lesion volume and improve its functional outcome. Since neonatal brain will be in developmentally active phase, better efficiency is observed with the transplantation of MSCs than in experiments involving the use of adult models of stroke. Also this was shown that treatment with MSCs after neonatal hypoxic-ischemic (hI) and brain injury resulted in axonal remodelling and enhanced neurogenesis thus underlying the improved functionality (Peltzer et al. 2018).

14.2.5 Progress in In Vitro Disease Models 3D bioprinting, an emerging and promising technology, can be used for the fabrication of complex tissue constructs with mechanical properties and biological components tailored to the need. Scientists are able to develop functional tissue models for use in in vitro drug screening and disease modelling by precisely positioning materials and cells due to advancements in this technology. With the use of 3D transformative technology, cells, growth factors and bioinks including hydrogels can be positioned precisely so as to create 3D in vitro culture environments (Jammalamadaka and Tappa 2018). This helps in developing biomimetic tissue models with native tissue architecture including vasculature and cellular composition for studying disease mechanisms using in vitro models (Markstedt et al. 2015).

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Gene Therapy and Its Applications

Any anomaly in the genetic makeup of an individual leads to genetic disease. Genetic diseases abnormalities mostly result due to mutations in genes. Changes in genetic makeup can range from discrete mutations of a single nucleotide base in the DNA of a gene to an entire abnormality in chromosome which involves the addition/subtraction of the whole chromosome or set of chromosomes. These changes can be of acquired in nature or mutations in pre-existing genes or genetically inherited disorders from parents. Genetic diseases like alpha- and beta-thalassemia, cystic fibrosis, Leber’s hereditary optic atrophy (LHON), sickle cell anaemia, Marfan syndrome and multifactorial disorders like heart disease, diabetes, high blood pressure, arthritis, cancer, Alzheimer’s disease, obesity, etc. are serious genetic disorders causing enormous health burden affecting more than 300 million people worldwide (Hussain et al. 2014). To get better knowledge of specific genetic factors in medical field, along with various challenges genetic diversity throws an opportunity in developing new diagnosis, differential treatment, risk factor identification, and eventually in effective prevention and cure of human chronic diseases (The International HapMap 3 Consortium 2010). During the 1960s and early 1970s, with the advent of recombinant DNA technology, it was shown that using cloning of foreign genes in mammalian cells in vitro, genetic defects and disease phenotypes can be corrected. This has paved the way to realize the concept of gene therapy which was broadly accepted by scientist and physicians in treating disorders through insertion of a gene into a patient’s cells rather than using drugs or surgery (Friedmann 1992; Selkirk 2004). Gene therapy is the approach of modulating genes for curing diseases. It has shown to be promising where traditional methods of medication/surgical treatment for a number of diseases mainly genetic diseases, viral infections, cancer and inherent disorders do not have cues. There are several methods of gene therapies which includes “knocking out”, “knocking in” and replacing the gene of interest/ disease. Although it is a promising technique to cure several diseases, it remains to be a risky procedure. Selection of gene delivery systems is the foremost thing for a successful gene delivery. For this complete knowledge of the targeting cell and the interactions between it and the delivery system is very important. Mostly these delivery systems consist of three parts, a gene regulatory system which regulates the expression of the gene in target cell, a gene of specific function/interest and a gene delivery system that delivers them to targeted cell. In addition, the stability of the foreign gene inserted by the gene delivery system into target cell is of utmost importance. There are different gene delivery systems important of which are viral and non-viral systems. Viral gene delivery systems include adenovirus, retrovirus, herpes simplex viruses and adeno-associated viruses, etc. Examples of DNA-based viral vectors are poxvirus, lentivirus, adeno-associated virus, adenovirus, human foamy virus (HFV), retrovirus and herpes virus (Gonçalves 2017). These DNA-based viral vector gene delivery systems integrate into the genome and last long. RNA based viral vectors for gene delivery are oncoretro-viral vectors, human

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foamy virus and lenti-viral vectors. In cancer treatment increased usage of oncolytic viruses (OVs) for gene therapy is under focus (Howells et al. 2017). A large spectrum of non-viral gene delivery systems are available of which cationic biomedical polymers like poly(L-lysine)(PLL) derivatives, polyethylenimine (PEI) and polysaccharides are important in the present context of this chapter. Various research groups have worked on the potential benefits of polymeric carriers as gene delivery systems. Cationic biopolymers like liposomes and chitosan derivatives are used for the purpose (Hulin-Curtis et al. 2016; Ginn et al. 2018). Others such as “lipoplexes”, cationic non-viral lipid-based gene carriers, are under clinical evaluation (Sung and Kim 2019). Both chitosan and pectin which are natural polysaccharides are studied as gene delivery systems by Morris et al. (2010). Chitosan has poly-N-acetyl-D-glucosamine, a polycationic derivative, has positively charged molecules for muco-adhesion whereas pectin has structurally segmented into smooth and hairy regions made of homo-galacturonan (HG) and rhamnogalacturonan (RG-I). Also, poly(L-lactide) which is a cationic polymer that can conjugate with the targeting ligands has been widely studied (Kabanov and Kabanov 1995; Stone 2010). Introduction of genetic material into the target cell through its cell membrane by means of physical methods is a mode of non-viral gene delivery. The physical method includes sonoporation, ballistic DNA injection, needle injection, magnetofection, photoporation and hydroporation. Target specific, non-immunogenic vectors are used as gene delivery systems capable of producing the desired effect by delivery of required amount of transgene expression is used for therapeutic applications (Katz et al. 2013). Mini plasmid DNAs have several advantages as delivery vehicles as they are minimally recognizable and can be designed in tissue/cell type specific manner and are easily inducible (Sum et al. 2014). Transposons are yet another gene delivery system capable of inserting the transgene permanently with higher biosafety norms which makes it advantageous to use instead of using other delivery systems which are non-integrating. Transposon system possesses comparatively superior biosafety profile and has the capability to insert transgene constructs permanently these properties make the SB approach more advantageous over viral gene delivery approaches and non-integrating non-viral vectors (Kebriaei et al. 2017; Tipanee et al. 2017). In other study, the authors have shown that DNA condensation and delivery can be carried out with polymeric gene carriers such as polyethylene glycol, dextran, hyaluronic acid (Hsu and Uludaǧ 2008; Dhanoya et al. 2011). Also, gene delivery systems are categorized into two types based on the targeted cells: (1) germline gene delivery systems, (2) somatic gene delivery systems. Due to ethical issues most studies in humans are confined to somatic gene delivery systems which are further categorized into in-situ, ex vivo and in vivo delivery systems.

14.3.1 Applications Gene therapy is the result of the advancements in genetics and bioengineering and technology. The most important application of gene therapy is in treating diseases

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with genetic basis by delivering the necessary genes to the target cells for expression of necessary proteins or by hindering the expression of a defective gene. Gene therapy has been proven to be a successful technique when compared to the protein therapeutics since it is easier to sustain the expression of therapeutic gene with longer life-span (Gonçalves 2017). Though research in gene therapy started over two decades ago, it is still considered to be in juvenile stage due to its enormous scope of development and applications. The first clinical trial of gene therapy approved by US-FDA was performed on 14 September 1990 on a ADA-SCID patient by replacing the defective gene in the blood cells of the patient with a functional variant (“The first gene therapy”. Life Sciences Foundation. 21 June 2011). Advent of gene therapy made its mark by providing cues in the treatment and prevention of a spectrum of genetic diseases such as cancer, diabetes, sickle cell anaemia, cystic fibrosis, haemophilia, muscular dystrophy, heart and Parkinson’s disease and some viral infections such as SCID. In vivo or ex vivo gene transfer involving CRISPR based gene editing and induced pluripotent stem cells has opened new vistas in genebased therapies in the recent times (Heyde et al. 2006).

14.3.2 GT in Tissue Engineering and Regenerative Medicine The real promise of gene therapy is that it rectifies the cause of disease at genomic level which could in fact result in a lifelong cure without the need to take pills every day or to have painful injections. The main aim of gene therapy is to deliver correct genetic material to target cells, to facilitate them with the missing/correct functional gene for curing the disease condition. Gene therapy is a very promising strategy and is being typically used to treat inherited diseases like haemophilia, cystic fibrosis, inherited retinal degeneration, severe combined immunodeficiencies, β-THALASSEMIA, infectious diseases, cardiac disease, neurological disorders, cancer, etc. (Farrar et al. 2012; Thrasher and Williams 2017; DiCarlo et al. 2018; Cavazzana and Mavilio 2018). Genetic diseases are caused due to errors or mutations in the genes which results in the loss or change in function of the gene products, i.e., RNA or protein molecules. These changes can be inheritable or can happen spontaneously. In gene therapy, the defective gene is replaced by healthy functional gene. Some of the therapeutic applications of gene therapy have been discussed below with respect to different disease conditions.

14.3.2.1 Heart Diseases Heart is a sort of strong muscular organ that pumps blood to complete body through the blood vessels of the circulatory system. It is a complex organ with multiple cell types that work in coordination to perform their individual specific task. In addition, they respond to external stimuli and systemic changes along with internal feedback mechanisms by sensing the humoral and neurogenic factors which simultaneously play important regulatory role. As heart is composed of various cell types like the cardiomyocytes, fibroblasts and the cardiac pacemaker cells, culturing of these cells in the lab do not make up for construction of a reliable heart disease model. In vitro

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models of heart help to understand the mechanistic functions of heart in health and disease, and to evaluate the safety and efficacy of potential therapeutics. Lot of work was carried over the years on the use of stem cells subjected to gene therapy to replace and restore damaged heart cells. The major concern in using this method is on improving cardiac cell function (Tilemann et al. 2012). The mechanism underlying the functional failure of heart has been elucidated at the level of expression of genes in cardiomyocytes. It has been attributed to abnormal Ca2+ handling and deficient contractility due to variations in calcium-cycling genes expression levels. Ca2 + ATPase (SERCA2a) plays an essential role in regulating the excitation/ contraction coupling. Systolic and diastolic dysfunction in heart failures is commonly driven by deficiency of SERCA2a (Zarain-Herzberg et al. 2012). This dysfunction of heart can be rectified by replacing of SCRCA2a through gene transfer technology using recombinant adeno-associated virus rAAV1/SERCA2a, which restores Ca2+ homeostasis resulting in improvement of cardiac function significantly. Normalizing SERCA2a deficiency via gene transfer rectifies the contractile abnormalities, decreases ventricular arrhythmias, reduces oxygen consumption of myocardiocytes and improves energetics, helps in electrical remodelling and survival in heart failure patients (Jessup et al. 2009; Lipskaia et al. 2010). Use of in vitro and in vivo models paved the way for the vast knowledge resourced towards the development of genetic modifications of the diseased heart. In order to overcome the challenges posed due to specific cell targeting, entry barriers, persistent expression and to circumvent the host anti-vector responses focus needs to be kept on developing improved delivery vehicles.

14.3.2.2 Lungs Diseases Respiratory illness is a common problem worldwide. The most common causes of respiratory diseases are smoking, infections, allergic reactions and pollution. Also, genetic conditions may lead to aggravation of lung diseases caused by environmental factors. Lung gene therapy has been advocated in a number of preclinical animal experiments such as targeting genetic diseases like cystic fibrosis and α1-antitrypsin deficiency, transplant rejection, lung injury, asthma, allergy, pulmonary arterial hypertension and lung cancer (Albelda et al. 2000; Kolb et al. 2006). Other diseases being targeted are pulmonary inflammation, surfactant deficiency and malignant mesothelioma (Curiel et al. 1996). Drumm et al. (1992) showed the possibility of delivery of a normal CFTR gene using retrovirus transduction into the adenocarcinoma cell of a CF patient. In CF epithelial cells, CFTR gene expression was linked to the regulation of cAMP-dependent Cl- channel. Rosenfeld et al. (1992) showed that adenovirus-mediated human CFTR DNA transfer into a cotton-rat model resulted in expression of the mRNA along with the expression of functional protein. Later, Zabner et al. (1993) worked on adenoviral gene therapy studies in humans. Though subsequent trials revealed the evidence for CFTR gene expression, they are limited in achieving clinical efficacy. Studies with shortened R domain deleted-CFTR using rAAV-mediated gene delivery have shown correction of CF phenotype in the intestinal organoids obtained from CF patients and in CF mice nasal mucosa (Vidović et al. 2016). Another application of gene therapy was its use in asthma.

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In particular its use in corticosteroid-resistant asthma patients and in patients with uncontrolled asthma where high doses of corticosteroids are to be administered is of at most importance. Gene therapy can be used to target the over expression of T-helper (Th) type 1 cytokines which modulate the Th2 cytokine reactions in bronchial asthma (Mathieu et al. 1999). Also it was shown that local antiviral immunity which was suppressed in Th2 dominant environment was restored due to over expression of IL-12 in bronchial asthma (Hogan et al. 1998). Although tremendous progress has been achieved in the area of lung gene therapy there are limitations for use in clinical application of this technology to human pulmonary disease due to various unresolved problems. With the evolution of delivery vector technology, gene therapy for a number of lung diseases can be a reality in near future (Pranke et al. 2019).

14.3.2.3 Liver Diseases Liver is the organ which detoxifies various metabolites, synthesizes proteins and produces biochemical entities necessary for digestion and growth. Metabolic and genetic defects due to inheritance are the cause for early chronic liver diseases. Majority of these diseases are caused due to defects in an enzyme or transport protein which modify the metabolic pathway thus executing a pathogenic action in the liver (Scorza et al. 2014). Deficiency of Alpha-1 antitrypsin is the most common hereditary genetic disorder which causes liver disease (Mitchell and Khan 2018). SERPINA1 gene expression leads to Alpha-1 antitrypsin (AAT) production in hepatocytes. SERPINA1 gene mutations are found to be clinically relevant and one such mutation Z (Glu342Lys) leads to misfolding of the alpha-1 protein, resulting in damage of the liver due to cirrhosis (Konietzko 2005). Gene therapy approaches and gene editing are used in the treatment for removal of the source of the toxic protein in the liver, along with ramping up the production of healthy alpha1 proteins (Borel et al. 2017). Gene therapy results in recapitulating the key functions of SERPINA1 gene product Alpha-1 antitrypsin and provides scope for disease treatment (Ercetin et al. 2019). 14.3.2.4 Kidney Diseases Over the past 20 years, acute and chronic kidney diseases like autosomal dominant polycystic kidney disease (ADPKD) and autosomal recessive polycystic kidney disease (ARPKD) as well as host rejection have seen little or no improvement in treatment methods. With the advent of gene therapy for use in kidney pathologies, delivery of genetic material to damaged cells in the kidneys resulted in differentiated expression of essential genes or suppressed the expression of pathogenic genes in different kidney cell types (Van Der Wouden et al. 2004). Targeted gene therapy could be an ideal treatment toward developing gene therapy to treat chronic kidney disease (Tomasoni et al. 2008; Ikeda et al. 2018; Davis and Park 2019). ADPKD occurs because of a mutation in one of two genes PKHD1 and PKHD2 which produces the two building block proteins of kidney, polycystin 1 and polycystin 2, respectively. Work by Ikeda et al. (2018) in kidney mesenchymal cells has showed that Gli2 when knocked out by injecting Anc80-Cre virus into Gli2-loxed

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homozygous or heterozygous mice results in the activation of canonical Wnt signalling transduction in kidney myofibroblast homozygous by Anc80 synthetic virus. This emphasized the importance of gene therapy in curing kidney diseases.

14.3.2.5 Brain Diseases Brain is the most complex organ present in human body. It serves as the central organ of the human nervous system. The developmental and functional aspects of brain are affected by the genetic disorders of brain. These genetic disorders are caused due to random mutations in the genes or mutations induced due to adverse environmental exposure like cigarette smoke, etc. These neurological disorders which include epilepsy, Alzheimer’s disease, multiple sclerosis, Parkinson’s disease and stroke result in various anomalies including loss of memory and ability to perform daily activities. A number of these disorders do not have proper medication and cure and require specific effective treatments. Gene therapy for these diseases offers an effective and promising cure for both genetic and acquired brain disease (Ingusci et al. 2019). Application of adeno-associated virus (AAV) based gene therapy for treatment of these diseases was shown to be promising along with metabolic disorders of the central nervous system like Canavan Disease, Phenylketonuria and Niemann–Pick disease (Gessler and Gao 2016).

14.4

Advances in Gene-Based Therapies and Its Applications in TE and RM

Gene therapy possesses the potential to cure large number of diseases which does not have any cues in the traditional methods of treatment and offers promising results. The potential developments in research on gene therapy in the last three decades have been promising and fruitful. Gene therapy is about delivery of specific nucleic acid segment/gene at specific location in targeted cell where it is needed to restore the function of that gene. Advances in research has paved the way of our understand of the usage of different nucleic acid derivatives such as RNA, non-coding RNAs and m RNA along with relatively older transposition mechanisms, pDNA and antisense oligonucleotides in gene therapies. Gene editing technologies like CRISPR/Cas9 system, designer nucleases, including transcription activator-like effector nucleases (TALEN) and zinc finger nucleases (ZFN) which induces double stranded breaks at the target site for insertion of gene of interest are some of the latest additions in the development of gene-based therapies (Urnov et al. 2010; Gaj et al. 2013; Jung and Lee 2018).

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14.4.1 Adenovirus as Gene Therapy Vectors 14.4.1.1 Adenovirus Based Therapy Using Antisense/Small Interfering RNA In gene therapy adenovirus vectors are used in gene silencing mechanisms. The gene of interest to be silenced is targeted with single-stranded antisense oligonucleotides (ASOs) delivered using Ads. The antisense RNA of the targeted gene is produced by transfecting the targeted cells with recombinant Ad-based vectors with partially cloned reverse sense sequences of the targeted gene. Adenoviruses expressing an antisense fragment of c-myc gene have been shown to induce differentiation cycles normally and suppress cell proliferation. 14.4.1.2 Cancer Vaccines Based on Adenoviruses Gene-based vaccination in the gene therapy of cancer has evolved as one of the most promising approaches in recent years. The various strategies that are being currently deployed for targeting cancer using cancer gene therapy include (a) inducing apoptosis by gene expression or enhancing the sensitivity of tumours to conventional drug/radiation therapy; (b) inserting wild genotype tumour suppressor gene; (c) using antisense (RNA/DNA) approach to block the expression of oncogenes; and (d) enhancing tumour cell immunogenicity to stimulate the recognition of immune cell. In the development of gene-based vaccines for cancer, Ads were the preferred choice of vectors for delivery of cancer cell-associated antigens and other immune stimulatory cytokines. The potential advantage of using Ads as tumour vaccine vehicles is in delivering high concentrations of localized immunomodulatory molecules and tumour-associated antigens, promoting the effective processing and expression of MHC and in stimulating cell-mediated immunity (Das et al. 2015). Adeno-associated virus (AAV) are small virus which can deliver DNA into cells. Gene therapy by AAV helps in addressing a number of different diseases because of their ability to deliver genes to a number of different tissues inside the body (Tani and Sufian 2011). Gene therapy has its major application in patients particularly suffering from inherited disorders. Also, malignant diseases and other acquired diseases have become the additional targets in gene therapeutic strategies (Mohr et al. 2002). 14.4.1.3 Gene Therapy: Applications with Haematopoietic Stem Cells Due to their higher potential to self-renovate and longevity, haematopoietic stem cells are preferred as ideal targets for gene transfer. Production of gene transfer vectors to develop induced pluripotent stem cells (iPSCs) with an added phenotype marks a good example of the use of gene therapy in combination with stem cells (Gonçalves 2017; Tachibana et al. 2017). A better explanation of this was illustrated from the studies involving transplantation of adult hepatocytes or iPSC derived hepatocytes for liver transplantation in patients with chronic liver disease and being infected by hepatitis virus. In such cases, along with gene transfer for conversion of stem cells to hepatocytes, a vector that encodes short hairpin RNA

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which is directed against the hepatitis virus is needed to avoid reinfection and to confer resistance to reinfection in transfected cells. These resistant cells are found to repopulate the liver in due course and restore the normal function of liver (Doss and Sachinidis 2019).

14.4.1.4 Gene Therapy and CAR-T Chimeric antigen recipient T (CAR-T) cell therapy is a form of immunotherapy which involves reprogramming of T lymphocytes of the patients themselves so that they recognize and attack the tumour T-cells. The main drawback of CAR-T therapy is in its identification of the target epitope by CAR expressed in non-tumour cells. 14.4.1.5 Gene Therapy in the Treatment of Adult-Onset Glaucoma Glaucoma refers to a set of eye diseases which are characterized by progressive and irreversible degeneration of retinal ganglionic cells. Since the axonal projections of these retinal ganglionic cells constitute the optic nerves, various strategies to inhibit apoptotic pathways of these cells were explored. Gene therapy approach for glaucoma aims at exploring the therapeutic potential of anti-apoptotic proteins which helps in inhibiting the progression of RGC apoptosis (Ratican et al. 2018).

14.5

Biomaterials in TE Based on GT

Both synthetic and natural biomaterials are being developed for use as gene delivery systems that can deliver the desired nucleic acid to the targeted site without getting impeded by the physiological mechanisms due to their packing of the agent to be delivered. These can enhance the functionality of delivered molecules and provide additional stability in physiological conditions due to the amalgamation of required factors/elements during their fabrication. These fabrication techniques allow the development of custom made biomaterials for various tissues ex vivo, and to develop new gene transfer techniques for constructing devices ex vivo for transplantation using modified cells (Tai and Sun 1993). In gene therapy, biomaterials are now being used to develop important technologies for transfer mechanisms. Different groups have worked on developing biomaterials and non-viral techniques which are less immunogenic and more safe than viral delivery systems for the delivery of nucleic acids. Liu et al. (2010) have demonstrated that nucleic acids when packed with cationic and lipidic biomaterials in combination have improved delivery over using either of the single moiety alone. For the first time CAR T-cells were developed in vivo by using a nanoparticle-based Poly(B-amino) ester polymer system (Smith et al. 2017). Olden et al. (2018) have shown that with the usage of cationic pHEMA-g-pDMAEMA polymer maximum transfection was achieved in primary T-cells with comb and sunflower shaped polymers. Smith et al. (2017) used alginate scaffold with collagen-mimetic peptide and adjuvant silica microparticles and observed that local and systemic response was elicited with the release of T-cells with adjuvant compounds. Enhanced migration of T-cells targeting glioblastoma was observed by Tsao et al. (2014) with the usage of PEG-g-Chitosan Hydrogel

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(PCgel). Polyisocyanopeptide hydrogel are used by Weiden et al. (2018) as injectable thermo-responsive scaffolds which allows in vivo survival and migration of T-cells.

14.6

Challenges and Future Prospects

Clinical safety is one of the major factors to be considered in biomaterials development. Biomaterials synthesized and fabricated for precise needs for use in therapeutic applications are the need of the hour. Responsive systems are to be developed for the systematic release of different molecules from the scaffolds. Although different delivery systems are available, there are both advantages and disadvantages of using them in gene therapy. Future research should be focused on further developing non-viral delivery vehicles based on biomaterials since they are more biocompatible and offer custom made delivery of nucleic acid agents. Use of advanced fabrications techniques for the development of in vitro disease models needs to be focused. This will help in creating biomimetic in vitro models thus providing wider scope to get proper insights into the aetiology of the disease. Both TE and RM can be reinvented into newer avenues with the application of gene-based therapies and use of in vitro disease models. Advanced gene editing technologies like CRISPER/Cas9 systems along with stem cell technology add new horns to TE and RM for use in therapeutic applications. More number of in vivo animal model and clinical studies need to be performed for proper evaluation of these systems. Gene therapy applied with TE and RM holds promising hope for treatment and cure of various diseases for which conventional medication is a distant sight. All said and done, ethical aspects need to be taken care of while using these technologies, since uncontrolled use of these technologies can prove to be disastrous.

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Nanobiomaterials in Tissue Engineering and Regenerative Medicine: Current Landscape and Future Prospects

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Nagaraju Shiga, Dumpala Nandini Reddy, Birru Bhaskar, and Vasagiri Nagarjuna

Abstract

It would be rather interesting aspect to pursue advanced, multidisciplinary, and emerging technologies/studies in any field of research. This chapter deals with such multidisciplinary advanced research topics that have emerged into prominence in the last decade or so. In this chapter we will be discussing about Nanobiomaterials in the context of Tissue engineering and Regenerative medicine, there present status, scope, advances, and future prospects. Biomaterials of nano basis have already made their mark in TE and RM and research in this area is thriving to meet the never-ending demand of human health care. The various types of nanobiomaterials, their use, and functional significance in different TE & RM aspects; the various methods/techniques used for the preparation of these nanobiomaterials have been detailed in this chapter. Overall, this chapter presents a vivid picture of nanobiomaterials use in TE & RM at present and their future prospects. Keywords

Nanobiomaterials · Nanotechnology · Encapsulation · Nanofibres · Nanopattern

Nagaraju Shiga and Dumpala Nandini Reddy contributed equally. N. Shiga · D. N. Reddy P.V. Narsimha Rao Telangana Veterinary University, Khammam, Telangana, India B. Bhaskar Prof. Brien Holden Eye Research Centre, LV Prasad Eye Institute, Hyderabad, Telangana, India V. Nagarjuna (*) Society for Biological Chemists India, Bangalore, India # The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 B. Bhaskar et al. (eds.), Biomaterials in Tissue Engineering and Regenerative Medicine, https://doi.org/10.1007/978-981-16-0002-9_15

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Introduction

Nanotechnology has been an emerging field with its footprints in various disciplines of science, technology, and research. With the growing number of applications to its credit, it has made its mark in Tissue Engineering, Regenerative Medicine, and Pharmaceutics in the last decade or so. Nanotechnology is a discipline that utilizes the properties of its nanometer size particles (1–100 nM) for various applications. Nanotechnology is also used to mimic the nanolevel topography of natural cells. Advances in developing new nanoscale materials together with stem cell biology research are paving the way in realizing the formation of different organs in vivo by tissue engineering. “Biomaterials are defined as natural or synthetic materials or constructs interacting with biological systems use multidisciplinary ideas and technologies from medicine, pharmaceutics, biology, tissue engineering, material science and chemistry” (Atala et al. 2010). Biomaterials of nanosize are termed as Nanobiomaterials. Nanobiomaterials have higher surface area to volume ratio facilitating more functions physically and chemically compared to bulk materials. Tissue regeneration capacity of tissues like bone, tendon, cartilage, cardiac muscle, and nerve is limited or absent if the degeneration is severe or defect is major (Liu et al. 2012). “Tissue engineering” deals with development and replacement or repair of biological tissues/organs using various cells, materials, and biological factors alongside engineering technologies. “Regenerative Medicine” is a branch of medicine that develops methods to regrow, repair or replace damaged or diseased cells, organs or tissues. Regenerative medicine includes the generation and use of therapeutic stem cells, tissue engineering, and the production of artificial organs. It can even correct congenital defects. The major constraints faced in tissue/ organ transplantation are immunological complications due to donor/recipient immunocompatibility and limited organ or tissue donor availability which can be surpassed through regenerative medicine. Nanobiomaterials play a pivotal role in current regenerative medicine strategies as they can simulate the native extra cellular matrix (ECM); modify cell topography, function and behavior thus mimicking the natural cell in all aspects. Regenerative medicine utilizes living cells, signaling molecules (Slaughter et al. 2009), and biomaterials as scaffolds (Bajaj et al. 2014) to rejuvenate its natural physiology by anatomically fabricating the deteriorated tissue (Liu et al. 2012). Scaffolding is a key component in tissue engineering as it provides structural support for growth and development of tissues (Chan and Leong 2008). Scaffolds have to be analogous to native ECM and must possess an architecture, cell and tissue compatibility, bioactivity, and mechanical property in order to provide vascularization, biological, and physical cues and stability to the native tissues, respectively (Chan and Leong 2008). The greatest challenge lies in the fabrication of a precise scaffold mimicking the native ECM, a multi-functional, dynamic and complexly composed 3D structure. Another limitation is that pre-fabricated scaffold often cannot accurately fit into irregular or defective tissues unless fabrication of scaffold is complexly designed. 3D porous structures support cells for adhesion, provide spatial and physical

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structure to the tissue for cellular activity. The complexity of scaffold design and surgical procedure can be minimized by using an injectable scaffold that fills into any geometry of the defect. Also, recovery time is also considerably reduced in comparison to implantation of scaffold (Zhang et al. 2016a, b). It was shown that when compared to conventional biomaterials, nanobiomaterials like carbon nanofibres possess improved properties to support the growth of osteoblasts (Price et al. 2004). The nanosize of biomaterials makes them highly reactive and exhibit improved electrical, electrochemical, optical, and magnetic properties (Singh 2011). For the purpose of scaffold fabrication different scaffolding approaches are used in tissue engineering like Pre-made porous scaffolds for cell seeding, Decellularized extracellular matrix for cell seeding, Cell sheet engineering, and Cell encapsulation in self-assembled hydrogel matrix. Pre-made porous scaffolds are generally made of porous biodegradable synthetic or natural materials on which cells are seeded. Immunogenicity is a major problem usually encountered in this approach. Use of Decellularized ECM from allogenic or xenogenic tissues is the most nature-mimicking scaffolding approach since it involves removal of cells from the matrix using various methodologies like physical, chemical, and mechanical means and retaining the ECM components. Antigens from tissues are removed and hence immunologically compatible tissues are replaced. With this approach tissues of heart valves (Knight et al. 2008), vessels (Borschel et al. 2005), nerves (Hall 1997), tendons and ligaments (Ingram et al. 2007) were engineered. Cell sheet engineering is a latest technology applied in regeneration of tissues without the use of a scaffold as in traditional therapy. With the help of temperature-responsive surface monolayer of cells are procured and cultured for regeneration of tissue (Egami et al. 2014). Cell encapsulation in self-assembled hydrogel matrix is yet another approach where cells are encapsulated in semipermeable membrane or homogenous solid matrix (Chan and Leong 2008). Advances in nanotechnology and nanofabrication techniques led to the development of biocompatible nanofibres, nanoparticles, nanopatterened surfaces, nanoporous scaffolds, nanowires, and carbon nanotubes which are used in advanced tissue engineering applications. Nanoparticles are used in scaffolds to improve biodegradation, corrosion rates, and mechanical properties. Nanoporous biomaterials are of great use in bone tissue engineering and they are known to exhibit greater surface area, enhanced diffusion through pores and adhesion to proteins and cell integrity. Nanopatterened materials also with their pillars and ridges like structures simulate topographical features in tune to mechanical and surface area of natural cells. These properties not only elicit stem cell differentiation but also prevent tissue fibrosis. Nanofibres are developed by electrospinning technique to mimic cell architecture and exhibit therapeutic properties. Carbon nanotubes in scaffolds provide mechanical strength, electrical conductivity aiding in the cardiac and neural tissue engineering (Padmanabhan and Kyriakides 2015). There are two types of nanobiomaterials; “Natural” are the ones derived from polysaccharides, proteins, and carbon-based nanomaterials. Polysaccharides include biopolymers like chitin, chitosan, and cellulose which are abundant in nature, biocompatible and available at low-cost. Chitin is the second most available

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biopolymer after cellulose. Its ferromagnetic property helps in drug-tracking (Jayakumar et al. 2010). Chitosan is derived from chitin and is superior to chitin, owing to its solubility in organic solvents and water. With chitosan, nanofibres and nanomaterials can be manufactured which possess relatively greater chemical, mechanical, and physical properties like tensile strength, porosity, high surface area, and conductivity in comparison to pure chitosan. Usage of chitosan nanofibers in wound dressing and as bioscaffolds in tissue engineering is in vogue. With advances in nanotechnology, chitosan-based nanomaterials with novel properties will make their way in future health care system. Cellulose is the most available biopolymer and has high surface area with negative charge providing lot of space for drugs to bind. Nanoscale cellulose fibers and composite materials with cellulose ensure desirable properties like stiffness and strength. They are used in dental tissue regeneration and assist in simplifying the surgeries involved. Cellulose nanofibres incorporation in polymer reinforcement is a new area of research (Siróand Plackett 2010). Protein-based nanobiomaterials aid in tissue repairing and regenerative medicine. Gelatin, collagen, silk, fibrin, and plant proteins are promising as nanobiomaterials due to their biocompatibility and other versatile properties. Collagen is a protein native to human body and constitutes 20–30% of the total protein and present everywhere in the body predominantly in skin and bone. Strong biocompatibility, weak antigenicity, and fast biodegradability make them good biomaterials with a wide variety of biomedical applications. Gelatin, also a protein is widely used in drug delivery and other applications. Proteins like Gelatin and Fibroin are widely used in drug delivery and other applications. Carbon-based nanobiomaterials like carbon nanotubes and graphene possess phenomenal physical, chemical, and mechanical properties like high surface area to volume ratio, electrical conductivity, stiffness and thermal conductivity. Carbonbased nanomaterials like graphene are used as bioscaffold due to its great mechanical property owing to strong covalent bonds and great cell adhesion potential. Graphene nanoparticles coated with nanometallic strontium have good potential in bone tissue engineering. Carbon nanotubes are deposited on neural electrodes to restore nerve function (Keefer et al. 2008). “Synthetic materials” other than carbon-based nanobiomaterials like polylactic acid (PLA), poly lactic-co-polyglycolic acid (PLGA), poly L-lactic acid (PLLA), poly acrylic acid (PAA), polyglycolic acid (PGA), polycaprolactone (PCL), poly vinyl alcohol (PVA), polyurethane (PU), poly l-lactic acid-co-polycaprolactone (PLLA-PCL), polyethylene oxide (PEO), polymethyl methacrylate (PMMA), and polyethylene terepthalate (PET) (Mogoşanu et al. 2016). Studies have shown that the use of combination of different types of nanobiomaterials (compound nanomaterials) is advantageous rather than using nanomaterials of a single category alone which have some limitations for themselves. Nevertheless, compound nanomaterials or modification of existing categories also enhances their potential in application. Nanoparticles with nanofibres incorporated and nanopatterns imprinted on electrospun nanofibres are studied for this purpose. These nanofibres are non-injectable, have insufficient porosity and pore

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size and hence advanced nanofibres with different composition like collagen, gelatin, silk fibroin, chitosan, etc., are utilized (Chen et al. 2020).

15.1.1 Bio and Immuno Compatibility of Nanobiomaterials A nanoengineered material with perfect architecture has to well integrate into the desired tissues and requires to be physiologically active for therapeutic and clinically purposes without eliciting any undesirable local or systemic responses. Testing procedures to verify these biocompatible properties as per FDA and ISO10993 guidelines or other similar standards are inevitable (Williams 2008; Hussein et al. 2016). In vitro analysis of these nanobiomaterials includes testing of tissue implantation, hemocompatibility, systemic toxicity, sensitization, cytotoxicity, genotoxicity, and subchronic toxicity to evaluate their biocompatibility (Williams 2003). Another priority is to look upon immunocompatibility as nanoengineered materials are transplanted with surgical procedures, a cascade of immune reactions which are complex, mount in response to the graft, corneal scaffolds or engineered tissue. Hence animal trials are a must prior to transplantation. Wide array of testing is essential to analyze the activation and differentiation markers of leukocyte subpopulation. The tests to perceive immune cell responses of sub types like B-lymphocytes, T-lymphocytes, granulocytes, and antigen presenting cells; and biocompatibility, physiological fluids interactions, local and systemic effects are good indicators of suitability of a nanobiomaterial. In vivo testing, i.e., preclinical trials on animal models after successful in vitro analysis is done to ensure bio and immunocompatibility of a novel biomaterial (Wicklein et al. 2019).

15.2

Nanobiomaterials in Tissue Engineering and Regenerative Medicine

Nanobiomaterials are widely used in tissue engineering and regenerative medicine due to their fabrication with latest methodologies involving phase separation, molecular self-assembly, and electrospinning. These techniques brought them regulation over size, shape, surface topography, physicochemical stability, mechanical strength and composition. Although this dynamic nature of nanobiomaterials simulated ECM of natural cells and made all types of interactions possible between cells, further studies on the toxicity of biomaterials used in TE and RM and measures to reduce it, along with development of highly biocompatible materials are the need of the hour (Mogoşanu et al. 2016). These novel nanobiomaterials are currently being used in anti-cancer therapy, drug delivery as polymers, gene delivery, molecular diagnostics, and stem cell engineering. Also, the use of nanobiomaterials helps in the controlled tissue assembly in tissue engineering therapies. Nanofibres are usually synthesized from compatible materials and are used as scaffolds in tissue engineering. Electrospinning is the latest and most common

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technology with which nanofibres are produced ranging in size from 50 to 5 μm. These nanobiopolymers are used in Neural Tissue Engineering (NTE) as bioengineered scaffolds or as vehicles in neurotrophic growth factors delivery. Electrospun Collagen nanofibres are used in repair of nerve tissue or spinal cord. They are shown to enhance myelin sheath formation and graft-host interfaces while repairing spinal cord injuries (Iannotti et al. 2003). Electrospun silk fibroin scaffolds are used in neural tissue regeneration since their mechanical properties helps in cell adhesion, growth, and proliferation (Hu et al. 2012). Nanofibrous electrospun biocompatible gelatin scaffolds are used for in vitro nerve stem cell differentiation and proliferation (Ghasemi-Mobarakeh et al. 2008). In vascular tissue engineering alginate hydrogels are used for controlled release of vascular endothelial growth factor which enhances angiogenesis and promotes normal tissue perfusion (Silva and Mooney 2007). Bilayered scaffolds, composed of poly-ε-caprolactone and type I collagen are used in the enhancement of engineering of human aortic endothelial cells (Ju et al. 2010). In cardiac tissue engineering, patches of alginate scaffolds that are nano wired with gold particles are used since they are found to enhance electrical communication with adjacent cardiac cells (Dvir et al. 2011). It was also shown that biodegradable synthetic nanofibres mimics extracellular matrix of cardiac tissue and helps in alignment of cells and their proliferation into the tissue filaments (Orlova et al. 2011). Muscle tissues with large defects are regenerated with highly aligned collagenPCL nanofibrous scaffolds containing skeletal muscle cells (Choi et al. 2008). Scaffolds of composite nature produced from PCL, multi-walled carbon nanotubes and PAA transmit electrical signals simulating contraction and relaxation of skeletal muscles (McKeon-Fischer et al. 2014). Biocompatible and biodegradable type II collagen nanobiomaterials are used as scaffolds in cartilage tissue engineering (Li and Hu 2015). It has been shown that in articular cartilage repair electrospun type II cartilage is used as scaffolds in the development of chondrocytes (Shields et al. 2004) and Electrospun hyaluronic acid nanofibers are used in articular cartilage repair (Lee et al. 2009). It was further shown that chondrogenic stem cell differentiation was also possible in porous scaffolds made of alginate, gelatin, and agarose (Awad et al. 2004). Hu et al. (2009) cited that chondrogenic and osteogenic differentiation of stem cells derived from mesenchymal cells of human bone marrow is possible on nanofibrous porous synthetic scaffold. Human anterior cruciate ligament (ACL) has poor healing capacity and hence ligament tissue engineering is mainly pertained to ACL regeneration with tissue grafts, stem cell therapy, and synthetic or natural scaffolds (Nau and Teuschl 2015). Another study has shown that Chitosan-polyethylene oxide electrospinned into nanofibres was found to be biocompatible with chondrocytes, providing great scope for cartilage tissue repair (Subramanian et al. 2005). Hydroxyapatite as nanobiomaterial is highly recommended as it is highly resorptive and used in material coating in dental tissue implants due to its crystalline nature. Titanium is also used as biomaterial in dental tissue engineering in combination with

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nanohydroxyapatite which is bioactive, owing to the metal’s non-reactive nature (Jimbo et al. 2012). Skin tissue engineering involves regeneration of dermal tissue or complete repair of a wound with nanobiopolymers that mimic human cells in producing ECM like growth factors, collagen, angiogenic factors (Mohd Hilmi and Halim 2015). For diabetic skin injuries, collagen and fibrin encapsulated with growth factors are used in active wound healing (Zahedi et al. 2010). For chronic wound healing, selfassembling elastin-like peptides act as keratinocyte growth factor (KGF) play a crucial role in reviving its normal structure (Koria et al. 2011). Scaffolds from silk fibroin-keratin from human hair are used in highly vascular dermal tissues which is biodegradable and biocompatible (Bhardwaj et al. 2015). Polycaprolactonecollagen-based nanomaterials as dermal substitutes provide good cell adhesion, proliferation, and movement of fibroblasts into nanofibrous matrix (Venugopal and Ramakrishna 2005).

15.2.1 Neural Tissue Engineering The scaffolds for nerve tissue engineering like any other tissue should not only be biocompatible, biodegradable, mechanically compatible, structurally porous or dynamic but also electrically permissive or stimulating as impulse conductivity is vital for good performance of nerve tissue (Cao et al. 2009; Ghasemi-Mobarakeh et al. 2011). Usage of scaffolds integrated with nanofibres, microfibres, and hydrogels, to simulate native ECM is in vogue (Fon et al. 2014). Nanomaterials are being widely studied for neural tissue engineering and schematic representation given in Fig. 15.1.

15.2.1.1 Types of Nano-Based Scaffolds Used in NTE Hydrogels They are formed by either physical or chemical crosslinking of polymers with specific designs to incorporate cell-binding sites that improve cell viability and differentiation; and neurotrophic factors to enhance nerve function (Bosworth et al. 2013). Albeit, the scaffolds exhibit poor mechanical strength and therefore has to be integrated with nanofibers, which has highly aligned chains of molecules and good surface topography with minimal defects (Boudriot et al. 2006). The mesh of nanofibers are identical to collagen in matrix, hence simulates native tissue environment, thus complements collagen and this is an inclusion to the existing properties of hydrogels. The nanofibers implanted into hydrogels can support biological activity with provision for sites of attachment to cells and are resistant to contraction or shrinkage during proliferation of tissue of interest (Butcher et al. 2014). There are direct applications of the composite hydrogels as they are fluid-gel free and hence can be used in therapies to regrow/repair/tissue engineer, injuries related to central nervous system (CNS): traumatic brain conditions, stroke, and spinal cord, rather than building nerve gaps in peripheral nervous system (PNS). Injectable fibrin hydrogels used in injured spinal cord are also known to enhance sprouting of nerve fibers and seals exposed dura mater (Johnson et al. 2009; Johnson

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Fig. 15.1 Application of various nanomaterials in neural tissue engineering. Reproduced from (Kumar et al. 2020) with permission # 2020 Elsevier B.V

et al. 2010; Taylor et al. 2006). HAMC, another injectable hydrogel is also known to reduce inflammatory reaction in the aftermath of spinal cord injuries (Gupta et al. 2006). Khaing et al. (2011) also cited that processed hyaluronic acid hydrogels that are implanted into spinal cord transection models inhibits macrophage infiltration and astrogliosis. Hydrogels act as vehicles for cell delivery to the target location post-spinal cord and brain injuries. When they are incorporated with embryonic human/rat stem cells into an injury of spinal cord are known to enhance growth, survival, and integration of neural tissue implanted, even regenerate 2 mm spinal resection lesion (Lu et al. 2012; Xiong et al. 2009). Self-assembling Scaffolds The spontaneous assembly of biomolecules like peptides, proteins, RNA, DNA, etc. into nanostructures like nanofibers, nanotubes, vesicles, helical ribbons, higher order structures like β-sheets and α-coils or even supercoils (Hosseinkhani et al. 2013), triggered by modified environmental parameters like pH, temperature, ionic strength, mechanical and electrical stimuli (Xu and Kopecek 2007), results in noncovalent interactions within molecules to form self-assembled scaffolds (Stephanopoulos et al. 2013). Zhang et al. (2015a, b) stated that the process in repair or regeneration of nerve tissue is engaging as selfassembling is an integral part in the development of native extracellular matrix at nanostructured levels.

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Peptide amphiphiles, as the name implies have hydrophilic and hydrophobic part, can also be tuned with amino-acids of required sequence which influences the secondary structures and enhances biocompatibility, are used in tissue engineering (Lock et al. 2013). Self-assembling peptide nanofibrous scaffolds (SAPNSs) have been studied for their use in PNS and CNS, both in vitro and in vivo (Saracino et al. 2012). RADA16-I, a self-assembling peptide in nanofibrous scaffolds showed positive results in directing the axons towards severed optic nerve thus play a pivotal role in reviving vision (Ellis-Behnke et al. 2006). SAPNSs incorporated with neural progenitor cells (NPC) and Schwann cells are implanted into a transected spinal cord. These scaffolds are shown to integrate into native tissue and have regrown the entire cavity without leaving any gaps as many host cells migrate into the scaffold. Angiogenesis also took place within the regenerating tissue. As NPCs are multipotent, they differentiate into neurons, astrocytes, and oligodendrocytes (Guo et al. 2007). SAPNSs can be administered into spinal cord lesions and injuries in brain in liquid form. Therefore they hold the advantage of not causing injury during injection; fills the entire lesion and incorporates water into it to form hydrogel (Holmes et al. 2000). This was investigated by Guo et al. (2009), who injected RADA 16-I into cortical lesions and observed that lesion was closed as cells migrated into it and reduced tissue loss. The disadvantage of RADA 16-I peptide is its inherent acidic nature which can damage tissue until it is brought to equilibrium before forming complex structures (Zhang et al. 2016a, b). Rosette nanotubes are also a type of self-assembling nanomaterials formed by integration of base-pair motifs, Guanine-Cytosine of nucleic acid, DNA, whose sequences align into nanotubes (Fenniri et al. 2002; Zhang et al. 2008). They are acclaimed for their tunability of size and surface chemistry, by linking peptide chains to G-C motifs (Zhang et al. 2015a, b), although their usage in neural tissue engineering is not studied practically, albeit they are potential candidates in tissue engineering (Sun et al. 2013). Electrospun Nanofibrous Scaffolds Electrospinning is a technique that is widely used in generation of nanofibrous scaffolds for tissue engineering purposes (Kanani and Bahrami 2010). This is a versatile technique which can be modified by employing diverse solutions and processing parameters to produce scaffolds with variety of mechanical and biological properties and morphologies (Vasita and Katti 2006; Sill and Von Recum 2008; Yang et al. 2006). More number of natural or synthetic polymers of potential used in preparation of scaffolds led to the eventually enhanced production of nanofibrous scaffolds creating scope for further development. In order to reduce cytotoxicity and improve bioactivity electrospun nanofibres can be integrated with self-assembling peptides. In neural tissue engineering, biodegradable polymers like polycaprolactone and polylactic acid are widely used as they are easy to regulate size and topography (Kwon and Matsuda 2005; Kumbar et al. 2008; Mobasseri et al. 2015; Hatamzadeh et al. 2016). The effects of variation in diameter of nanofibers in electrospun nanofibres were investigated by various scientists to find the phenomenal influence

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of diameter on neurite proliferation and migration with reduced aggregation in accordance with reduction in the diameter (Christopherson et al. 2009). Another topographical feature studied is fiber orientation which guides neurite regeneration in the direction of target tissues or organs. Neurites of primary dorsal root ganglia regenerated on electrospun PCL nanofibres showed orientation as per the nanofibres orientation. Parallel oriented Nanofibres are shown to guide the neurites along the fibers whereas randomly oriented nanofibres showed random growth of neurites (Xie et al. 2009). Carbon-based Scaffolds Carbon-based nanomaterials are unique and versatile, produce carbon nanotubes (CNTs), carbon nanofibres (CNFs) (Tran et al. 2009), graphene nanomaterials (Wang et al. 2011; John et al. 2015) that are potent for applications in tissue engineering. They are known to exhibit thermo-electric conductive properties. Graphene is a one carbon atom thick sheet in 2D hexagonal lattice (Katsnelson 2007) which is the thinnest and strongest nanomaterial known at present (Giem 2009). Graphene and its nanomaterials exhibit good physical properties like planar surface area, electronic flexibility, high mechanical strength, and exceptional thermal conductivity (Bhattacharya et al. 2016). Carbon nanotubes are secondary structures that are formed by graphene sheets rolled into single-walled and multiple-walled nanotubes (Dai 2002). The incorporation of CNTs into nanoscaffolds exhibit an improved interfacial area, conductivity, and electrochemical properties (Keefer et al. 2008). Cytotoxicity of CNTs at present is a major limitation in their use in tissue engineering applications. Cytocompatibility is dependent on properties like length of tube, concentration and dispersion of CNTs, hydrophobicity of nanomaterials which determine their long-term usage in vivo. Even though structural modifications in CNTs change the properties of electrical conductivity and thus improve biocompatibility and bioactivity, regulation of cell adhesion, proliferation, differentiation, matrix remodeling is possible with modification of surface chemistry and conjugation of signaling molecules onto CNTs (Yang et al. 2007). Neurons grown on 3D graphene films improve cell conductivity; promote cell proliferation and properties specific to it like differentiation into astrocytes and electrical signaling in differentiated cells along with those grown on 2D graphene films (Li et al. 2013). Electrical signaling has the potential to stimulate myelin sheath formation and enhance impulse conduction incase the material is used in vivo. Single-walled nanotubes integrated with polyethylene glycol, investigated in vitro, exhibited long neurite outgrowth, promotion of axonal repair and regeneration of injury related to spinal cord (Roman et al. 2011).

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Nanobiomaterials and Bone Tissue Engineering/ Regeneration

Bone tissue engineering is a complex and dynamic process. Orthopedic procedures like implantation repair and regeneration of bone have seen a rise with aging population. Bone is primarily composed of organic (mainly collagen) and inorganic (mainly hydroxyapatite) components with a hierarchical organization of molecules from nano to macroscale. Nanostructured scaffolds with architecture close to native ECM of bone are considered potential candidates for repair or replacement of tissues since the mechanical stimulation needed for bone TE is brought about by the use of magnetic nanoparticles that attract charged molecules into the cells altering cell physiological and biochemical environment (Balasundaram et al. 2014). Bone composition as three-dimensional nanostructures, when fabricated in accordance to the native bone tissue architecture is considered an ideal surface in order to restore functional segmental deossification.

15.3.1 Nanobiomaterials Used in BTE Natural and synthetic polymers like gelatin, chitosan, collagen, silk fibroin, graphene, carbon nanotubes (CNTs), poly (l-lactic acid) (PLLA), poly (lactic-coglycolic acid) (PLGA), and polyamide are the organic materials used in scaffolds for applications in bone tissue regeneration/engineering. The inorganic materials used are Hydroxyapatite, biphasic calcium phosphate (BCP), calcium phosphate cement (CPC), glass fiber, titanium dioxide and silver. These biomaterials are nanoscaled and used in various combinations to obtain optimal results which have been discussed below. Advanced regenerative medicine approaches incorporate stem cells into scaffolds composed of these nanobiomaterials, and such implants or substitutes are used for repair of bone tissue.

15.3.2 Nanohydroxyapatite (nHA) Hydroxyapatite, a principal inorganic mineral component of bone and dentine of both animal and human is mostly insoluble in solution. Nevertheless, in its nanoparticulate structure, i.e., as nanohydroxyapatite (nHA), when used as a biomaterial can be precipitated or incorporated into scaffolds to promote varied bone tissue applications like stem cell differentiation, skeletal defects, internal fixation, spinal fusion, etc., which are detailed below.

15.3.2.1 nHA in Stem Cell Differentiation During Bone TE Nanohydroxyapatite (nHA), titanium, calcium phosphate, graphene oxide, carbon nanotubes or the combination of these biomaterials (Sun et al. 2013; Xu et al. 2014; Liu et al. 2015) fabricated as nanostructures can biomimic natural ECM of bone tissue in cell culture and as nanocompound tissue promote stem cell differentiation

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into osteogenic lineage, thus regenerating bone tissue (Pan et al. 2017). It was also shown that a novel eri-tasar silk (ET) fibroin when fabricated into nanofibrous scaffold and surface precipitated with nHA is shown to promote osteogenic differentiation without growth factor stimulation. This regenerative capacity is brought about by the enhanced physical chemistry and biological properties of the scaffold. Further, it was shown that an electron spun scaffold composed of biomaterials, nHA, fibrous collagen, and high molecular weight poly-L-lactide (PLLA) can help repair bone in stem cell therapies as they are capable of inducing differentiation of human mesenchymal stem cells (hMSCs) into bone stem cells (Raghavendran et al. 2014). Hu et al. (2014) also added that conventional biphasic calcium phosphate (BCP) ceramics possess properties like conduction and biocompatibility but no osteoinduction. The intrinsic property of osteoinduction of nHA, together with inherent properties of BCP makes them a potential scaffold combination in bone regenerative medicine. nHA coated BCP scaffolds exhibited enhanced adhesion, proliferation, and osteogenic differentiation of MSCs in comparison to traditional BCP ceramics. nHA composites as nano3-D structures promote cell organization, proliferation and transport nutrients to the developing tissues during regeneration. A comparative study work done by Ji et al. (2015) with sphere- and rod-shaped nanoHA-chitosan-gelatin composite 3-D porous scaffolds, showed that sphere shaped nanocomposite scaffold accelerated osteogenic differentiation of fibroblasts derived pluripotent stem cells than rod-shaped scaffold.

15.3.2.2 nHA in Skeletal Defects Restoration Permanent defects in bone can be resolved with artificial bone, which is modified to exhibit exceptional osteoinduction and biomechanical stability. Lin et al. (2013) demonstrated in vivo that bone like substitutes composed of nHA type I collagen beads, platelet rich plasma (PRP), and bone mesenchymal stem cells (BMSCs) in rabbit model enhanced the regeneration of new bone material. Thus, the substitute proved a potential in bone repair or in its regeneration. 15.3.2.3 nHA in Internal Fixation Internal fixation is an orthopedic surgical intervention where bone plates, screws or devices are implanted into fractured bones to repair. A compound incorporated with different polymers like nHA, polyamide 66, and glass fiber is not only biocompatible and exhibited good biomechanical strength but also enhanced MSCs attachment and proliferation. The composite does not negatively effects either mineralization in matrix or MSCs differentiation towards osteogenic lineage (Qiao et al. 2014). 15.3.2.4 nHA in Spinal Fusion Spinal fusion is a procedure usually performed to alleviate the pain caused due to degradation of intervertebral discs. The therapy involving bone graft transplantation has several drawbacks and hence an alternative to it investigated on rabbit was a graft composed of novel mineralized collagen matrix and nHA-collagen-polylactic acid combined with autologous adipose-derived mesenchymal stem cells (ADMSCs).

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ADMSCs on the scaffold proliferated well and posterolateral spinal fusion was witnessed in Rabbit, by the Tang et al. (2012).

15.3.3 Nanostructured Calcium Phosphate (CaP) The CaP biomaterials or scaffolds are biocompatible and are chemically/crystallographically analogous to inorganic component of bone and hence are a subject of interest in bone tissue engineering. CaP biomaterials are more preferred over Silica, Carbon nanotubes, Quantum dots or magnetic particles due to their less toxicity. García-Gareta et al. (2013) demonstrated that when nano-sized CaP coated on biomimetic metallic discs showed greater proliferation of MSCs when compared to uncoated and electrochemically coated discs. Calcium phosphate cement (CPC) as nano-apatitic biomaterial can be easily assimilated in regenerative bone tissue therapy. Xu et al. (2010) demonstrated in vitro that human umbilical cord mesenchymal stem cells (hUCMSCs), cultured on CPC scaffolds proliferated well as it offered good cellular adherence and later in the initiation of differentiation of hUCMSCs towards osteogenic lineage. CPC has high dissolution/resorption capacity, cell adhesion and stem cell delivery potential and hence is a recommended scaffold. Its application is extended to a wide range of load-bearing locations due to its inherent mechanical properties and hence may find its application in maxillofacial and orthopedic operations, if investigated in vivo. CPC alone has low mechanical strength but an additive effect can be noticed with chitosan incorporation into its scaffolds. When MSCs, cultured in vitro on CPC-chitosan scaffold and CPC, differentiation to bone cells occurred in both types but higher alkaline phosphatase activity (ALP), a bone marker, was observed in former compared to latter, conferring high strength to the composite scaffold (Moreau and Xu 2009).

15.3.4 Graphene Nanobiomaterials Graphene is a subject of scientific interest due to its phenomenal properties like mechanical strength, superior electron transport, and high surface area. Due to its presence as thin atomic sheets other polymers can be incorporated into it to enhance their physical properties. Graphene oxide (GO), a derivative of graphene is electrospun with PLGA into porous 3-D scaffolds mats. GO exhibits functions like enhancing hydrophilic performance of scaffold, adhesion, proliferation and inducing differentiation of human mesenchymal stem cells (hMSCs) (Luo et al. 2015a, b). A superior osteogenic differentiation was observed when MSCs cultured on mechanically stiff substrates. Graphene provides such substrate where an enhanced differentiation of hMSCs into bone cells was observed after their proliferation. Graphene-based substrates have good cellular adhesion when compared to ones

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made of silica (Shadjou and Hasanzadeh 2016). Studies were done by Duan et al. (2015) on carbon nanomaterials(CNM), carbon nanotubes (CNTs) and graphene incroprated. The nanofibrous architecture alone provided a good substrate for cell adhesion, proliferation and osteogenic differentiation, incorporation of CNM especially graphene further enhanced the differentiation process in vivo.

15.3.5 Titanium Nanobiomaterials Titanium dioxide (TiO2) nanoparticles are known to enhance cellular migration according to their effective concentration and size (Abou Neel et al. 2008; Hou et al. 2013). They activate macrophages that prompt cell migration to repair/regenerate bone tissues (Chamberlain et al. 2011). Tetra sulfonatophenyl porphyrin (TSPP) in combination with TiO2 nanowhiskers showed good proliferation of bone marrow stem cells and hence are considered potential candidate to treat rheumatoid arthritis (Rehman et al. 2015). A composite nanofibrous scaffold of Zein–polydopamine impregnated with bone morphogenic protein—2 (BMP) peptide conjugated with TiO2 nanoparticles is fabricated in a sustainable way to biomimic the natural topography and produce biochemical cues due to cell-biomaterial interactions that usually occurs in native tissue. The scaffold investigated for its osteogenic differentiation potential has expressed osteogenic markers (human fetal osteoblasts) and proved to be a good biomaterial substrate in bone regeneration applications (Babitha et al. 2018). Kim et al. (2016) demonstrated the potential of another composite scaffold fabricated with silk fibroin into which nanoparticles of TiO2 and HA are incorporated. The scaffold investigated for osteogenic differentiation potential of rat bone mesenchymal stem cells in vitro, indicated with sufficient alkaline phosphatase activity and osteogenic gene expression.

15.3.6 Silica Nanobiomaterials In vitro studies by Tarpani et al. (2016) on nanostructured silica (SiO2) condensed with amine functional group (SiO2-N) and adsorbed with silver nanoparticles (SiO2Ag) to observe the proliferation, survival, and differentiation of human bone marrow mesenchymal stem cells (hBM-MSCs) and adipose stem cells (ASCs) cultured on them have shown them as functional substrates for delivery of molecules to cells. Luo et al. (2015a, b) used a composite mesoporous silica nanoparticle scaffold laden with bone morphogenetic protein (BMP), a transforming growth factor (TGF), which triggered osteogenesis with controlled release of the BMP.

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15.3.7 Bioactive Glass Nanobiomaterials When a trimodal (macro/micro/nano) porous scaffold (TMS) was laden with recombinant human bone morphogenetic protein-2 (rhBMP-2) developed on mesoporous bioactive glass (MBG) and studied in vitro, superior cell attachment, in growth and sustained release of rhBMP-2 prompted osteogenesis was observed. In vivo studies on rabbit also exhibited appealing bone regeneration capacity when rhBMP-2 when used on trimodal scaffold rather than on bimodal (macro/micro) one. Hence clinical application of the substitute is recommended in bone regeneration applications (Tang et al. 2016).

15.4

Nanobiomaterials in Tissue Engineering of Bone Associated Tissues

Tendon-ligament tissue engineering is a must after trauma and multi-walled carbon nanotubes (MWCNTs) when incorporated with polymeric nanofibers are known to enhance tensile strength and biocompatibility by enhancing fibroblasts attachment and hence are considered ideal substitutes in tendon-ligament repair (Sheikh et al. 2015). In cartilage tissue engineering, use of poly glycerol sebacate (PGS) reinforced with nanobioactive glass fiber composite scaffold has shown enhanced tensile strength, tailor the biodegradability of polymer and the bioactivity of the composite is observed when silica is incorporated into glass fiber (Souza et al. 2017).

15.4.1 Craniofacial and Dental Tissue Engineering Tissue engineering in dentistry pertains to three regions, i.e., hard tooth/bone and soft oral tissues. Current research is mainly oriented towards repairing of alveolar bone lost due to periodontal diseases and bone defects in maxillofacial region. Li et al. (2017) emphasized that for craniofacial and dental tissue engineering, nanomaterials like nanoparticles, nanofibers, nanotubes, and nanosheets are commonly used. Nanofibers are porous in nature, simulate the native ECM, aid in cell invasion and their proliferation, hence used as scaffolds in bone, cartilage, and tooth regeneration. Nanotubes and nanoparticles enhance cell attachment, migration, and provide environment viable for tissue regeneration. Nanofibers and nanoparticles even play a pivotal role in controlled delivery system as a bioactive agent carrier in bone and tooth regeneration. The pace at which bioactive agent is released coincides with the matrix degradation thus advancing tissue regeneration procedure. Dental pulp stem cells (DPSCs) located in soft pulp tissue of tooth are highly proliferative and self-renewing cells (Markovic et al. 2010). They are analogous to BM-MSCs sharing a gene profile that initiate mineralization and bone homeostasis (Shi et al. 2001). Galler et al. (2012) addressed pulp injuries with a nanofibrous selfassembling peptide incorporated with DPSCs and diverse growth factors like

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fibroblast growth factor 2 (FGF2), TGF β1, and vascular endothelial growth factor (VEGF). The nanofibrous structure ensures cross-linking which simulates natural ECM and promotes cell to cell and cell to matrix interactions. Encapsulation also allows for a sustainable and controlled release of the factors that regulate angiogenesis and their respective functions. The scaffold when implanted into artificial teeth of mice modal induced angiogenesis in vivo. Periodontal Ligament Stem Cells (PDLSCs), present in periodontium, have the capacity to differentiate into cementoblasts, odontoblasts, and fibroblasts (Seo et al. 2004). BM-MSCs can be used as an alternative to PDLSCs for promoting osteogenesis (Lee et al. 2013). PDLSCs have been demonstrated for bone formation capacity on surface coated with Titanium (Heo et al. 2011) and evaluated human PDLSCs biocompatibility when grown on graphene-based scaffolds. Dental follicle progenitor cells, confined to dental follicle, a tissue surrounding enamel and dental papilla exhibit multi-potency and differentiate into cementoblasts, odontoblasts, and osteoblasts besides being immunomodulatory (Morsczeck et al. 2005).

15.4.1.1 nHA in Dental Restoration Studies by Uezono et al. (2013) established the role of nHA in dental restoration. They demonstrated that by coating nHA and collagen onto titanium (Ti) rods. The composite scaffold when placed under the periosteum of a rat calvaria showed complete formation of new bone surrounding Ti rod after 4 weeks of implantation. nHA when incorporated with poly ethylene glycol (PEG) polymer was found to significantly increase its mechanical strength due to interaction of nHA with polymer. Cell adhesion and spreading was favored due to addition of nHA particles. These bioactive nanomaterial elastomeric composite scaffolds are suitable for use as injectable matrix to promote periodontal tissue regeneration (Gaharwar et al. 2011a, b). 15.4.1.2 Nano-Titanium in Dental Regeneration Ti is commonly used in dental regeneration and load-bearing applications due to its toughness, osteointegration, corrosion resistance, and biocompatibility. Despite formation of an inert layer of TiO2 when in contact with microenvironment in vivo, there are limitations like encapsulation of Ti implants by fibrous tissue, etc., which could be overcome by nanostructured Ti implants which produce osteoinductive signals that facilitate cell adhesion. Nanotopography of the implant influences physicochemical and biochemical properties, promoting ECM formation or mineralization of dental implant surfaces (Mendonca et al. 2008). Jiang et al. (2013) evaluated bimodal (nano/micro structured) Ti implants as clinically usable endoosseous tissue grafts, due to their compound topography which enhances surface area and hydrophilicity thus increasing protein adsorption eventually improving bioactivity of the implant. 15.4.1.3 Synthetic Silicate Nanoparticles in Dentistry Synthetic silicate nanoplatelets are capable of differentiating hMSCs even in the absence of growth factors inducing osteogenesis like BMP-2 or dexamethasone.

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Moreover, its capacity to interact with both natural and synthetic polymers makes it an injectable matrix for cell therapies (Gaharwar et al. 2011a, b, 2014).

15.4.1.4 Graphene in Craniofacial Bone Tissue Engineering Xie et al. (2017) reported that graphene produced by chemical vapor deposition promoted osteogenic and not odontoblastic differentiation when incorporated with DPSCs which was due to downregulation of odontoblast genes and upregulation of osteogenic genes on its surface.

15.4.2 Cartilage Tissue Regeneration (Temporomandibular Joint) As a part of craniofacial bone disorders, i.e., to temporomandibular joint (TMJ) disorder, nanomaterials employed for cartilage regeneration can be used for TMJ disc cartilage regeneration. When electrospun nanofibrous PCL scaffolds incorporated with nHA particles induced differentiation of MSCs and when gold nanoparticles incorporated into porcine collagen hydrogels enhanced mechanical strength of the scaffold and proliferation of chondrocytes (Chiu et al. 2016). Erisken et al. (2011) have demonstrated the ability of human adipose-derived stromal cells to differentiate into chondrogenic lineage. Nanofibrous scaffold composed of poly ɛ-caprolactone seeded with human adipose-derived stromal cells when distributed with insulin and β-glycerophosphate (β-GP) by twin-screw extrusion and electrospinning method, chondrogenic differentiation and β-glycerophosphate (β-GP) induced mineralization was simulated by insulin. Nanostructured PLGA/ TiO2 composite scaffolds exhibited improved cell adhesion of chondrocytes and osteocytes (Kay et al. 2002). Chahine et al. (2014) demonstrated in vitro the biocompatibility of single-walled carbon nanotubes (SWCNTs) in articular cartilage tissue engineering with bioactive functional groups. They noticed enhanced chondrocyte viability and ECM deposition in 3-D porous scaffolds with charged functional groups in comparison to other group tested.

15.5

Nanobiomaterials in Corneal Tissue Engineering

Cornea is the most powerful focusing tissue of the eye which provides 80% of the eyes refractive power. Corneal damage caused due to chemical, physical or by infections leading to diseases like trachoma, onchocerciasis, corneal ulceration, corneal dystrophies, and xerophthalmia cause visual impairments, and may even result in blindness, by affecting cornea, that needs to be corrected (Sommer 1982). From traditional eye medication which implies significant risk to US FDA recently approved collagen cross-linking used to strengthen cornea, several advances took place in corneal research with the aim to develop procedures that are less invasive and restore or preserve vision (Jeng et al. 2016). Corneal tissue is the most common tissue transplanted worldwide. The major limitations of conventional corneal transplantation are donor graft rejection or at times occurrence of or late graft failure and

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limited donor availability. As per World Health Organization (WHO) reports availability of donors is only 15–20% of patients requiring transplantation (Whitcher et al. 2001). It was also noted that if a corneal tissue is transplanted and is rejected immunologically, then re-transplantation of the same is risky and painful (Panda et al. 2007). Keratoprosthesis, a procedure to replace the diseased cornea with artificial cornea, pioneered by Mannis and Mannis (1999) do not opacify and are used for the purpose of transplantation. Although it is clinically successful host rejection is relatively high in these transplants. Also, they have other disadvantages which include complications such as infection, extrusion, glaucoma, and retinal detachment. Corneal tissue engineering overcome these limitations and provides the best platform as an alternative to cadaveric corneal transplantation to replace either an individual layer of epithelium, stroma, or endothelium or regenerate the entire tissue and restore its functional properties (Hong Sheng 2015). Ocular surface regeneration, as indicated in several articles, is the bioengineering of the outer most layer of eye and utmost care has to take to reconstruct it with good optical and biomechanical properties. Cornea has to withstand an intraocular pressure of 10–20 mm Hg (Mobaraki et al. 2019). The tissue contains stratified lamellae that exhibits topographical features at nano and microlevel. Therefore, any engineered product should be in compliance with those features. There are four main types of approaches in corneal tissue engineering, i.e., full-thickness, stroma, epithelial, and endothelial regeneration. Hong Sheng (2015) emphasized that tissue engineering of either epithelium, stroma, or endothelium alone can regenerate cornea restoring its physiological activity. Nevertheless, all those approaches require scaffolds with different cells incorporated into it. Ruberti and Zieske (2008) stated that cells, scaffolds, and biomolecules are used to produce corneal substitutes that are used in regeneration or repair of injuries or corneal diseases. Tissue engineering of cornea is less complicated as it is mostly avascular and exposure to immune cells is a rare phenomenon. However, there exists contrasting features which are to be focused while engineering corneal structures like transparency, protection, and substantial refractive power when compared to native tissue (Ghezzi et al. 2014). Nanostructured scaffolds of natural or synthetic origin or in combination simulated native ECM in 3D corneal tissue formation and novel cell implantation techniques to facilitate proper alignment lead this technique to replace corneal transplantation (Fu et al. 2010; Zhang et al. 2015a, b). Cornea is known for its mechanical strength and transparency and restoring these properties, besides refractive power is crucial in its reconstruction. Inflammatory reactions, neovascularization, and limbal cell deficiency determine corneal clarity. Immune privilege, i.e., limitation and modulation in immune response, produces little inflammatory reactions during a corneal injury preventing angiogenesis, a process that results in neovascularization (Mobaraki et al. 2019). Another limitation is to transfer cells of interest to specific layers of cornea, their viability and preservation of its function which can be overcome by usage of scaffolds that are key in supporting attachment and in the survival, proliferation, and alignment of transplanted corneal cells. Scaffolds need to be mechanically strong and transparent

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in long-run along with conventional properties like biodegradability, biocompatibility, and good topography (Carlsson et al. 2003). Work of Zarrintaj et al. (2017) on stimulating the natural body repair mechanisms to restore functional healing of damaged tissues and organs with an approach integrating corneal transplantation and tissue engineering is now considered a gold standard for advanced corneal tissue engineering/transplantation. The existence of stem cells in limbus epithelium, present between the cornea and the sclera which have the capacity to regenerate cornea and are anticipated to provide fine healing and integration provides the basis towards working on this approach. As such is the scenario, researcher is being focused with the aim of producing a greater number of novel biomaterials that simulate the topography and morphology of corneal tissue and aid in its repair.

15.5.1 Natural Polymers Collagen, chitosan, gelatin, hyaluronic acid, silk fibroin, polyarginine, and hydroxyapatite are used in combination with diverse synthetic polymers to produce varied scaffolds and nanobiomaterials for applications in corneal tissue engineering.

15.5.2 Synthetic Polymers PGS (Polyglycerol sebacate), PCL (Poly ε-caprolactone), Polyglycolic acid, Poly ester (urethane) urea, and surface modified PCL are some examples of synthetic polymers used in corneal tissue engineering. Studies by Reimondez-Troitiño et al. (2016) on Polyarginine, a cationic peptide, as nanocarrier aids in wound healing by interfering with the transforming growth factor beta/SMAD (TGF-β/SMAD) signaling pathway. A biodegradable nanopolymer, alpha 1–6 polycarboxymethyl-sulfate which is also a pharmaceutical agent, was investigated as a biomimetic polymer, mimicking heparan sulfate, a cell signaling protein of the extracellular matrix (ECM), using epi-off corneal collagen cross-linking technique was found to accelerate the process of corneal re-epithelization and stromal healing (Gumus et al. 2017).

15.5.3 Nanobiomaterials in Corneal Epithelial Tissue Engineering Ma et al. (2006) investigated the potential of mesenchymal stem cells (MSCs) in repair of corneal tissue formation. They transplanted human MSCs into rats and observed the expression of epithelial cell markers, thus proved the differentiation of MSCs into corneal epithelial cells on human amniotic membranes. Studies on acellular porcine cornea on which amniotic epithelial cells were grown found to develop strata of epithelial cells that is used in lamellar keratoplasty which readily integrated with host cornea (Luo et al. 2013). Application of amniotic membranes in corneal regeneration is limited due to its different cell morphology compared to

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corneal tissue even though it has high anti-angiogenic and anti-inflammatory properties (Connon et al. 2010). To overcome the various limitations of conventional scaffold materials, nanobiomaterials like PGS, PGA in combination with collagen, chitosan, etc., are used in various studies (Lawrence et al. 2009; Nakajima et al. 2015).

15.5.4 Nanobiomaterials in Corneal Endothelial Tissue Engineering Elastomeric biodegradable nanofibrous scaffold composed of PGS (Polyglycerol sebacate) -PCL (Poly ε-caprolactone) that are semi-transparent is used to engineer endothelial tissue (Salehi et al. 2017). Rizwan et al. (2017) cited that sequential hybrid cross-linking of nano-gelatin methacrylate as hydrogel in tissue repair is a good nutrient conduit, provides mechanical strength besides being transparent.

15.5.5 Nanobiomaterials in Corneal Stroma Tissue Engineering One of the resident cells of cornea is keratocytes that are present in stroma. Stroma of cornea is complex and unique affecting the mechanical strength and transparency of this tissue. Hence a wide range of nanomaterials are studied for the purpose to develop engineered corneal stroma (Torbet et al. 2007; Lawrence et al. 2009; Mimura et al. 2008). Gelatin hydrogels are used to engineer corneal stroma and introduce into rabbit stroma for therapeutic purposes (Mimura et al. 2008). Amphiphilic self-assembling nanopeptides, laminin, and fibronectin derived substances can be integrated into nanofibrous scaffold, which acts as a substrate for corneal replacement. In cell culture on laminin derived nanofibrous, growth and differentiation of human keratocytes was observed in vitro. The same was observed in vivo in rabbit model (Uzunalli et al. 2014). Poly ester (urethane) as a nanofibrous scaffold supported differentiation of human corneal stem cells into stromal cell lineage (Wu et al. 2013).

15.5.6 Cell Sheet Engineering in Corneal Tissue Engineering Koizumi and Okumura (2019) reviewed that conventional tissue engineering using bio scaffolds has been successful in various tissues for over three decades. Nevertheless, in reconstruction of an ocular surface there are certain limitations that can be corrected by cell sheet engineering. There is a high risk of infection as biomaterials are employed and improper attachment to replacement site results in degradation and loss of transparency, an important factor in corneal construction. Cell sheet assemblies can resolve those difficulties (Fagerholm et al. 2014). Desired cells are grown on specific cell sheets that avoid cell adhesion and it is transplanted to the target site as a whole entity without using any scaffolds. The technology utilizes external stimuli like specific enzymes, temperature variation, magnetic force,

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electrochemical polarization, pH variation, polyelectrolytes, and illumination to promote detachment of cells from surface (Owaki et al. 2014). Recent developments in the technology with advances in fabrication of sheets to produce cell sheet assemblies of good quality and achieve replacements with minimal biological contamination includes “Cell cartridges,” a novel closed culture system, devised to regenerate corneal epithelial cell sheets (Kobayashi et al. 2013; Nakajima et al. 2015). Another latest technology is self-lifting analogous tissue equivalent (SLATE), a bio-fabricated, scaffold-free system that engages peptide amphiphile as surface template to regulate its morphology and physiology and eventually aids in repairing corneal tissue. Studies on rabbit model implanted with SLATE have exhibited no rejection and complete integration into native tissue by 9 months (Gouveia et al. 2017).

15.6

Limitations and Future Prospects

There is tremendous scope to develop different nano biomaterials for use in various applications in TE and RM. Since use of nanobiomaterials is in its early stages and a lot of research is being focused to develop scaffolds that can biomimic, nanobiomaterials took the center stage due to their properties which gives them edge over normal biomaterials. Advances in nanobiomaterials have paved the way for new multidisciplinary research areas in biology, medicine, material sciences, and nanotechnology. This led to reshaping the ideas of application of materials in biology and medicine and furthering our understanding of the interactions of materials with biological systems. But our current understanding of nanobiomaterials, their efficiency and biocompatibility, applications are limited given the large spectrum of nanobiomaterials being synthesized by different techniques with variation in their properties with respect to evaluating their use in biological systems, especially TE & RM. As the field of nanobiomaterials is entering into new era of smart and intelligent biomaterials a better understanding of their properties and interactions with biological systems is of utmost priority for their proper and efficient usage in different aspects of TE & RM. Both in vitro and in vivo studies are to be carried out before their clinical usage to evaluate their safety and toxicity. A proper investigation into their interactions within the systems and the environmental aspects needs to be carried out. Future scope in nanobiomaterials relays in the development and use of novel fabrication techniques for developing of smart and intelligent biomaterials which are simulated to perform specific functions at desired micro-environmental conditions. This can pave the way for automatic simulative response nanobiomaterials that can be used to biomimitic the microenvironment of cells and tissues. Another aspect where focus should be made is on developing integrated nanobiomaterials with electronic devices like nanochips, etc., which will open up the scope of the research to new avenues. Although some work has been done in this area, it is still at its infancy. Development of multifunctional nanobiomaterials is another area where much emphasis is to be done since realizing this means creating the path for

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immunocompatible, bioactive, system deliverable nanobiomaterials. Last but not least is the aspect of CAD, fabrication and simulation of nanobiomaterials that needs to be focused since it will not only save time and money but also helps in development of precise novel nanobiomaterials.

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Intelligent Biomaterials for Tissue Engineering and Biomedical Applications: Current Landscape and Future Prospects

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M. S. Anju, Deepa K. Raj, Bernadette K. Madathil, Naresh Kasoju, and P. R. Anil Kumar

Abstract

Intelligent biomaterials continue to find emerging applications in the field of medical devices, drug delivery, tissue engineering and beyond. A major approach to design such advanced materials is to incorporate specific functional groups into the molecular structure of a biomaterial that can respond and be modulated by appropriate alternations in the microenvironment. Typically, these intelligent materials change their morphological, physical or chemical properties in a controlled manner in response to external stimuli that are either physical such as light, temperature, etc., chemical such as pH, ionic strength, etc. or biological such as enzyme, stress, etc. In this chapter, we give a brief historical account of intelligent biomaterials, followed by basic concepts and current trends in the field of shape changing materials, thermoresponsive, light-sensing, pH-reactive, magnetostimuli and bio-active biomaterials. Further, we also detail some other unique and emerging intelligent biomaterials under development including ultrasoundreactive, dual or multi-responsive biomaterials. Research and development activities in the field of intelligent biomaterials have intensified in recent years and we envisage a greater scope for these biomaterial systems in tissue engineering and beyond.

M. S. Anju, Deepa K. Raj and Bernadette K. Madathil contributed equally. M. S. Anju · D. K. Raj · N. Kasoju (*) · P. R. Anil Kumar (*) Division of Tissue Culture, Department of Applied Biology, Biomedical Technology Wing, Sree Chitra Tirunal Institute for Medical Sciences and Technology, Thiruvananthapuram, Kerala, India e-mail: [email protected]; [email protected] B. K. Madathil Advanced Centre for Tissue Engineering, Department of Biochemistry, University of Kerala, Thiruvananthapuram, Kerala, India # The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 B. Bhaskar et al. (eds.), Biomaterials in Tissue Engineering and Regenerative Medicine, https://doi.org/10.1007/978-981-16-0002-9_16

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Keywords

Smart materials · Phase transformation · Shape memory effect · Medical devices · Tissue engineering · Drug delivery · Biosensing · Diagnostics

Abbreviations aPLA-co-TMC bFGF BIB CHC ECFCs ECM GelMA hTMSCs iPSCs LCST MeHA MNPs NGMA NIR PCL PDLLA PEG PEI PNIPAAm PPS ROS SiRNA SMMs SMPs UCST

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acrylated poly-L-lactide-co-trimethylene carbonate basic Fibroblast Growth Factor 1,4-Bis(imidazol-1-ylmethyl) Benzene Carboxymethyl-Hexanoyl Chitosan Endothelial colony forming cells Extra cellular matrix Methacrylated gelatin human Tonsil-derived stem cells induced Pluripotent stem cells Lower critical solution temperature Methacrylated-hyaluronic acid Magnetic nanoparticles Poly N-Isopropylacrylamide-co-Glycidylmethacrylate Near infrared Polycaprolactone Poly(ethyleneglycol)-Poly(D,L-lactide) Polyethylene glycol Polyethylenimine Poly(N-isopropylacrylamide) Poly(Propylene Sulphide) Reactive oxygen species short interfering RNA Shape memory materials Shape memory polymers Upper critical solution temperature

Introduction

In the year 1976, first Consensus Conference of European Society for Biomaterials (ESB) defined a biomaterial as ‘a nonviable material used in a medical device, intended to interact with biological systems’. Over the decades advances in materials science, microfabrication technologies as well as biomedical engineering have led to the development of next generation biomaterials that are capable of modulating their

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Fig. 16.1 Schematic of intelligent biomaterials: Various biomaterials are being developed that would act intelligently towards a particular condition and elicit a desirable response

properties and functions in response to exogenous stimuli (Kowalski et al. 2018). This new class of materials, termed as intelligent biomaterials, emerged with the discovery of the shape memory effect in AuCd alloy in the 1930s and over the last few decades has evolved rapidly and found numerous application in various biomedical fields such as medical devices, tissue engineering, drug delivery, biosensing, bioimaging, immune engineering and 4D printing. These materials are made from metals, ceramics, polymers and composites; though, a majority of them are polymer based since polymers are relatively easy to process into various forms such as gels, films, sponges, etc. Intelligent biomaterial systems usually involve one or more moieties that are integrated within the molecular structure and respond to the stimuli, either in a reversible or irreversible manner. These materials can be made to sense, react or adapt to various stimuli (Fig. 16.1) that are broadly categorized into (a) physical stimuli such as light, temperature, magnetism, ultrasound, electricity, etc. (Ju et al. 2009a), (b) chemical stimuli such as solvents, pH, ionic strength, etc. (Ju et al. 2009b) and (c) biological stimuli such as enzymes, glucose level, hypoxia, oxidative stress, etc. (Lu et al. 2017). Multifunctional or multi-reactive biomaterial systems are also emerging in recent years in order to achieve the desired function in a particular physiological or pathological conditions (Keane and Badylak 2014). In this chapter, we begin with a historical account and progressive evolution of intelligent biomaterials from shape memory metallic alloys to multi-stimuli biomaterial systems (Sect. 16.2). This is followed by Sect. 16.3 which details about shape memory effect in metal alloys and subsequent innovations in developing shape memory polymers. In Sect. 16.4, we describe about biomaterials that respond to temperature, with an emphasis on Poly N-Isopropylacrylamide (PNIPAAm) and their applications in cell and tissue engineering. Subsequently, Sect. 16.5 details the biomaterials that respond to light across various wavelengths of the

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electromagnetic spectrum, with a focus on much-acclaimed Methacrylated gelatin (GelMA). Section 16.6 gives insights into pH-responsive biomaterials and their application in drug delivery and tissue engineering. Section 16.7 details about magneto-sensing biomaterials that are typically synthesized by the incorporation of polymer materials with magnetic nanoparticles. Section 16.8 explains about the electroactive biomaterials that includes polymers, electrets, piezoelectric and photovoltaic materials. In Sect. 16.9, we describe about biomaterials that react to physiological and pathological conditions in the body, with a focus on enzyme-reactive, oxidative stress-reactive and immune-responsive biomaterials. In the following Sect. 16.10, we describe about the other unique stimuli-responsive materials such as ultrasound responsive and emerging dual/multi-responsive biomaterials. We conclude this chapter with Sect. 16.11 describing a brief summary and a note on future prospects.

16.2

Historical Account of Intelligent Biomaterials

The history of intelligent biomaterials commenced as early as 1932 with the discovery of rubber like elastic behaviour of AuCd alloy at room temperature. The shape memory effect of metallic alloy was started as a new scientific area by Chang and Read in 1951. Subsequently pseudo-elastic property of Cu-Zn alloy was reported by Reynold and Bever in 1952. A breakthrough in the field of intelligent biomaterials came in 1963 with the discovery of NiTi alloy by William J. Buehler. Since the discovery of NiTi alloy, and particularly in the early 1970s, many studies have reported on the application of NiTi in medical applications (Cutright et al. 1973; Iwabuchi et al. 1975). In the early 1990s first commercial stent was developed using NiTi, resulting in a big leap in the field of cardiovascular biology and radiology (Ryklina et al. 1996). Reports on the shape memory polymers have emerged in the year 1980s. One of the early publication mentioning ‘shape-memory’ effects in polymers was reported by L. B. Vernon claiming a dental material (methacrylic based) with elastic memory that resumes to its original shape upon heating (Vernon and Vernon 1941). In the following years, heat shrinkable polyethylene tubing and films were also reported for potential applications in biomedical field (Rainer et al. 1964). The next generation intelligent biomaterials have emerged from 1990s onwards with the reports of multi-stimuli materials. The thermoplastic polyurethane shape memory polymer originally invented by Dr. S. Hayashi in 1990s was reported to be both thermoresponsive and moisture responsive. Additionally, the shape memory hybrid composites provide a new paradigm in the material–structure interactions. Carbon nanotubes incorporated polyurethane materials were found suitable for potential application in small load baring devices like tracheal and laryngeal stent (Bellin et al. 2006). In the following years, triple or multi-shape memory or stimuli biomaterials were reported (Bellin et al. 2006). Two-way shape memory semicrystalline network polymer enabled the development of costly shape memory alloys and liquid crystalline elastomers (Chung et al. 2008). Recently, the application of 3D

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Fig. 16.2 Timeline of events in intelligent biomaterial systems: With early reports on pseudoelastic nature of Au-Cd alloy in 1930s, the field of intelligent biomaterials has grown over the years, and they are now widely explored in various biomedical applications

printing to create multi-shape memory polymeric composite was reported and thereby further expanding the horizons of intelligent biomaterials in new generation biomedical engineering applications (Wu et al. 2016). Developments on selfdeployable devices are gaining momentum with introduction of materials by copolymerization of 2-methoxyethyl acrylate and N-methylol acrylamide (Liu et al. 2012). Overall, the field of intelligent biomaterial systems have evolved quite extensively over the last few decades (Fig. 16.2), while some of them are already used in clinical setups, several others are under development.

16.3

Shape Changing Materials

The shape memory effect was first discovered by Chang and Read (1951) in the year 1932 in gold–cadmium alloys (Chang and Read 1951). Shape memory materials (SMMs) can be categorized as a group of intelligent materials with mechanical, functional, or simply, materials with an ability to change shape/form. SMMs are generally composed of metals, polymer and composites. For a metal to exhibit shape memory, it requires thermo-elastic martensitic phase transformation. Martensite phase of a material is formed through diffusion-less transformation, i.e. phase transformation from one to another through subtle atomic rearrangements rather than diffusion. Over the years, the shape memory metallic alloys have found its applications in biomedical engineering. One of the popular SMMs is equiatomic NiTi, also known as nitinol, first reported by Buehler et al. (Buehler et al. 1963). Because of the shape memory and super elastic property, nitinol—that is composed of nearly equal amount of nickel and titanium—has been widely explored as a biomaterial as compared to other conventional materials. Nitinol has great advantage in biomedical application since the transformation occurs close to body temperature,

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thermo-elastic martensitic phase and reverse transformation to parent austenite behaviour upon heating (Yahia et al. 2009). In the context of tissue engineering, nitinol is widely used in reinforcing bone owing to the low elastic modulus, high tensile strength and high compressive strength close to bone material. Besides, it has many applications in the medical field as guided wire, heart valve tool, joining fractured bone (Niinomi 2003). Polymers that are capable of similar shape memory effect in a predefined manner in response to appropriate stimulus can be classified as shape memory polymers (SMPs). The SMPs mostly rely on the molecular architecture and do not require specific chemical structure thereby make it possible to tune the shape by adjusting the ratio of monomer to co-monomer. Shape memory is not an intrinsic property and polymers do not exhibit this by themselves. It requires special processing to acquire phase change properties that sense and respond to a stimulus. For instance, the processing can either be heating up of sample or cooling at very low temperature across a transition temperature Ttrans. Most of the SMPs are with dual shape effect which may be temporary or permanent. For dual shape memory effect the polymer requires chemical and/or physical cross-links (crystallites for semi-crystalline polymers and entanglements for amorphous polymers) to memorize the initial shape (Ortega et al. 2012). For instance, thermally activated poly(norbornene)based SMPs, discovered by Echigo et al. (Echigo et al. 1990), were succesfully used to occlude cardiac ductus arteriosus. Bio-derived SMPs are widely used in biomedical applications, including drug delivery carriers, self-tightening sutures, fasteners, drug eluting stents, aneurysm occlusion, endovascular clot removal and orthodontic appliances (Yang et al. 2013). For instance, Chen et al. 2009 developed a drug-loaded chitosan stent that was made to crimp to allow easy deployment into the artery (Fig. 16.3).

16.4

Thermoresponsive Biomaterials

One of the widely explored intelligent biomaterials are those that sense and respond to heat. These materials, termed as thermoresponsive, exhibit a phase change at a particular temperature. Thermoresponsive polymers are classified into two types, namely those that follow lower critical solution temperature principle (LCST) or those that follow upper critical solution temperature principle (UCST). LCST polymers become insoluble above a critical temperature and the other will precipitate below the critical temperature. LCST is an entropically driven effect while UCST is an enthalpically driven effect (Ward and Georgiou 2011). Studies have shown that both hydrogen bonding and hydrophobic interactions in the polymer solvent system are responsible for the phase transition (Teotia et al. 2015). Widely used thermoresponsive polymers include poly(N-isopropylacrylamide) (PNIPAAm) with an LCST of around 32  C, poly(N,N-diethylacrylamide) with an LCST of 25–32  C, poly(N-vinylcaprolactam) with an LCST of 25–35  C, poly [2-(dimethylamino)ethyl methacrylate with an LCST of around 50  C. Owing to ethylene oxide units in a chain, size of the side chains, copolymerization partner and

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Fig. 16.3 Schematic illustration showing shape memory cycle and shape memory effect in a SMP stent: Sirolimus-eluting chitosan based stent was made to crimp so as to suit deployment via a French sheath in an artery and then to expand with hydration upon deployment. Reproduced with permission from (Chen et al. 2009). Copyright # 2008 Elsevier Ltd

molecular weight, PEG shows LCSTs in a wide range of temperatures from 20 to 85  C (Kim and Matsunaga 2017). Of these PNIPAAm is the most exhaustively studied thermoresponsive polymer with potential applications in biomedical research owing to its thermoresponsiveness between room temperature and physiological temperatures. It has been employed to design a variety of smart biomaterials for a wide spectrum of applications in advanced biomedical applications including drug delivery, cell sheet engineering and micro-carriers for large scale cellular proliferation. The most successful application of PNIPAAm is in the technique of cell sheet engineering (Fig. 16.4). Here PNIPAAm is grafted onto tissue culture polystyrene dishes to generate thermoresponsive substrates. By variation of temperature, cell adhesion and detachment can be controlled as the hydrophilic/hydrophobic nature and conformation of the polymer changes at temperatures below and above its LCST. The technique avoids the use of proteolytic enzymes, like trypsin, or mechanical procedures, like scraping. Cell viability and re-adhesion of cells can be enhanced while growing in the polymer because ECM of the cell layer stays in place around the cells. The cell sheet can be transplanted to the diseased site and has the advantage of being suture-less and scaffold free, hence reducing the chances of inflammation and immune rejection (Tekin et al. 2011). Cell sheets have been generated for engineering various tissues such as cardiac, hepatic, skin, periodontal area and cornea. Different cells can be co-cultured together in a single cell sheet. Individual cell sheets can be stacked to form thicker functional tissues. Poly NIsopropylacrylamide-co-glycidyl methacrylate (NGMA) with an LCST of about

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Fig. 16.4 Schematic of thermoresponsive biomaterials in various biomedical applications: PNIPAAm based intelligent substrates were successfully used for cell sheet engineering, large scale cell expansion and bioseparation. Reproduced from (Nagase et al. 2018). # 2017 The Authors. Licensee Elsevier Ltd

31  C has been developed in our lab by copolymerizing glycidyl methacrylate with NIPAAm using the technique of free radical polymerization. These NGMA substrates have been used to generate intact and viable rabbit native as well as stem cell derived corneal epithelial, stromal and endothelial cell sheets (Nithya et al. 2011; Madathil et al. 2014; Ramesh et al. 2014; Raj et al. 2018; Venugopal et al. 2020). We also explored its use in the myocardial cell sheet engineering for potential applications in cardiac tissue engineering (Nair and Kumary 2016). Ours as well as various other studies highlight the potential of thermoresponsive polymers in engineering cell sheets as an advanced cellular therapy strategy (Ebara et al. 2004; Arisaka et al. 2016). Thermoresponsive polymers are often used as in situ injectable hydrogels, nanogels or micro-sized bulk hydrogels. They find widespread biomedical applications as they are in the liquid state outside the body, but on injection or application to a part of the human body, i.e., at 370c the increase in temperature results in formation of a gel at the desired site (Fitzpatrick et al. 2010). Thermoresponsive injectable carboxymethyl-hexanoyl chitosan (CHC) nanoscale hydrogel is observed to increase the viability and CD44+ proportion of induced pluripotent stem cells (iPSCs). The delivery of iPSC/CHC hydrogel enhanced corneal reconstruction by down regulating oxidative stress and recruiting endogenous epithelial cells to restore corneal epithelial thickness in alkaline induced corneal injury (Chien et al. 2012). Stimuli-responsive porogens, such as gelatin, alginate and hyaluronic acid, are used to create hydrogels with tunable porosity. Such microporous hydrogels provide for efficient transfer of nutrients, ECM and promote cell proliferation, migration, cellular interactions and vascularization (Han et al. 2013). For instance, Lee et al. have reported the use of solvent-spun PNIPAM microfibers as temperature-responsive templates in gelatin hydrogels encapsulating human

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neonatal dermal fibroblasts. To generate capillary-like micro-channel vascular, the PNIPAM microfibers are removed by immersion in cell culture medium at room temperature (Lee et al. 2016). Another widely explored field of thermoresponsive polymers is micro-carriers based cell expansion, wherein cells such as stem cells are cultured on these intelligent micro-carriers and are subsequently harvested without using any enzymes (Tamura et al. 2012). The field of thermoresponsive polymers is still evolving and is now finding relationship with advanced technologies such as microfluidics and 3D printing to further its applications in biomedical field (Kim et al. 2015; Celikkin et al. 2018).

16.5

Photoresponsive Biomaterials

Intelligent biomaterials that use light as a trigger are highly attractive as the light can be exposed remotely, with a control over its wavelength, intensity and area of focus. Typically, light sensitive intelligent biomaterials are made by incorporation of photoreceptive moieties into the molecular structure of the material. The most commonly used photo liable groups are o-Nitrobenzyl ester, Coumarin, Thiol-ene, Azobenzene, Spiropyran, Trithiocarbonate (Li et al. 2019). The most common photochemical reactions that are employed to create sophisticated biomaterial scaffolds are photo-initiated chain polymerization, thiol-ene and thiol-yne photoclick chemistry, sequential crosslinking strategies, photo-cleavage reactions and photo-caging reactions (Brown and Anseth 2017). Typically, the photochromic molecule initially captures the optical signal and converts it into a chemical signal through a photoreaction involving isomerization (cis-trans, open-close), cleavage or dimerization, possibly in reversible manner (Mantha et al. 2019). Subsequently, this will lead to light-induced transition of hydrophobicity-hydrophilicity in a biomaterial, cleavage and release of biomolecules bound to a material surface, crosslinking of biomaterials to yield a functionalized surface or in situ gelling matrices, or lightinduced localized heating (Indermun et al. 2018). Light across various wavelengths of the electromagnetic spectrum has been explored including UV range across 250–380 nm and near infrared (NIR) range from 700 to 900 nm. In certain cases, owing to the photo-toxicity associated with the light in these ranges, biomaterials that respond to light within visible range are also being developed. One of the highly popular photo-reactive polymers is methacrylated gelatin, widely abbreviated as GelMA, which is a UV-crosslinkable polymer that is synthesized by adding methacrylate groups to the amine-containing side-groups of gelatin. GelMA has profound use in cell therapy and precision medicine because of its convenient synthesis, biocompatibility and 3D printability. Since its first mention by Van Den Bulcke, GelMA has been at forefront of photo-reactive biomaterial research and therefore an array of reports described various modes of GelMA synthesis and functionalization in order to achieve control in physico-chemical and mechanical properties suitable for cell and tissue engineering applications (Sun et al. 2018). For instance, Lin et al. reported on the use of cell laden injectable GelMA that can be photo-cross-linked in vivo in an immune-deficient mice using transdermal

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Fig. 16.5 A representative example of photo-reactive biomaterial: GelMA-based hydrogel was used to encapsulate ECFCs and MSCs by either (a) ex vivo and (b) in vivo photo-crosslinking approach (c—cell viability, d—tubule formation imaging and e—mouse receiving UV light and vascularized graft after day 7). Reproduced with permission from (Lin et al. 2013). Copyright # 2013 Elsevier Ltd

exposure to UV light. It was reported that within 7 days the human blood derived endothelial colony forming cells (ECFCs) and bone marrow-derived MSCs in the GelMA developed vascular networks throughout the hydrogel and formed functional anastomoses with the host vasculature (Fig. 16.5) (Lin et al. 2013). In order to impart electrical conductivity, UV-cross-linkable gold nanorod-incorporated GelMA hybrid hydrogels have been proposed for cardiac tissue engineering. The hydrogel can be cross-linked to modulate mechanical stiffness while the gold nano rods promote electrical conductivity. The hybrid hydrogel proved to be an efficient substrate for the culture of cardiomyocytes in vitro and hence finds potential as functional cardiac patches in regenerative medicine (Navaei et al. 2016). Kwon et al. have also reported on the osteogenic potential of nanobioglass embedded GelMAbased cryogels. In vitro this hydrogel induced osteogenic differentiation of human tonsil-derived stem cells (hTMSCs), further, in vivo studies using calvarial defect model in mice have shown that the bioglass embedded GelMA cryogel laden with hTMSCs had promoted bone regeneration of the defected area. The cryogel exhibited osteoinductive properties and was bio-active as evinced by the formation of hydroxyapatite on its surface (Kwon et al. 2018). Various other photo-reactive biomaterial systems are under investigation for potential applications in drug delivery and tissue engineering. Lin et al. described methacrylated-hyaluronic acid (MeHA) hydrogels to be ideal 3D scaffolds for

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proving synergistic effects of chondrogenic preconditioning and mechanical loading on the chondrogenic differentiation of encapsulated bone marrow-derived MSC. These preconditioned cell laden photo-polymerized hydrogels showed improved chondrogenesis and neocartilage formation in rat osteochondral defects (Lin et al. 2017). Sun et al. have developed a poly(ethylene glycol)-poly(D,L-lactide) (PDLLAPEG 1000) hydrogel which used the visible-light sensitive initiator lithium phenyl2,4,6-trimethylbenzoylphosphinate for crosslinking. They reported a biomaterial, PDLLA-PEG 1000 having mechanical properties in the range of native articular cartilage. Within the PDLLA-PEG 1000 hydrogel in vitro studies were carried out which showed both high cell viability and chondrogenic potential of seeded human MSCs (Sun et al. 2017). Stefani and Cooper-White had reported the fabrication of composite tubular polymer scaffolds made from polycaprolactone (PCL) and acrylated L-lactide-co-trimethylene carbonate (aPLA-co-TMC) by electrospinning. The addition of acrylate groups facilitated UV crosslinking of the aPLA-co-TMC chains during the electrospinning process. In vitro studies using human stem cells showed the fibres to be cytocompatible and the cells to be aligned with the fibres. These multicomponent, elastomeric electrospun polymer scaffolds thus have potential to be used in vascular tissue engineering applications (Stefani and Cooper-White 2016). Otsuka and Barrett had used the technique of electrospinning to fabricate nano/macro-fibres of a photo-responsive cellulose Avicel. The cellulose derivative is made photo-responsive by functionalizing with a reversible molecular photo-switch, azobenzene (Otsuka and Barrett 2019). Recently, the field of photo-reactive biomaterials have received tremendous attention owing to its importance in 3D bioprinting applications. Several photo-polymerization chemistries are currently under investigation and therefore we envisage further development of this field in near the future (Bagheri and Jin 2019).

16.6

pH-Responsive Biomaterials

pH is one of the critical physiological parameters that defines or indicates a particular condition in cell and tissue microenvironment. The pH value varies at different sites in the human body from being highly acidic (pH 1–5) in the gastric fluids to being alkaline (pH 5–8) in the intestines. Tumours are also reported to have acidic microenvironment due to their high metabolic and cellular proliferation rate while chronic wounds exhibit pH in the range of 7.4–5.4 (Ju et al. 2009a, b). These pH variations form the basis for the development of pH-responsive biomaterials. Typically, a pH-responsive moiety may be incorporated into the polymer structure or chemical conjugation of pH-liable linkage between polymers and drugs may be used or destabilizing a self-assembled polymeric aggregate with pH change may be adopted for designing such polymers (Gil and Hudson 2004). pH-responsive biomaterials are mainly polymers that have weak acidic or basic groups that gain or release protons in response to changes in the environmental pH leading to alterations in their physical properties like solubility, swelling and chain conformation. Changes in the environment such as pH, ionic strength and type of counter-ions

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Fig. 16.6 pH-sensitive biomaterial in action: pH-sensitive liposomes modified with poly(ethylene glycol) were developed for intracellular delivery of bleomycin for cancer chemotherapy. Reproduced from (Yuba et al. 2018). # 2018 The authors. Licensee MDPI, Basel, Switzerland

modulate the electrostatic repulsion and effect the transition between a tightly coiled and expanded conformation of the polymer. At low pH polyacid polymers will be collapsed as the acidic groups will be protonated and unionized. Negatively charged polymer will swell when the pH range shifts from acidic to basic. Polybasic polymer swells when the pH range shifts from basic to acidic because the ionization of the basic groups will increase when pH decreases (Reyes-Ortega 2014). Smart biomaterials can be modulated by changes in pH have found significant applications in the fields of drug delivery and tissue engineering (Fig. 16.6). pH-responsive materials were typically made to contain cleavable bonds such as hydrazine, imine, oxime, amide, ketal, orthoester and phenylvinylether (Zhang et al. 2019). The pH-responsive polymers used for biomedical applications include natural polymers like chitosan, albumin, gelatin, alginate and synthetic polymers like poly (acrylic acid), polyacrylamide, poly(methacrylic acid), poly(diethylaminoethyl methacrylate), poly(dimethylaminoethyl methacrylate) and their copolymers (Mantha et al. 2019). pH-responsive hydrogels as non-viral gene carriers have been extensively researched over the last decade. Gene therapy as a treatment mode is directed at introducing a new gene, replacing a mutated gene or knocking out a mutated gene. Non-viral gene delivery is purported to overcome the safety issues of viral vectors. Here nucleic acids are complexed with either cationic liposomes to form lipoplexes or with cationic polymers to form polyplexes (Mathew et al. 2017). For instance, Guan et al., had complexed plasmid DNA with

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polyethylenimine (PEI)—poly(L-glutamate) (PLG) which was then tightened with PEG. Between the aldehyde group of PEG and the amino group of PEI a schiff base is formed which was stable at physiological pH (7.4), but cleavable in acidic pH, which can lead to shedding of the PEG shielding there by facilitating gene delivery. This system of PEG and PEI showed high transfection efficiency in mouse colon carcinoma CT-26 cell lines (Guan et al. 2016). Similarly, Cao et al. developed a pH-responsive micelle based on polycaprolactone-block-poly 2-(dimethylamino)ethyl methacrylate cationic copolymer to deliver short interfering RNA (siRNA) for silencing interleukin 8 (IL-8) gene in hepatoma cancer cells SK-Hep1 (Cao et al. 2019). pH-sensitive biomaterials have also been extensively used in the field of cell and tissue engineering. For instance, in the case of acute myocardial infarction, the infarcted area has a lower pH than normal cardiac tissues because of the inflammation generated. Therefore, a hydrogel that is responsive to the pH difference is suitable for catheter-based delivery. Garbern et al. developed basic fibroblast growth factor (bFGF) loaded injectable polymer made of poly(N-isopropylacrylamide-copropyl acrylic acid-co-butyl acrylate) gels at pH 6.8 within ischemic myocardium and release bFGF. The in vivo studies in rats have demonstrated increased angiogenesis and improved cardiac function (Garbern et al. 2011). pH-responsive polymers for dental applications have also been reported. Weir et al. had developed dental resin composites comprised of nanoparticles of amorphous calcium phosphate and tetracalcium phosphate for tooth cavity restorations that exhibiting limited ion release at pH 7 and substantial calcium ion release at a pH 4. The in vitro studies using this composite have shown that it achieved about 50% remineralization for the dentin lesions within 8 weeks of time (Weir et al. 2017). Apart from these, another critical use of the bio-responsive materials is to construct smart coatings or surfaces on medical implants. Accumulation of microbes on the biomedical device surface and the subsequent formation of biofilms are a potential health hazard and can lead to implant failure. To this end, pH-responsive polymers have been used in the design of such surfaces to prevent implant-related infections. For instance, Wang et al. coated antibiotics loaded nanotubes on the surfaces of titanium implants. The nanotubes were then capped by coordination linkage using 1,4-bis(imidazol-1-ylmethyl) benzene (BIB) and metal ions. In presence of any bacteria, the acidic metabolism products break the coordination linkage and release the loaded antibiotics (Wang et al. 2017).

16.7

Magneto-Responsive Biomaterials

Magnetic responsive biomaterials are the topic of immense research due to their potential applications in various fields like biomedical, microfluidics and microelectronics (Thévenot et al. 2013). These magneto-responsive materials have evolved and emerged as an active scaffold which can be manipulated spatio-temporally via an external magnetic field for the applications (Fig. 16.7), particularly (a) to offer a controlled mechanical stimulation of tissues thereby enhance the healing response,

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Fig. 16.7 Schematic of magneto-active biomaterials: Magnetic responsive elements such as magnetic nanoparticles are being incorporated into several biomaterial structures for applications in drug delivery and tissue regeneration. Reproduced from (Bruno Lodi and Fanti 2019) # 2019 The Author(s). Licensee IntechOpen

(b) to develop a smart and reliable magnetic drug delivery system and (c) to carry out local hyperthermia against cancer cells by generating therapeutic heat (Adedoyin and Ekenseair 2018; Bruno Lodi and Fanti 2019). Several strategies have been used to develop magneto-responsive biomaterials that include blending, in situ precipitation, covalent bonding and dip coating. Magnetic nanoparticles (MNPs) made of inorganic matter, mostly superparamagnetic iron oxide Fe3O4 or γ-Fe2O3 also known as ‘soft’ metallic iron, and some ‘hard’ magnetic materials, e.g. Co, Ni, FeN, FePt, FeP, doped polymeric materials can create a potential impact on the future of biomedical research (Thévenot et al. 2013). Overall, these magnetoreactive materials can be categorized into (a) those that have the ability to deform, (b) those that offer magnetic guidance remotely and (c) those that induce pseudoshape memory effect via actuation or other phenomenon (Thévenot et al. 2013). In the context of tissue engineering, cells can be either mixed with the biomaterial prior to solidification or be seeded onto the scaffold surface after solidification (Li et al. 2013). Magneto-responsive biomaterials with specific multifunctional properties are subjected to a static magnetic field to induce localized forces on cells there by initiating a controlled therapeutic action with different intensities over a continuous period of time (Bruno Lodi and Fanti 2019). Langer and co-workers proposed the concept of magnetically activated delivery of drugs (Langer and Peppas 1981) where magneto-responsive hydrogels of ethylene vinyl acetate were loaded with insulin. Since then the magneto-responsive biomaterials loaded with drugs or other therapeutic compounds were explored and manipulated spatiotemporally to release drugs at specific concentrations on-demand via internal/

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external magnetic signaling (Qiu and Park 2001). Recently, Adedoyin and Ekenseair described extensive use of MNPs in drug delivery, hyperthermia, magnetic resonance imaging, cellular targeting, tissue engineering and beyond (Adedoyin and Ekenseair 2018). For instance, Huges et al. demonstrated that to activate mechanosensitive ion channels located on the cellular membrane they used MNPs thereby activated the biochemical pathways that directly influenced osmoregulation (Hughes et al. 2005). Lately, magneto-responsive materials are being explored in 3D printing in order to expand the field into 4D bioprinting. For instance, a polymeric ink infused with magnetic iron oxide nanoparticles was formulated by Wei et al., and it is utilized for direct-write printing to fabricate tubular structures which are magnetically guidable and have shape recovery properties for the creation of magneto-responsive constructs using additive manufacturing technology (Wei et al. 2017).

16.8

Electro-Responsive Biomaterials

Flexibility in material properties allowed the development of a wide range of materials with electro-activity that includes conductive conjugated polymers, percolated conductive composites and ionic conductive hydrogels. These versatile electroactive materials are designed to respond under the influence of an electric stimulus, there by triggering outstanding properties suitable for biomedical applications (Sequeira 1983). When compared to other categories of biomaterials, electroactive smart materials have been less studied for biomedical applications (Tandon et al. 2018). Generally, electroactive biomaterials allow the direct delivery of electrical, electrochemical and electromechanical stimulation to cell (GhasemiMobarakeh et al. 2011; Rivers et al. 2002). Major group of electroactive biomaterials include conductive polymers, electrets, piezoelectric and photovoltaic materials (Fig. 16.8). Among these groups, electrets and piezoelectric materials do not require an external power source to allow delivery of an electrical stimulus but control over the stimulus is limited (Rivers et al. 2002). Electroactive polymers can be classified into two types based on the mechanism of conduction, viz. (a) ionic and (b) electric conductive polymers. Typically, the ionic conductive polymers show conductivity due to the ionic groups present in their main chain structure or by the presence of electrolytes in the medium. The electric conductive polymers are classified into intrinsic and extrinsic, based on their mechanism of electron conduction. These polymers appear to be highly conductive perhaps due to relatively high electron movement caused by constitutive bonds between atoms or the conductive elements present (Palza et al. 2019). Electro-stimuli biomaterials are not only being explored to stimulate cells and tissues in the field of tissue engineering and regenerative medicine, but are also being explored to develop smart drug delivery systems and as artificial muscle (Rivers et al. 2002; Balint et al. 2014). Novel electroactive materials are produced totally based on the bactericidal effect of electrical stimulation. It has the capacity to prevent biofilm formation and future bacterial infections in the host. Therefore, we can

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Fig. 16.8 Schematic of electroactive biomaterials: Conductive, piezoelectric, electret and photovoltaic biomaterials are explored in biomedical field, including drug delivery and tissue regeneration. Reproduced with permission from (Tandon et al. 2018). Copyright # 2017 Elsevier Ltd

merge a polymer capable to deliver electric stimulus and the necessities needed for any biomaterial designed for tissue engineering, i.e. a biomaterial that promote cellular adhesion and proliferation while capable of avoiding biofilm formation through a bactericidal effect (Sequeira 1983). In the context of drug delivery, electroactive biomaterial holds great potential for controlled therapeutic delivery when subjected to electrical stimulation. With this approach, one can explore delivery of various biochemical molecules such as genes, proteins and RNA molecules, both in a localized, controlled and efficient manner. Tandon et al. demonstrated that through binding biologically important molecules into the polymer, physical properties of electroactive biomaterial can easily be optimized for a specific application. Sponge-like structures were created by polymerizing polypyrrole a conductive polymer around sacrificial nano or microbeads. Such nanostructures were used for the release of rhodamine B, dexamethasone, fluorescein chlorpromazine and risperidone (Tandon et al. 2018).

16.9

Bio-Responsive Biomaterials

Smart biomaterials are also designed to adapt their physico-chemical properties in response to changes in physiological parameters (Fig. 16.9). These bio-responsive polymers are designed to elicit specific biomedical functions in response to physiological and pathological conditions in the body such as cell/tissue/organ level pH, local redox potential, enzyme activity, body temperature, etc. (Hu et al. 2014; Kowalski et al. 2018). The choice of specific stimulus will always depend on the nature of target tissue. Among these biomaterials that respond to physiological pH

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Fig. 16.9 A representation of various biological stimuli: physiological temperature, pH, reactive oxygen species, local oxygen concentration, glucose levels and several enzymes can be considered as stimuli to design bio-responsive intelligent systems

and temperature are most widely explored and are covered in earlier sections. In this section, we cover enzyme-responsive, stress-responsive and immune-responsive biomaterials.

16.9.1 Enzyme-Responsive Biomaterials In living organisms almost all physiological activities are controlled by various enzymes. Use of enzymes as the trigger to change the properties of a particular polymer has received increasing attention due to the specificity and selectivity of enzymes for the substrates. Enzyme-responsive materials are a new class of smart materials which are triggered by selective catalytic actions of enzymes thereby undergoing macroscopic transitions. The major reason that makes enzymeresponsive systems especially suitable for biomedical application is the high catalytic efficacy and inherent biocompatibility (Zhang et al. 2019). The use of enzymes as stimuli to trigger mechanical responses in materials opens up a number of possible applications (Ulijn 2006; Zelzer et al. 2013). The enzyme-responsive materials are composed of two units: (a) an enzyme sensitive unit, which is a substrate or substrate mimic based on a biomolecule (such as a peptide, a lipid or a polynucleotide) and (b) a unit that directs and control changes in non-covalent interactions that cause macroscopic transitions (Hu et al. 2014). Proteases, endonucleases, kinases, phosphatases and quite a few other enzymes can be explored for developing enzyme-responsive materials (Cinay et al. 2017). The ability of proteases to cleave peptides and endonucleases to cleave oligonucleotides is typically used to degrade or disassemble enzyme-responsive materials. The physiological processes are accurately regulated to control level and activity of enzymes, however, pathological conditions may see altered expression of enzymes. These changes can be explored as the trigger while designing enzyme-responsive systems (Zelzer et al. 2013). Enzyme-responsive hydrogel based chemical hydrogels is the most commonly used one; Sperinde and Griffith demonstrated the formation of a hydrogel by the crosslinking of PEG-Gln-NH2 with poly(Lys:Phe) and developed methods to predict gelation kinetics in these systems (Sperinde and Griffith 1997). Very high sensitivity biosensing can be achieved by using enzyme-responsive biomaterial,

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for example, sensing of the anthrax lethal factor protease relevant to bioterrorism detection (Guarise et al. 2006).

16.9.2 Stress-Responsive Biomaterials Oxidative stress is associated with a number of pathologies, making reactive oxygen species (ROS) targeting an efficient strategy to selectively deliver a therapeutic pay load to diseased cells or tissues especially in inflammatory tissues. Based on how the polymer react to oxidative stimuli, two classes are identified, viz. (a) those with an oxidation-induced polymer solubility switch and (b) those that undergo oxidation-induced polymer degradation (Lee et al. 2013). The polymers which are capable of undergoing a ROS-mediated solubility switch include poly (propylene sulphide) and selenium-linked polymers. They are typically formulated as hydrophobic materials, which upon oxidation by cell-generated ROS become hydrophilic and thereby disintegrate to trigger drug release or other associated function (Napoli et al. 2004; Vasir and Labhasetwar 2007). Many studies and trials explored oxidation-responsive biomaterial drug delivery for ROS-mediated release of therapeutics in both intracellular and extracellular environments (Lee et al. 2013; Martin and Duvall 2016; Xu et al. 2016; Yao et al. 2019). One of the popular ROS-sensitive materials investigated so far are the thioether-containing polymers that undergo phase transition from hydrophobic sulphide to more hydrophilic sulphoxide or sulphone during oxidative condition. To design ROS-responsive polymeric vesicles thioether based block polymer consisting of poly(ethylene glycol) (PEG) and poly(propylene sulphide) (PPS) commonly known as ABA block polymer were used (Napoli et al. 2004). Another class of ROS materials undergo degradation and leave porous structure behind. This property can be very well explored to develop ROS-responsive polymeric scaffolds and hydrogels that could modulate the cell response in the field of tissue engineering. For example, an ABC triblock polymer poly [(propylene sulphide)-b-(N,N-dimethylacrylamide)-b-(Nisopropylacrylamide)] was synthesized as a ROS-degradable thermoresponsive hydrogel which is widely used nowadays (Gupta et al. 2014).

16.9.3 Immuno-Responsive Biomaterials Biomaterials of varying composition, structure and properties are widely explored in several cell- and immune-therapies. But the clinical successes of these approaches are rare since our immune system adversely reacts to the use of biomaterials, recognizing it as a foreign subject and leading to graft rejection. Therefore, it is very critical to design biomaterials or implants thereof that could potentially evade immunological rejection (Andorko and Jewell 2017). Recently smart biomaterials which are capable of modulating the immune system are under intense investigation to improve the overall clinical success in cell- and immune-therapies (Fan and Moon 2015; Kowalski et al. 2018). Typically, immune-responsive biomaterials can be

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designed to evade immunity or to activate immunity (Lee et al. 2013; Martin and Duvall 2016; Xu et al. 2016; Andorko and Jewell 2017). Recent endeavours in cancer research have mainly focused on cancer immunotherapy involving sustained release of therapeutic molecules there by activating the patient’s immune system. Smart biomaterials can potentially promote immunomodulation for personalized cancer treatment. A bio-active polymeric implant scaffold integrated with collagen-mimetic peptide unit has been developed by Stephan et al. which can bind to lymphocyte surface and is able to deliver, expand and disperse T cells reactive to tumours (Stephan et al. 2015). In the field of tissue engineering, exploring the intrinsic immunomodulatory properties of mesenchymal stem cells, several investigators are working on the development of MSC-laden scaffolds to reduce the immunological reaction towards the graft (Le Blanc and Ringdén 2007; Molina et al. 2015). In the field of cell therapy, particularly, pancreatic islet transplantation to treat type I diabetes, several interesting immune-protective biomaterials that can protect the transplanted islets from recurrent auto-immune action are emerging (Anderson et al. 2008; Ward 2008).

16.10 Other Stimuli-Responsive Biomaterials Apart from the above categorized intelligent biomaterial systems, there are quite a few other intelligent biomaterial systems that are emerging in recent years. In the field of oncology, ultrasound therapy is being explored as a means to locally heat up the cancer tissue and thereby eradicate the cancer tissue. Along with this hyperthermia treatment, ultrasound can also be explored as a means to stimuli materials and thereby control drug release and delivery. Typically, application of ultrasound offers more than one stimulus that includes thermal induction, mechanical stimulus or gas vaporization. Besides influencing material properties, ultrasound also influences the tissue by enhancing tissue, cell and specialized barriers permeation that include the blood–brain barrier (Phenix et al. 2014). Besides such unique single-responsive systems, other multi- or combinatorial systems are also emerging to achieve the desired goals overcoming the challenges posed by the complex physiological microenvironment (Fig. 16.10) (Zhang et al. 2019). Typically, three kinds of multi-stimuli-responsive polymer materials are available that includes: (a) multistimuli-responsive particles-micelles, micro/nanogels, vesicles and hybrid particle, (b) multi-stimuli-responsive films polymer-brushes, layer-by-layer polymer films and porous membranes and (c) multi-stimuli-responsive bulk gels—hydrogels, organogels and metallogels (Zhang et al. 2019; Cao and Wang 2016). Current multistimuli-responsive systems are generally designed by incorporating additional responsiveness property to a pH-responsive system, including pH-reduction responsive, pH-light responsive, pH-diol responsive and pH-temperature responsive systems (Fu et al. 2018). A representative example of multi-responsive material includes a temperature and pH-responsive hydrogel which was synthesized by PNIPAM and positively charged poly(diallyldimethyl ammonium chloride)

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Fig. 16.10 An illustration showing dual and multi-responsive intelligent biomaterials: A combination of two or more physical, chemical and biological parameters have been used as trigger in developing intelligent biomaterial systems for an array of biomedical applications including tissue engineering

and which was assembled on a layer of gold-coated poly dimethyl siloxane (Zhang et al. 2019).

16.11 Summary and Future Prospects In recent years, there has been a paradigm shift in the development of biomaterials from inert materials to intelligent biomaterials or structures in order to achieve a desired material function in the context of tissue engineering and beyond. Starting with shape memory alloys in the 1930s, the field of intelligent biomaterials has evolved over the years with introduction of several biomaterials that sense and react to various physical, chemical and biological stimuli. In this chapter, we have attempted to provide a history of intelligent biomaterials, followed by state of the art in shape changing materials, thermoresponsive, light-sensing, pH-reactive, magneto-stimuli and bio-active biomaterials. We have also detailed other distinctive and emerging intelligent biomaterials under development including ultrasoundreactive, dual or multi-reactive biomaterials. We envisage further development in the field of intelligent biomaterials since the research and developmental activities across the globe are intense in recent years. Particularly, there is great potential for the dual or the multi-responsive intelligent biomaterial systems in various contexts of biomedical engineering. Further, artificial intelligence, internet of things and many innovations in other fields may also see interesting relationships with intelligent biomaterials in the near future. However, as one imparts intelligence to the biomaterials and thereby gives biological relevance and anticipate a drug-like or other biomimetic action, there will be significant regulatory implications. Although the field is still evolving, we feel it is important to focus on such regulatory as well as other associated complications in order to streamline the clinical translation of intelligent biomaterial systems.

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Acknowledgements Authors acknowledge Sree Chitra Tirunal Institute for Medical Sciences and Technology (An Institution of National Importance), Thiruvananthapuram and Department of Science and Technology, Government of India for funding (TRC-P8223).

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3D Bioprinting in Tissue Engineering and Regenerative Medicine: Current Landscape and Future Prospects

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J. Anupama Sekar, R. K. Athira, T. S. Lakshmi, Shiny Velayudhan, Anugya Bhatt, P. R. Anil Kumar, and Naresh Kasoju

Abstract

Tissue engineering and regenerative medicine typically involve design and development of a biocompatible and bioresorbable scaffold, followed by seeding and culturing of cells within the scaffold, and subsequent maturation into a tissue of interest. However, inherent diffusion restrictions in such top-down tissue engineering approach make the construction of large-scale tissues far from reality. 3D bioprinting technology has emerged as an innovative bottom-up tissue engineering approach which aims to biofabricate clinical scale tissues/organs with intricate details. This fascinating computer-aided additive manufacturing process typically involves extrusion of cell-laden hydrogels, popularly termed as bioinks, through a nozzle of definite diameter and collection or otherwise printing on a base in the form of a 3D structure of interest. In this chapter, we start with (a) historical background, typical setup, work flow, types and working principles of 3D bioprinting, (b) insights into bioinks including ideal characteristics of a

J. Anupama Sekar, R. K. Athira and T. S. Lakshmi contributed equally. J. Anupama Sekar · R. K. Athira · P. R. Anil Kumar (*) · N. Kasoju (*) Division of Tissue Culture, Department of Applied Biology, Biomedical Technology Wing, Sree Chitra Tirunal Institute for Medical Sciences and Technology, Thiruvananthapuram, Kerala, India e-mail: [email protected]; [email protected] T. S. Lakshmi · A. Bhatt Division of Thrombosis Research, Department of Applied Biology, Biomedical Technology Wing, Sree Chitra Tirunal Institute for Medical Sciences and Technology, Thiruvananthapuram, Kerala, India S. Velayudhan Division of Dental Products, Department of Biomaterials Science and Technology, Biomedical Technology Wing, Sree Chitra Tirunal Institute for Medical Sciences and Technology, Thiruvananthapuram, Kerala, India # The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 B. Bhaskar et al. (eds.), Biomaterials in Tissue Engineering and Regenerative Medicine, https://doi.org/10.1007/978-981-16-0002-9_17

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bioink, various types of polymeric tissue specific and combinatorial formulations, and (c) various approaches in 3D bioprinting such as those based on one component, multi-component, sacrificial component based formulations and other combinatorial modes including electrospinning. The chapter ends with a summary of the key advances up to date and a note on future prospects. Keywords

Biofabrication · Bioinks · Bottom-up tissue engineering · Additive manufacturing · Organ printing

Abbreviations 3D CAD ECM FDM HUVEC RGD SLA UV VEGF

17.1

Three dimensional Computer aided design Extracellular matrix Fused deposition modeling Human umbilical endothelial cells Arginine-glycine-aspartate Stereolithography Ultraviolet Vascular endothelial growth factor

Introduction

The increasing incidences of vital organ failure have led to the concept of organ transplantation which was indeed a captivating revolutionary change. This was accomplished successfully for the first time by Drs. Joseph Murray and John Merrill of Peter Bent Brigham Hospital, wherein they undertook the challenge of transplanting a kidney from a healthy donor to its identical twin (Tan and Merchant 2019). However, as contended by Dr. Murray himself that every organ transplant comes with its own challenge and the list goes on from organ rejection to finding a suitable donor. Majority of these issues can be overcome by building tissues from scratch using patient specific cells that not only prevent organ rejection but can also meet the rising organ transplant demand (Dzobo et al. 2018). This new generation medicine termed as tissue engineering typically involves development of a biocompatible scaffold followed by seeding of cells on top of scaffold, which then gradually proliferate and mature into a tissue of interest (Fig. 17.1) (Eltom et al. 2019). This approach has been clinically successful in the development of tissues such as skin. Unfortunately, construction of large organs such as liver and kidney is far from

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Fig. 17.1 Overview of tissue engineering: cells, scaffolding materials, and bioactive cues are combined together to engineer artificial tissues that can be transplanted into patients ((a) overall scheme, (b) top-down approach, and (c) bottom-up approach)

reality. The major bottleneck with conventional top-down tissue engineering is the restricted diffusion of nutrients and gases beyond 200 μm due to lack of proper vasculature-like channels (Lovett et al. 2009; Rademakers et al. 2019). Therefore, a bottom-up tissue engineering approach involving assembling of micro-scale building blocks, including those that create vasculature-like channel network, in an orderly fashion has been attempted. One such highly attempted yet still emerging bottom-up tissue engineering approach is 3D bioprinting (Elbert 2011). This is a fascinating computer-aided additive manufacturing process where cell-laden hydrogel formulations called bioinks are loaded into a cartridge and are printed on a substrate to yield 3D objects as per a computer-aided design (Li et al. 2020). In this chapter, we begin with a brief background on 3D bioprinting (Sect. 17.2), followed by insights into bioinks for 3D bioprinting (Sect. 17.3), details of various approaches in 3D bioprinting (Sect. 17.4). In particular, Sect. 17.2 of the chapter gives a historical background of 3D printing with details of earlier conceptualization and its eventual evolution into 3D “bio” printing, followed by details of typical set up as well as workflow of 3D printing and modifications thereof towards 3D bioprinting, and details of three popular modes of 3D bioprinting approaches

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including their working principles. Section 17.3 describes the details of the most critical element of 3D bioprinting, i.e., the bioink. The sections highlight the features of an ideal bioink, then gives an outline of various formulations of polymeric bioinks, followed by the developments in the field of tissue specific bioinks using compounds derived from decellularized extracellular matrix, and various other formulations of combinational character. Section 17.4 gives insights into the various bioprinting approaches in order to develop tissue engineering constructs. It covers various reports based on one component bioink formulations, followed by two or multi-component bioink formulations including sacrificial elements and other emerging approaches. We conclude the chapter with Sect. 17.5 describing a summary of the developments so far and a note on emerging as well as future prospects.

17.2

Background to 3D Bioprinting

17.2.1 Historical Account of 3D Printing/Bioprinting Started as a naïve concept way back in 1970s, the 3D printing has evolved rapidly over the last few years and became a state-of-the-art manufacturing technology (Fig. 17.2). The concept of 3D printing envisaged back in 1970s, but the experiments are dated from 1981. Japanese inventor Hideo Kodama, in 1981, through an additive manufacturing process used UV to harden polymers and created solid objects (Savini and Savini 2015). This was a major step forward in the field of stereolithography (SLA). Subsequently, in 1983, Charles Hull created smaller version 3D objects by layer-by-layer and UV light exposure using 3D printing technology. Other developments in the field of additive manufacturing in the same timeline, such as laminated object manufacturing technology which involves fusion of multiple cross sectional laminates of a 3D object, and selective laser sintering which involves fusion deposition of powdered polymer to create 3D object were not as successful

Fig. 17.2 Timeline of events in 3D printing technology: the additive manufacturing technology has evolved from 1981 with a report 3D additive fabrication to date with various reports of innovative high-tech printing machines

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as SLA or 3D printing (Shakor et al. 2019; Jiménez et al. 2019). Inspired by inkjet technology, first developed by Canon Co. in 1979, Massachusetts Institute of Technology developed and popularized 3D printing technology in the early 1990s. Subsequently, the 3D printing technology has evolved and expanded its footprint from conventional manufacturing field to bio-manufacturing with a great promise of printing human organs (Starly and Shirwaiker 2015). Here, the conventional 3D printing machine was adopted and modified to use human cells—containing hydrogel as bioink to fabricate live 3D objects. To this end, pioneering efforts of Wake Forest Institute for Regenerative Medicine on bioprinting involving 3D printing of a synthetic scaffold of human bladder using a 3D printer and culturing of patient specific cells led to the successful biofabrication of tissues and thereby opened up new avenues in regenerative medicine (Munoz-Abraham et al. 2016).

17.2.2 Set Up and Work Flow of 3D Printing/Bioprinting 3D printers are available in a broad range but in a quintessence, they are all computer controlled additive manufacturing machines (Starly and Shirwaiker 2015). There are mainly four processes of 3D printing, viz. digital light processing using planar projection, fused deposition modeling using filaments (FDM), inkjet printing using microspheres and selective laser sintering using powders (Zeming et al. 2019). However, FDM approach has gained tremendous popularity among all, owing to its simplicity and versatility. In a typical FDM setup (Fig. 17.3), a 3D printing unit consists of (a) a print bed with/without cooling pad and (b) an extrusion type print head having an inlet for a polymeric filament, a heating unit to melt the filament and a nozzle to extrude the molten polymer. Typically, print head and print bed have motion controls in X, Y, and Z axis and are controlled by software (Munoz-Abraham

Fig. 17.3 Schematic of 3D printer vs 3D bioprinter: components and setup for both 3D printer and 3D bioprinter are nearly same, except that the 3D bioprinter involves cell-laden bioink and the entire process is ready in aseptic conditions

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et al. 2016; Carneiro et al. 2015). As for the workflow, a typical 3D printing process involves generation of a digital model using computer-aided design (CAD); uploading and digital sectioning of the model into thin layers in 3D printer software, followed by computer-aided printing and post-fabrication processing if necessary. This typical 3D printing setup and workflow have been conveniently evolved into 3D “bio” printing wherein cell-laden hydrogel formulations were used as “bioinks” to fabricate living structures for potential tissue engineering and regenerative medicine applications (Walker and Humphries 2019). Major alterations in the setup include introduction of ascetic chamber and modifications to print head to handle bioink loaded cartridges, however, the overall workflow of 3D bioprinting is very much similar to conventional 3D printing (Hospodiuk et al. 2017).

17.2.3 Types and Principles of 3D Bioprinting Based on this typical setup and workflow, a wide variety of 3D bioprinting approaches are emerging (Munoz-Abraham et al. 2016), however, the most popular of all are extrusion printing, inkjet printing, and laser assisted printing (Fig. 17.4).

Fig. 17.4 Typical modes of 3D bioprinting: the printing of bioinks to design of user choice is done by either extrusion mode (a), ink-jet mode (b), or laser assisted mode (c)

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17.2.3.1 Extrusion Bioprinting It is widely used because of its simplicity and affordability. Here, the cell-laden bioink is extruded as a continuous filament through a nozzle and deposited on a substrate to form desired structures. Depending on the actuating modes of liquid dispensing system, extrusion printing can be categorized into pneumatic-, piston-, and screw-driven. In pneumatic driven system, compressed air drives the bioink through the nozzle, in piston-driven system, a mechanically driven syringe pump pushes the bioink, and in screw-driven system the bioink slides through a screw that is connected to the motor (Pati et al. 2014). This approach has been widely explored in bioprinting field including creation of constructs with channeled network and even microfluidic chips. This approach is also popular because of its versatility to create structures using multiple materials at once. In spite of such impressive versatility, the extrusion-based bioprinting suffers from relatively poor printing resolution. Besides, it is challenging to formulate a bioink that is able to maintain mechanical integrity while not causing shear stresses on cells while printing. 17.2.3.2 Inkjet Bioprinting It is yet another most commonly used type of modality in 3D bioprinting field. Here, the cell-laden bioink is passed through a specialized nozzle wherein thermal or electric forces release the bioink in the form of droplets (Shintaroh et al. 2015). Thermal inkjet printers use electricity to heat the print head and produce pulses of vapor pressure to break the bioink flow into droplets, whereas piezoelectric valve based inkjet printers use a valve that creates acoustic wave inside the print head to form the bioink droplets and the disposition of the cells on to the scaffold (MunozAbraham et al. 2016; Zeming et al. 2019). Inkjet bioprinting offers excellent printing resolution of up to 50 μm and it was also reported to give constructs with a very high cell viability. Besides, this also enables printing multiple cell types at once. However, not all materials can be printed with this approach because of limitations on the viscosities of the bioinks, therefore, this approach may have to be combined with other approaches to make complex multi-cellular tissue structures. 17.2.3.3 Laser Assisted Bioprinting It is based on the principle of laser-induced forward transfer. Briefly, it functions using focused laser pulses to generate a high-pressure bubble that propels cell-laden bioink toward the collector. Although relatively less common, this technique has advantage of working in non-contacting and nozzle-free printing process; further, it is compatible with a range of viscosities (1–300 mPa/s) and can print mammalian cells with negligible effect on cell viability and physiological functionality (Guillemot et al. 2010). This particular approach has gained much attention from scientific community because of lack of shear stress on cells which means the cell viability would be unaffected, also, there is no restrictions on viscosities of bioink. However, the use of lasers to print and the use of UV and near UV light to crosslink bioink may cause photo-toxicity to cells including potential DNA damage. Nevertheless, several researchers are working on developing photoinitiators that work in visible range.

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Bioinks in 3D Bioprinting

Bioink is the key component in the process of 3D bioprinting. It represents a cocktail of cells or cell aggregates in a suitable medium with or without an externally provided extra cellular matrix (ECM) that are used for biofabrication of 3D constructs using a 3D bioprinter (Whitford and Hoying 2016). The ECM component in the bioink would be generally water-based hydrogel systems that functions like a biocompatible cell delivery agent and provide the necessary mass transfer (Tibbitt and Anseth 2009). Bioinks can be defined as “a formulation of cells that is suitable to be processed by an automated biofabrication technology” (Groll et al. 2018). The “bio” in the term bioink indicates living cells and hence the formulation will qualify to be called as a bioink only if cells are included. Thus, high concentration of cells or aggregates of cells in an appropriate culture medium can also been suggested as bioink. Such bioinks are used in dispensing bioprinters that works on extrusion principle. This approach takes the advantage of spontaneous cellular assembly and self-organizing capability of cells rather than providing an additional biomaterial for biofabrication of tissues (Guillotin et al. 2010). The biomaterial used in bioinks should possess properties such as biocompatibility, printability, and mechanical property before and after the process of bioprinting (Fig. 17.5). Since the bioink performs the function of ECM in the bioprinted construct, the standard evaluation processes that apply for a tissue engineering scaffold will also be required here. The physico-chemical characteristics that would be general for a wide range of polymeric hydrogel based bioink are viscosity, viscoelasticity, shear thinning, gelation kinetics, and degree of hydration (Gao et al. 2019). Certain important biological evaluations that can be adopted for biological evaluation of bioinks can be listed as cytotoxicity,

Fig. 17.5 Features of a bioink: a bioink should typically meet several physico-chemical and biological requirements in order to be used in bioprinting process

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cell adhesion, cell proliferation, cell encapsulation, in vivo biocompatibility, and tissue specific functionality of bioprinted construct (Ashammakhi et al. 2019). Different kinds of biomaterials have been used for bioink designed to be used with one cell type (homogeneous bioink) or multiple cell types (heterogeneous bioink) (Colosi et al. 2017). Various polymers have been extensively studied for application as bioink. It includes natural and synthetic polymers which basically form a hydrogel (Turunen et al. 2018). The bioinks are generally classified based on its rheological properties as non-shear thinning and shear thinning bioinks. Shear thinning bioinks set back and attain the physical gelation quickly after being extruded by the bioprinter (Highley et al. 2015). The non-shear thinning bioinks need to be stabilized by crosslinking externally. The crosslinking of bioinks can be of physical crosslinking or chemical crosslinking depending upon the bulk polymeric biomaterial used in it. The bioinks are physically crosslinked by ionic gelation, thermocondensation, and self-assembly (Ashammakhi et al. 2019). Crosslinking alginate with calcium ions is an example for physical crosslinking (Lee and Mooney 2012). The formation of covalent bond represents chemical crosslinking as in the case of modified gelatin known as gelatin methacryloyl (Perera and Ayres 2017). Examples for synthetic polymers used as bioink are poly (ethylene glycol) and poly (ethylene glycol) diacrylate (Zhu 2010). The natural biomaterials that are used in bioinks are collagen, gelatin, fibrinogen, alginate, chitosan, hyaluronic acid, and silk. Gungor-Ozkerim et al., have reported an extensive and elaborative review on various synthetic and natural biomaterials classifying with respect to bioprinting method, cell type, and target tissue (Gungor-Ozkerim et al. 2018). Precise positioning of cells is critical in preparing organized tissue. Similarly, organization of ECM in its sophisticated manner is essential for the function of the tissue and organs (Murphy and Atala 2014). Extracellular matrix provides structural integrity to tissues, controlled proteolytic turnover, appropriate Cell–ECM interactions, and cellular responses (Choudhury et al. 2018). This complex interaction has been illustrated by Chen et al., showing the Cell–ECM relationship in the context of cell proliferation, development, and differentiation (Chen and Liu 2016). This complex property is achieved due to the presence of attaining a threedimensional framework supplemented with collagen, glycosaminoglycan, growth factors, elastin, and other signaling molecules (Gopinathan and Noh 2018). Lack of understanding on the accurate composition of the components and the inability to control the precise deposition of ECM and cells remain as the biggest challenge in the biofabrication of the tissues. Depending upon the printing technology the accurate deposition of the cells can be achieved to an extent. However, mimicking the components of ECM and the soluble and insoluble stimulatory molecules within the bioink is not easy. ECM isolated from specific tissue of interest can be used as the bulk component of bioink so as to provide cues of tissue microenvironment. The ECM molecules can be used in bioink with or without modification for crosslinking. Ali et al., reported modification of decellularized ECM into a photocrosslinkable form to bioprint renal tissues (Ali et al. 2019). The decellularized heart tissue ECM was used as bioink, along with vascular endothelial cell growth factor and Vitamin B2, to 3D bioprint vascularized heart construct (Jang et al. 2017). Pati et al., have

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shown that decellularized adipose, cartilage and heart tissues provide crucial cues for cells encapsulation, survival, and functions (Pati et al. 2014). Liver is a complex organ with multitude of physiological functions. Lee et al., successfully bioprinted liver construct using decellularized porcine liver and have shown in vitro maturation and differentiated function of liver (Lee et al. 2017). Kim et al. proposed a novel formulation of microparticles from liver ECM incorporating microparticles within it for enhancing the 3D printability and mechanical properties (Kim et al. 2020). Availability, extraction, and storage of safe tissues for decellularization are also proposed to be a necessary aspect in this area (Kim et al. 2017a). Storage of available tissues, qualified for clinical application with the required regulatory approvals needs new policies and guidelines in future for the extensive use of ECM as bioink.

17.4

Approaches in Bioprinting

17.4.1 Single Component Bioink Based Approaches There are many aspects for the successful bioprinting of a tissue, the most important being the choice of bioink used. There are numerous bioink formulations that include different types of biocompatible materials for bioprinting applications. A researcher decides to choose the formulation for bioink based on the intended structure and function of the organ or tissue to be printed. Bioinks are synthesized from one or more base polymers or other biomaterials based on the intended application. Single component collagen based bioink has been used in many 3D bioprinting applications due to its uniqueness in mimicking as the extracellular matrix that holds the cells together and thus provides a native tissue environment. In a study done by Koch et al., fibroblast and keratinocytes are embedded in collagen using laser assisted 3D bioprinting to develop construct that resemble skin tissue. The study showed that printing cells on collagen matrix have reproduced a skin construct that shows intracellular adhesion and formation of gap junction which are important for cell proliferation and differentiation (Koch et al. 2012). Single component nanocellulose bioink was synthesized by another study group whereby the bioink was 3D printed with customizable infill and seeded with cardiomyoblast (Ajdary et al. 2019). The single component systems thus were able to generate basic models of tissue, however, the choice of biomaterials that can be used for bioink generation using single component is limited and is often sub-optimal. For example, solubility at physiological temperature and pH and presence of RGD moieties make gelatin an excellent candidate, but it is prone to rapid biodegradation in vivo. Similarly, alginate possesses excellent gel properties to allow diffusion of media inside out, but it has weak mechanical stability and lacks RGD moieties.

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17.4.2 Multi-component Bioink Based Approaches In contrast, multi-component system involves two or more polymers and negates the limitations of either biopolymer. For instance, Jia et al. demonstrated that RGD-peptides conjugated alginate dialdehyde mixed with human adipose derived stem cells (hADSCs) printed on gelatin substrate led to precise formation of lattice structure with excellent hADSCs proliferation and spreading (Jia et al. 2014). In another study, methacrylated pectin incorporated with bioactive peptide ligands was subjected to UV photopolymerization and forms dermal layer of skin. It was reported that the bioink system provided necessary cues to the fibroblast cells to produce their own ECMs thereby exhibiting its potential use as a skin substitute (Pereira et al. 2018). In a study done by Ng et al., human dermal fibroblast, keratinocytes, and melanocytes were 3D printed using collagen based bioink and compared it with the manual printing process. The study showed a significant difference in melanin deposition and the uniformity of the pigmentation in 3D printed construct compared to manual printing (Hu et al. 2018). In another study by Giglio et al., a multicompartment 3D printed design was constructed to incorporate human-TERT mesenchymal stem cell and human umbilical vein endothelial cell co-cultured in gellan gum-based hydrogel system that enhanced cell viability and support osteogenesis (De Giglio et al. 2018). As shown in Table 17.1, multi-component systems enhance properties of bioink and make it suitable for 3D bioprinting application. The involvement of different components gives additional property to the bioink in view of improving shear thinning, viscosity, printability, and adaptability to the cells being used. Further, these multi-component systems with the inclusion of different biomaterial components have also enabled inclusion of multiple cell types towards biofabrication of multi-cellular tissues/organs.

17.4.3 Approaches Involving Bioinks with Sacrificial Elements The multi-component systems have further evolved to generate 3D living objects with intricate structural details that offer a solution to one of the biggest problems of restricted nutrient flow in tissue engineering constructs. In native state, the mass transfer of nutrients, metabolic waste and gases, distribution of biochemical signals including growth factors from one cell/tissue to another are achieved by a complex network of vasculature. Unfortunately, in conventional top-down tissue engineering approach, the normal diffusion limit of nutrients and gases required for the metabolic activity is limited to 200 μm in thick tissues due to lack of proper vessel-like network (Carmeliet and Jain 2000). This would further prevent integration of implanted tissue construct with the host tissue environment. To this end, the bottom-up 3D bioprinting offers many potential ways to generate systemic lumen network and maintain the nutrient and gas flow for cell viability throughout the construct. One of the popular approaches to incorporate channel like structure within the construct is the use of sacrificial component which leaves behind a lumen when dissolved (Ji et al. 2019). For instance, sacrificial ink was printed as a sandwich between

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Table 17.1 List of multi-component based bioink system and advantages thereof in comparison to single component system. Bioink formulation Alginate/ Nanosilicate clay, alginate methyl cellulose dialdehyde Nanofibrillated cellulose Polylactide fibers Hydroxyapatite

Gelatin Collagen and fibrinogen Collagen/ gelatin

Methacrylated gelatin and collagen microfibers Alpha tricalcium phosphate Fibrin and vascular endothelial growth factor Atelocollagen and hyaluronic acid

Agarose

Decellularized ECM and silk fibroin Collagen and fibrinogen

Sodium alginate

Carboxymethyl chitosan and alginate

Advantage of multi-component bioink system Improved printability, shape fidelity with controlled release of bioactive agents Provide structural and mechanical support for support cartilage formation Increased hydrogel strength for cartilage tissue engineering Enhanced mechanical property and good cell viability for more than 2 weeks Fabrication of capillary like structure with different diameter Allowed for fine tuning of gelation time and stiffness of gel for better cell proliferation Allowed modification of hydrogel for controlled release of bone morphogenic protein 2 3D bioceramic that showed significantly higher cellular activity and mineralization Better migration of neural cell due to VEGF and fibrin network

Allows heterogeneous tissue regeneration by two different ECM materials Better cell proliferation and mechanical strength compared to collagen alone Stiffer gel was formed. Capillary like network formation was achieved (HUVEC and human dermal fibroblast) Improved mechanical strength but lower cell proliferation of chondrocytes Stable gel with rapid crosslinking for expansion and differentiation of stem cell

References Ahlfeld et al. (2017)

Nguyen et al. (2017)

Kosik-Kozioł et al. (2017) Raja and Yun (2016), Bendtsen et al. (2017) Ruther et al. (2018) Montalbano et al. (2018) Du et al. (2015)

Kim et al. (2017b)

Lee et al. (2010)

Shim et al. (2016)

Lee et al. (2018)

Kreimendahl et al. (2017)

Yang et al. (2018)

Gu et al. (2016, 2017)

(continued)

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Table 17.1 (continued) Bioink formulation Cellulose Nanofibrillated cellulose and alginate Cellulose nanofibrils and polyurethane Cellulose nanofibrils and GelMA

Chitosan

Cellulose nanofibrils, bioactive glass, alginate and gelatin Alginate Gelatin

Hydroxyapatite

Advantage of multi-component bioink system Exhibited better printability at room temperature and low pressure and enhanced crosslinking after printing Generated high viscosity that can be tuned and directly printable hydrogel Enhanced mechanical and structural stability due to efficient crosslinking by Ca2+ and UV Improved printability and better mechanical stability

Improved viscosity and printability with fast gelation Allows fine tuning of gelation time, printability and shear thinning property of gel Formed stable hydrogel with good printability and enhanced cell proliferation and differentiation

References Möller et al. (2017), Markstedt et al. (2015), Martínez Ávila et al. (2016) Chen et al. (2019a)

Xu et al. (2019)

Ojansivu et al. (2019)

Liu et al. (2018) Roehm and Madihally (2017) Demirtaş et al. (2017)

layers of intended cell-laden bioink and subsequently the sacrificial ink was washed away to form the channel. In this study Pluronic F-127 was used as sacrificial layer which was dissolved by placing the construct in phosphate buffer saline solution (Fig. 17.6) (Ji et al. 2019). In a similar approach, Lee et al. used human umbilical vein endothelial cell (HUVECs) loaded gelatin as sacrificial ink and printed within a collagen matrix. Once the gelatin was melted and washed away, the HUVEC cells were deposited on the surrounding collagen matrix (Lee et al. 2014). It is therefore evident that the bottom-up 3D bioprinting approach using sacrificial component has the scope of developing clinical size tissue or organ constructs by overcoming diffusion barrier.

17.4.4 Combinatorial Approaches in 3D Printing/Bioprinting Besides the above-mentioned typical approaches, lately, the capacity of 3D bioprinting has been expanded by following combinational approaches, particularly from electrospinning technique. In broad sense, electrospinning is also a polymer filament producing technique similar to 3D printer; however, electrospinning uses electric charge to generate polymer jet through a nozzle and often results in micro to

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Fig. 17.6 A schematic of 3D bioprinting with sacrificial bioinks: In order to create channels within the bioprinted construct, one can use sacrificial inks which can be dissolved by immersing in appropriate solution. Reproduced with permission from Ji et al. (2019). # 2019 Elsevier Ltd.

nano scale fibers. This technique has been modified into a coaxial setup wherein two immiscible polymer solutions were passed through a two-capillary spinneret to fabricate hollow or core-shell filaments. Inspired by this approach, 3D bioprinting setup was conveniently modified to hold two bioinks to develop 3D objects with hollow vasculature-like lumen network. For instance, a coaxial extrusion-based bioprinting was developed by Gao et al., where alginate was used as bioink that flows externally while calcium chloride solution was allowed to flow internally and thereby creating an endogenous channel that supported cell viability much better than constructs without such channels (Gao et al. 2015). In a similar study, He et al. fabricated coaxial multi-cellular structure using HUVECs and HSMCs to construct a vascular co-culture model (He et al. 2018). Further, combinational approaches of 3D bioprinting with electrospinning and electrospraying are also emerging. For instance, Chen et al., for the first time processed electrospun fibers into an ink, and demonstrated successful fabrication of scaffolds with controlled shapes and large

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Fig. 17.7 A schematic of combinatorial approaches in 3D bioprinting: electrospun fibers were processed into a bioink and successfully demonstrated for its use in cartilage tissue engineering. Reproduced with permission from Chen et al. (2019b). # 2019 Elsevier Ltd.

pores which when combined with chondrocytes resulted in appreciable cartilage regeneration in vivo (Fig. 17.7) (Chen et al. 2019b). In another interesting report, to confer structural precision and mechanical strength to the construct, Yoon et al. used electrospun nanofibers sheets during 3D bioprinting process in a layer by layer fashion (Yoon and Kim 2011). Likewise, the field of 3D bioprinting is evolving by convoluting with several other interesting techniques.

17.5

Summary and Future Prospects

In recent times, the bottom-up tissue engineering approach has gained more attention as it could overcome the inherent issues in the conventional top-down tissue engineering such as low vascularization, low cell density, non-uniform cell distribution, and diffusion limitation. The bottom-up approach mainly involves the assembly of micro-scale cell-laden modules to form larger structures, and in this context, different modes of tissue fabrication have been explored. One such highly attempted yet still emerging approach is 3D bioprinting, which has originally evolved from a non-biological 3D printing technique. 3D bioprinting is a computer-aided additive manufacturing process where cell-laden hydrogel formulations called bioinks are

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loaded into a cartridge and are printed on a substrate to yield 3D objects as per a computer-aided design. From a naïve idea in 1970s, the technology has grown rapidly till date and it is now standing as one of the state-of-the-art manufacturing technologies. Several innovations are reported in the way the technique is adopted to print 3D tissues, viz. extrusion, inkjet, and laser assisted bioprinting. One of the most intensely attempted fields of research in 3D bioprinting is the formulation of bioinks since the characteristics of bioinks decide the fate of the success of 3D bioprinting. A range of polymers such as collagen, gelatin, silk fibroin, etc. have been modified and explored in various combinations and permutations to make a printable formulation. So far, approaches involving single component of cells or bioink, multiple components, and sacrificial components have been reported. Lately, 3D printing/ bioprinting is finding interesting relationships with other equally fascinating techniques such as electrospinning in order to develop novel biomaterial and tissue structures. Based on the literature survey and the developments taking place in the field, we envisage scope for further research and development in (a) bioink formulations, (b) innovations in 3D printer units, (c) innovative approaches in 3D printing/bioprinting and thereby expand its applications further in tissue engineering, drug delivery, and beyond. Acknowledgements Authors acknowledge Department of Science and Technology, Government of India and Sree Chitra Tirunal Institute for Medical Sciences and Technology (An Institution of National Importance), Thiruvananthapuram for funding liver and skin 3D bioprinting programs (TRC-P8141 and TRC-P8137).

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Index

A Acrylatedpoly-L-lactide-co-trimethylene carbonate (aPLA-co-TMC), 545 Activated protein C (APC), 456 Adeno-associated virus, 492, 495, 497 Adenovirus, 492, 498–499 Adipose derived stem cell (ADSCs), 236, 238, 315, 392, 407 Adult stem cells, 9, 143, 364, 471, 472, 491 Alginate (Alg), 97, 98, 111, 122, 138, 141, 144, 150, 167, 185–187, 191–193, 197, 202, 223, 225, 229, 230, 253, 281, 282, 294, 316, 353, 354, 360, 363, 367, 368, 383, 391, 392, 399, 402, 403, 405, 409–411, 426, 434, 441, 447, 450, 452, 453, 475–477, 488, 499, 510, 542, 546, 569–574 Alkaline phosphatase (ALP), 27, 67, 74, 200, 319, 329, 517, 518 Allogenic, 169, 203, 255, 397, 507 Allografts, 6, 8, 14, 75, 140, 349, 360, 370, 371, 387, 390, 391, 396, 428, 475 Aluminium (Al), 24 Angiogenic peptide (AP), 366 Angiopoietin-1 (ANG-1), 225 Anterior cruciate ligament (ACL), 46, 48, 510 Arginine-Glycine-Aspartate/Arg-Gly-Asp (RGD), 31, 180–182, 184, 248, 251, 275–277, 282, 283, 294, 412, 570 Artificial nerve grafts (ANG), 145, 428 Atomic layer deposition (ALD), 246 Autologous, 8, 14, 15, 75, 98, 99, 101, 169, 181, 218, 253, 349, 364, 368, 387, 390, 392, 396, 397, 399, 402, 427, 429, 435, 445, 516

B Bacterial cellulose (BC), 121, 132, 191, 193, 200–202, 330, 332 Barium titanate (BaTiO3), 319, 320 Basic fibroblast growth factor (bFGF), 74, 102, 223, 332, 392, 409, 410, 547 Bioactive, 9–12, 31, 32, 42, 46, 47, 73, 76, 93, 102, 105, 106, 108, 122, 124, 125, 146, 150, 170, 174, 180, 188, 194, 221–223, 231, 239, 247, 251, 256, 273–284, 294, 296, 298, 299, 308, 348, 350–353, 356, 359, 360, 362, 363, 367, 390, 393, 399, 402, 412, 427, 440, 441, 443, 447, 457, 490, 511, 519–521, 526, 563, 571–573 Bioadhesives, 181, 184, 189, 194, 276, 277 Bioceramics, 62–76, 178, 188, 285, 296, 351, 356, 363, 371, 572 Biocompatibility, 5–7, 21, 22, 24–27, 31, 34– 36, 39, 41–48, 63, 94, 99, 106, 166, 167, 180, 181, 187, 188, 194, 195, 200, 202, 217, 218, 221, 222, 228, 230, 238, 244, 245, 247, 248, 252, 254–257, 275, 281, 282, 284, 285, 287, 294, 298, 308, 314, 317, 320, 327, 330, 335, 337, 350, 356, 363, 370, 383, 389, 399, 427, 433, 434, 437, 441, 442, 444, 450, 452, 456, 470, 473–475, 487, 488, 508, 509, 513, 514, 516, 519–521, 523, 525, 543, 551, 568, 569 Biodegradability, 5, 22, 24, 32, 38, 48, 92, 93, 95, 121, 122, 126, 129, 133, 144, 166, 167, 180, 187, 188, 195, 202, 256, 273, 274, 285, 308, 314, 327, 336, 351, 353– 355, 369, 434, 452, 470, 473, 508, 519, 523 Biofilms, 191, 221, 248, 547, 549, 550

# The Author(s), under exclusive license to Springer Nature Singapore Pte Ltd. 2021 B. Bhaskar et al. (eds.), Biomaterials in Tissue Engineering and Regenerative Medicine, https://doi.org/10.1007/978-981-16-0002-9

581

582 Bioinert, 105 Biomaterials, 3–16, 21–49, 63, 70, 90–112, 120–151, 165–203, 217–257, 272–299, 308–337, 348–371, 382–413, 426–458, 470–478, 485–500, 506–510, 515–521, 523–525, 536–554, 568–570, 576 Biomimetic gelatinmethacrylamide (Bio-GelMA), 238 Biomimic, 515, 518, 525 Biopolymers, 92–98, 143, 166, 168, 169, 176, 178, 179, 181, 182, 185, 186, 189–191, 194, 195, 202, 203, 222, 253, 254, 363, 435, 452, 493, 507, 508, 571 Bioresorption, 167 Biosensors, 24, 49, 144, 151, 179, 474 Biosorbable, 122 1,4-Bis(imidazol-1-ylmethyl) Benzene (BIB), 547 Bisvinyl sulfonemethyl (BVSM), 433 Bonemarrow cells (BMCs), 388 Bone marrow mesenchymal stem cells\Bone marrow-derived mesenchymal stem cells (BMSC), 251, 291, 332, 367, 396, 434, 439, 440, 447, 450, 516, 518 Bone morphogenetic protein (BMPs), 350, 361, 362 Bone morphogenetic protein-2 (BMP2), 31, 294, 365, 368, 397, 519 Brain-derived neurotrophic factor (BDNF), 433, 439 Bromodeoxyuridine (BrdU), 27 B-tricalcium phosphate (b-TCP), 367 Bulk metallic glass (BMG), 41, 44, 149 C Calcitonin gene-related peptide (CGRP), 433 Calcium (Ca), 38, 62–64, 66–76, 123, 124, 137, 140, 142, 171, 186, 188, 189, 200, 222, 229, 238, 240, 244, 255, 296, 324, 325, 327, 348, 354, 356, 363, 449, 547, 569 Calcium chloride (CaCl2), 181, 368, 392, 574 Calcium phosphate (Ca(PO4)2/CaP), 62–64, 66, 68, 69, 75, 188, 246, 356, 360–363, 366, 367, 395, 474, 477, 515–517, 547 Calcium phosphate cements (CPCs), 72, 73, 236 Carbon nanofibers (CNFs), 326, 328, 514 Carbon nano tubes (CNTs), 326–329, 444, 445, 514, 515 Carboxymethylated chitosan (CCS), 239

Index Carboxymethyl chitosan (CMCS), 391, 450, 572 Carboxymethyl-Hexanoyl Chitosan (CHC), 542 Cardiovascular Diseases, 387, 475 Carrageenan (CRG), 187–188, 197 Cartilage oligomeric protein (CGN), 285 Cellulose (C), 92, 99, 121, 128, 132, 134, 135, 149, 150, 168, 177, 178, 191–193, 196, 197, 200–202, 230, 282, 329, 330, 332, 353, 397, 450, 475, 477, 488, 507, 508, 545, 572, 573 Central nervous system (CNS), 145, 281, 427, 440, 476, 497, 511, 513 Chemical vapour deposition (CVD), 37, 244, 245 Chemical vapour infiltration (CVI), 37 Chimeric antigen recipient T (CAR-T), 499 Chitosan (CS), 39, 42, 92, 96–97, 103, 105, 111, 131, 132, 138, 140, 141, 144, 145, 147, 150, 166, 168, 177, 184, 185, 188, 194, 197, 198, 202, 225, 227–230, 233, 238–241, 251, 253, 254, 278, 281, 282, 285, 288, 292, 295, 317, 322, 329, 330, 332, 354, 360, 363, 365, 366, 391, 392, 399, 405, 409, 426, 451, 475, 488, 493, 507–509, 515, 517, 523, 524, 540, 541, 546, 569, 573 Chitosan oligmer methacrylate (ChitoMA), 188 Chondroitin sulfate (CS), 185, 186, 316, 393, 394, 396, 397, 399, 401, 403, 405, 410, 431, 433, 435, 440–442, 447, 450, 452 Chondroitin sulfate-incorporated starPEG nanocoating (CS-PEG), 456 Chondroitin sulfate proteoglycan (CS-PG), 252 Chromium (Cr), 34, 76 Ciliary neurotrophic factor (CNTF), 439, 441 Citalopram-loaded gelatin nanocarriers (COMP), 282, 285 Cluster of differentiation (CD), 491 Cobalt (Co), 21, 35, 137, 327 Cold-gas spraying method (CGSM), 243 Collagen (Col), 37, 46, 71, 73, 74, 92, 94, 95, 97, 101, 106, 120, 122, 131–133, 138, 141, 143–145, 166–168, 171, 172, 175, 177–180, 196, 203, 219, 222, 223, 225– 228, 233, 234, 237, 239, 241, 246, 248, 251, 252, 254, 274, 276, 278, 281, 284– 286, 291, 292, 295–297, 314, 316, 317, 321, 323, 327, 329, 331, 332, 348, 353, 354, 356, 358, 360–363, 366, 367, 370, 371, 383, 386, 388, 389, 391–396,

Index 400–407, 409, 411, 412, 426, 427, 429, 431, 433, 437–441, 445–447, 449, 451– 455, 473, 475–477, 488, 490, 508–511, 515, 516, 520, 521, 523, 524, 569–573, 576 Collagen-glycosaminoglycan matrix (CGA), 433, 439 Computed tomography (CT), 44, 457 Computer aided design (CAD), 526, 566 Conducting polymers (CPs), 312–317, 442, 443 Controlled-atmosphere plasma spraying (CAPS), 243 Copper (Cu), 24 CRISPR, 494, 497 Cytocompatibility, 36, 43, 44, 46, 48, 49, 74, 97, 104, 133, 194, 280, 282, 351, 353, 354, 360, 393, 443, 514 D Decellularized extracellular matrix (DECM), 168–177, 405, 410–412, 477, 507, 564 Deferoxamine (DCF), 293 Demineralized bone matrix (DBM), 350, 397 Demineralized dentin matrix (DDM), 362 Dendrimers, 288, 291–292 Dental pulp stem cells (DPSCs), 362, 364, 440, 519, 521 Deoxyribonucleic acid (DNA), 29, 92, 171, 172, 179, 187, 289, 295, 314, 321, 434, 445, 492, 493, 495, 498, 512, 513, 546, 567 Dermal fibroblasts (DDF), 325, 329, 334, 543 Detonation-gun spraying (D-GUN), 243 Diamond-like carbon (DLC), 326 Diclofenac (DFO), 293 Dielectric, 320, 321, 333 Dimethylaminoethyl methacrylate (DMAEMA), 334, 335, 546 (3-(4,5-Dimethylthiazol-2-yl)-2,5Diphenyltetrazolium Bromide) (MTT), 27 Direct metal laser sintering (DMLS), 37 Dorsal root ganglion (DRG), 286, 444, 447 Drug eluting absorbable metal scaffold (DREAMS), 47 E Elastin-like polypeptides (ELPs), 233, 275 Electrical stimulation (ES), 9, 136, 309, 311, 313–317, 320, 324, 325, 327–330, 336, 369, 442–444, 448, 449, 549, 550

583 Electrophoretic deposition (EPD), 222, 243 Electroslag remelting (ESR), 34 Electrospinning, 99, 101–105, 109, 111, 127–129, 131, 133, 146, 147, 196, 274, 279, 282–284, 286, 315, 316, 319, 325, 328, 329, 356, 357, 382, 384, 388, 391, 413, 429, 437, 441, 443, 444, 447, 507, 509, 513, 521, 545, 573, 574, 576 Electrostatic, 70, 180, 221, 241, 246, 276, 279, 320, 321, 331, 334, 395, 456, 546 Embryonic stem cells (ESCs), 74, 144, 187, 286, 439, 472 Enamel, 63, 142, 348, 358, 359, 361, 364, 366, 369, 520 Encapsulation, 97, 98, 106, 108, 196, 197, 231, 276, 291, 293, 294, 426, 427, 430, 439, 444, 451–454, 456, 458, 473, 475, 476, 507, 520, 569, 570 Endothelial Colony Forming Cells (ECFCs), 544 Ethylenediaminetetraacetic acid (EDTA), 172 Expanded polytetrafluoroethylene (ePTFE), 250, 283, 429 Extra cellular matrix (ECM), 15, 90, 91, 96, 102, 109–111, 150, 165–173, 175, 176, 179, 182, 185, 188, 202, 203, 219, 221, 223, 238, 240, 241, 248, 250–257, 274, 276, 277, 279, 281, 284, 285, 289, 296, 297, 308, 316, 322, 327, 333, 350, 353, 354, 382, 383, 389–391, 394, 397, 401, 403–405, 407, 412, 426–428, 431, 434, 435, 437, 438, 443, 446, 448, 451–456, 458, 470, 473, 474, 477, 486, 487, 490, 506, 507, 509, 511, 515, 519–523, 541, 542, 568–570, 572 Extracellular vehicles (EVs), 439 F Fabrications, 4, 9, 12, 16, 21, 22, 31, 35, 37, 47–49, 64, 71, 76, 90, 91, 93, 99, 101, 103, 106–111, 126–131, 133, 134, 137, 138, 140, 141, 143, 146–148, 165, 166, 174, 175, 180, 181, 185, 187–189, 195, 196, 217, 232, 236, 239, 246, 255, 256, 286, 292, 309, 316, 319, 320, 322, 325, 328, 332, 349, 353, 356, 357, 362, 365, 369, 384, 387, 388, 401, 413, 426, 428– 430, 432, 433, 436–438, 441, 443–445, 447, 450, 453, 454, 457, 458, 473, 475, 477, 478, 487, 491, 499, 500, 506, 507, 509, 525, 526, 545, 564, 572, 574, 575 Fetal pulmonary cells (FPC), 235

584 Fibroblast growth factor (FGF), 31, 248, 278, 408 Fibroblast growth factor 1 (FGF1), 440, 453 Fibroblast growth factor 2 (FGF-2), 194, 363, 520 Fibronectin (FN), 143, 181, 230, 241, 248, 250–252, 274, 276, 348, 431, 434, 438, 444, 446, 447, 454, 455, 524 5-fluorouracil (5-FU), 323 Food and drug administration (FDA), 4, 14, 15, 99, 100, 103, 141, 146, 149, 176, 182, 195, 196, 337, 433, 509 Four dimensional (4D), 111, 127, 369, 537, 549 Fused Deposition Modeling (FDM), 129, 365, 565 G Gelatin (Gel), 31, 95, 97, 102, 104, 106, 111, 131, 144, 145, 177, 178, 180–181, 184, 194, 203, 219, 223, 225, 227, 228, 230, 231, 233–235, 238, 239, 241, 248, 253, 254, 282, 285, 286, 329, 353, 354, 368, 383, 389, 391–394, 431, 433, 473, 477, 508–510, 515, 523, 524, 542, 543, 546, 569–573, 576 Gelatin methacrylamide/methacryloyl/ methacrylatedGelatin (GelMA), 143, 181, 188, 238, 328, 392, 439, 444, 447, 538, 543, 544, 569, 573 Gene therapies, 14, 15, 123, 166, 486, 488, 492–500, 546 Gingival mesenchymal stem cells (GMSCs), 447 Glial cell line-derived neurotrophic factor/Glial cell-derived neurotrophic factor (GDNF), 278, 433, 439–441 Glycosaminoglycan (GAG), 138, 172, 186, 254, 298, 397, 399, 402, 454, 569 Gold nanoparticles (GNPs), 132, 289, 319, 360, 442, 521 Good manufacturing practices (GMP), 109 Granulocyte-macrophage colony-stimulating factor (GMCSF), 278 Graphene (G), 103, 132, 133, 185, 286, 315, 319, 326, 328, 329, 365, 433, 442, 444, 445, 447, 449, 474, 508, 514, 515, 517, 518, 520, 521 Graphene-based nanocomposites (GNs), 445 Graphene oxide (GO), 133, 315, 319, 329, 445, 450, 515, 517 Growth associated protein-43 (GAP-43), 443, 450

Index H Heparin sulfate proteoglycans (HS-PG), 252 1, 1, 1, 3, 3, 3-hexafluoro-2-propanol (HFP), 274 High nitrogen steel (HNS), 34, 35, 45 High velocity oxy-fuel spraying (HVOF), 243 Human adipose-derived stem cells (hASC), 99, 286, 365 Human aortic smooth muscle cells (HASMC), 100 Human dental pulp stem cells (hDPSCs), 360, 362, 366, 368 Human embryonic palatal mesenchymal (HEPM), 95 Human nasal inferior turbinate tissue-derived mesenchymal stromal cells/Human Tonsil-derived stem cells (hTMSCs), 367, 544 Human umbilical endothelial cells/Human umbilical vein endothelial cells (HUVEC), 34, 35, 476, 572–574 Hyaluronans, 98, 138, 150, 185, 199, 225, 251, 252, 278, 401 Hyaluronic acid (HA), 31, 38, 42, 72, 75, 97– 99, 111, 131, 138, 166, 167, 185, 186, 191–193, 198, 203, 229, 237, 238, 246, 251, 254, 256, 282, 286, 288, 297, 319, 321, 322, 330, 353, 354, 356, 360, 363, 367, 383, 393, 397, 403, 405, 406, 409, 412, 431, 435, 447, 449, 488, 490, 493, 510, 512, 518, 523, 542, 569 Hyaluronic Acid-g-Poly(2-hydroxyethyl methacrylate) (HEMA), 254 Hydrogels, 31, 45, 95, 97–99, 102–106, 111, 129, 131, 132, 138, 140, 141, 143–145, 147, 149, 150, 174, 176, 180, 182, 184– 188, 194, 197, 198, 200, 223, 225, 227– 232, 234, 237–239, 250, 251, 253, 254, 274, 275, 279, 281, 282, 285–288, 294– 296, 314, 317, 327, 328, 331–336, 352– 356, 361–368, 371, 382, 384, 388, 391, 392, 396–398, 401–405, 409, 410, 413, 427, 431–435, 437, 439, 444, 446–454, 458, 473–476, 488, 490, 491, 499, 500, 507, 510–513, 521, 524, 542, 544–549, 551–553, 563, 565, 566, 568, 569, 571– 573, 575 Hydrophobic, 124, 145, 241, 281, 283, 289, 331, 392, 438, 453, 513, 540, 541, 552 Hydroxyapatites (HAp/HA), 25, 31, 44, 63–68, 71, 74, 75, 99, 124, 125, 131, 132, 137, 140–142, 188, 189, 196, 200, 222, 223, 237, 238, 291, 292, 294–296, 319, 321, 322, 330, 332, 348, 351, 352, 354, 356,

Index 360, 362, 363, 366, 368, 473, 474, 510, 515, 544, 572, 573 3-Hydroxybutyrate-co-3-hydroxyvalerate (PHBV), 74, 104, 191, 195, 196, 449 4-Hydroxybutyrate/poly 4-hydroxybutyrate (P4HB), 191, 195 2-Hydroxyethyl methacrylate (HEMA), 143, 317, 335 Hydroxyethyl Methacrylate-Alginate-Gelatin (HAG), 255 I Immunocompatablity, 506 Implantation, 7, 9, 11, 15, 21, 22, 29, 41, 44, 46, 47, 64, 72, 74, 90, 91, 97, 100, 101, 104, 105, 144, 146, 174, 195, 200, 219, 229, 243, 244, 273, 279, 284, 292, 294, 322, 324, 335, 353, 359, 365, 366, 388, 389, 398, 399, 402, 403, 405, 406, 409–412, 429, 444–446, 451, 454, 507, 509, 515, 520, 522 Induced pluripotent stem cells (iPSC), 9, 10, 439, 470–472, 474, 475, 490–491, 494, 498, 542 Instantaneous blood-mediated inflammatory reaction (IBMIR), 456 Insulin-like growth factor (IGF), 67, 144, 226, 408, 433 Insulin-like growth factor-1(IGF-1), 144, 226, 362 Interferon γ (IFN γ), 278 Interleukin-8 (IL-8), 547 Interpenetrating network (IPN), 104 Iron (Fe), 24, 27, 32, 40–43, 47, 149, 280, 327, 332, 450, 548, 549 K Keratose (KOS), 437 L Laminin (LN), 106, 241, 248, 251, 252, 276, 386, 429, 431, 438–440, 446, 451, 452, 524 Laser engineered net shaping (LENS), 139, 145, 256 Layer-by-layer (LbL), 39, 176, 239, 246, 256, 276, 293, 443, 445, 457, 553, 564 Lead Zirconatetitanate (PZT), 320 Leu-Asp-Val (LDV), 184 Linear ordered collagen fibrous scaffold (LOCS), 446 Liposomes, 288–290, 294, 493, 546

585 Liquid crystalline elastomers (LCEs), 333, 538 Lower critical solution temperature (LCST), 281, 331, 540, 541 M Magnesium (Mg), 24, 29, 33, 38, 40, 42, 189, 222, 367 Magnetic field (MF), 32, 150, 327, 330, 332, 368, 450, 451, 547, 548 Magnetic nanoparticles (MNPs), 132, 327, 330–332, 450, 451, 515, 538, 548, 549 Magnetic Resonance Imaging (MRI), 24, 25, 32–33, 40, 44, 397, 549 Magnetic scaffold/stimulation (MS), 309, 327, 330, 451 Matrix metalloproteases (MMP), 274, 329, 403 Mesanchymal stem cells (MSC), 294, 298, 358, 470–472, 474, 476, 545 Methacrylated-Hyluronic Acid (MeHA), 544 Methacryloyl-substituted tropoelastin (MeTro), 275 Microporous annealed particle (MAP), 336 Molybdenum (Mo), 24, 34, 38 Mouse embryonic fibroblasts (MEF), 448 Multi walled carbon nano tubes (MWCNTs), 328, 329, 444, 450, 519 Myocardial infarction, 133, 144, 251, 275, 284, 289, 547 N Nano bioglass (nBG), 99, 434 Nanobiomaterials, 327, 328, 336, 506–519, 521–526 Nanocomposites, 102, 131, 133, 141, 202, 250, 254, 282, 289, 296, 322, 332, 363, 367, 368, 437, 450, 451, 516 Nanocrystalline diamond (NCD), 326 Nanofibres, 144, 507–511, 513, 514 Nano-hydroxyapatite (n-HA), 38, 120, 141, 358 Nanoporous cellulose gels (NCG), 450 Nanoscaffolds, 445, 514 Nano silver (nAg), 434 Nanotechnologies, 13, 90, 109, 179, 250, 257, 289, 431, 506–508, 525 Near Infrared (NIR), 325, 543 Neovascularization, 275, 294, 362, 391, 522 Nerve growth factor (NGF), 278, 328, 434, 436, 439–441, 443, 445, 447, 450, 451 Nerve-guidance-channels/conduits (NGC), 145, 148, 428–431, 433, 435–441, 444– 446, 451

586 Neural progenitor cells (NPC), 282, 449, 472, 513 Neural stem cells (NSCs), 286, 328, 439–441, 443, 444, 447, 449 Neurotrophic factors (NTFs), 200, 429, 431, 433, 434, 436, 439, 441, 448, 511 Neurotrophin-3 (NT-3), 441, 446 Nickel (Ni), 25, 34, 286, 327, 539 Nitrobenzene (NB), 334 N,N0 -dioctyl-3,4,9,10-perylenedicarboximide (PTCDI-C8), 325 Noncollagenous proteins, 348 Non-mulberry silk fibroin (NMSF), 182 Non-steroidal anti-inflammatory drugs (NSAIDs), 404 O Odontoblasts, 348, 361, 364, 520, 521 Olfactory ensheathing cells (OECs), 196, 439 Oncolytic viruses (OVs), 493 One dimensional (1D), 326 Osteoblasts, 26, 27, 29, 37, 38, 44, 45, 67, 106, 132, 141, 202, 242, 277, 279, 295, 322, 325, 332, 348, 350, 355, 360, 364, 367, 395, 471, 475, 507, 518, 520 Osteoconduction, 63 Osteogenesis, 24–26, 32, 35, 37–39, 42, 44, 48, 63, 102, 175, 237, 238, 248, 286, 350, 365, 474, 475, 518–520, 571 Osteogenic peptide (OP), 366 Osteoinduction, 31, 42, 46, 63, 70, 73, 189, 296, 358, 516 P Peptide amphiphiles (Pas), 334, 513 Periodontal ligament stem cells (PDLSCs), 75, 447, 520 Peripheral nervous system (PNS), 145, 281, 427, 440, 511, 513 PGA tube (PGAt), 440 Phenyl-C61-butyric acid methyl ester (PCBM), 324, 325 Phosphate glass microfibers (PGFs), 445 Photovoltaics, 312, 321–325, 336, 538, 549, 550 Physical vapor deposition (PVD), 222, 242 Piezoelectricity, 314, 318 Plasma-enhanced chemical vapor deposition (PECVD), 245 Platelet-derived growth factor (PDGF), 248, 362, 392

Index Pluripotent, 295, 364, 449, 471, 472, 475, 477, 490, 516 Poly(γ-glutamic acid) (γ-PGA), 191 Poly(ε-caprolactone) (PCL), 254, 431 Poly(ε-L-lysine) (ε-PL), 191 Polyacrylamide (PAM), 450, 490, 546 Poly(acrylic acid) (PAA), 233 Polyacrylonitrile (PAN), 128, 286, 316, 328 Polyamidoamine (PMAM), 291 Polyaniline (PANI), 144, 254, 314–317, 442, 443, 449 Polycaprolactone (PCL), 42, 46, 93, 99, 101–104, 106, 110, 111, 125, 128, 131, 132, 138, 144, 145, 147–149, 166, 198, 222, 233, 274, 280, 283, 292, 293, 296, 315, 325, 329, 353, 355, 361, 365– 367, 384, 388, 389, 393, 394, 399, 402, 403, 406, 410, 411, 436–438, 442, 445, 453, 475, 508, 510, 513, 514, 521, 523, 545 Poly(caprolactone-co-lactide) (PCLA), 431 Polycaprolactone fumarate (PCLF), 328, 436, 441 Polydimethylsiloxane (PDMS), 328, 452 Poly(D)L-lactide-co-ε-caprolactone) (PLCL), 431, 440 Poly dopamine (PDA), 239 Poly electrolyte multilayer (PEM), 246, 276 Poly(ether ether ketone) (PEEK), 111 Poly(3)4-ethylenedioxythiophene) (PEDOT), 314–316, 443, 449, 450 Polyethylene glycol/Poly(ethylene glycol) (PEG), 31, 99, 100, 102, 105, 106, 111, 125, 138, 143, 196, 231, 235, 237, 250, 252, 274, 285, 334–336, 355, 363, 367, 437, 448–450, 453, 454, 457, 473–477, 490, 493, 514, 520, 541, 547, 552 Poly(ethylene glycol)-co-poly(glycerol sebacate) (PEGS), 106, 317 Poly(ethylene glycol) methacrylate (PEGMA), 296 Poly(ethyleneglycol)-poly(D)L-lactide)/poly (D) L-lactic acid) (PDLLA), 100, 128, 278, 442 Polyethylene oxide (PEO), 105, 111, 147, 250, 331, 508 Poly(ethylene oxide)-poly(propylene oxide)poly (ethylene oxide) (PEO–PPO–PEO), 331 Polyethylene terephthalate (PET), 140, 143, 227, 254, 283, 508 Polyethylenimine(PEI), 251, 493, 547 Poly(glycerol sebacate) (PGS), 254, 399

Index Poly(glycolic acid)/Poly (glycolic acid)/ Polyglycolic acid/Polyglycolide (PGA) (PGA), 93, 94, 99, 101, 105, 123–125, 145, 148–150, 192, 194, 195, 235, 254, 274, 281, 284, 353, 360, 363, 389, 406, 407, 431, 435, 436, 440, 453, 508, 524 Polyhedral-oligomeric silsesquioxane-poly (carbonate-urea) urethane (POSS-PCU), 250 Poly(3-hexylthiophene) (P3HT), 325 Polyhydroxyalkanoates (PHA), 66, 93, 149, 191, 192, 195–196 Polyhydroxy butyrate/Poly-3-hydroxybutyrate (PHB), 131, 191, 192, 195, 196, 240, 434, 437 Poly(hydroxybutyrate)/chitosan (PHB/CTS), 282 Poly(3-hydroxybutyrate-co-3hydroxyvalerate)/Poly (3-hydroxybutyric acid-co-3-hydroxy valericacid)/Poly (3-hydroxy butyrateco-3- hydroxyvalerate) (PHBV), 105, 319, 437 Poly(3-hydroxydodecanoate) (P(3HDD)), 191 Poly (2-hydroxyethyl methacrylate) (PHEMA), 140, 143, 254, 317, 335, 473, 499 Poly (lactic acid)/Poly(lactide)/Polylactic acid/ Polylactide (PLA), 71, 74, 93, 94, 97–106, 111, 123–125, 128, 131, 135, 138, 139, 144, 145, 148–150, 166, 175, 222, 223, 226, 230–233, 235, 236, 240, 247, 250, 253, 254, 274, 286, 294, 314–317, 319, 325, 329, 331–333, 353–356, 360, 363, 366–369, 384, 393, 399, 403, 406, 407, 431, 436, 437, 440, 442, 443, 448, 449, 453, 473, 474, 508, 513, 537, 540, 541, 545–547, 551–554, 572 Poly(lactic acid-co-glycolic acid)/poly(lacticco-glycolic acid)/poly(lactic-coglycolic)/poly(glycolic-lactic acid)/ polylactic-coglycolic acid/polylactide co-glycolide (PLGA), 93, 99, 101–102, 111, 138, 144, 145, 147, 148, 175, 223, 226, 229, 231, 232, 235, 236, 254, 274, 278, 284, 322, 353, 355, 363, 367, 384, 406, 411, 412, 431, 436, 439, 441, 442, 446, 448, 473, 476, 508, 515, 517, 521 Poly(lactic-co-glycolic acid)/hyaluronic acid (PLGA/HA), 335 Poly(l/d-lactic acid) (PLDLA), 445 Poly(L-lactic acid)/poly (lactic-co-glycolic acid)/poly-L-lactic acid (PLLA), 99,

587 100, 131, 140, 145, 147, 148, 233, 236, 237, 284, 286, 315, 317, 388, 399, 402, 431, 436, 438, 445, 451, 476, 508, 515, 516 Poly l-lactic acid-co-poly-caprolactone (PLLAPCL), 508 Poly(L-lactic acid)–co–poly(ε-caprolactone) (PLPCL) smooth muscle cells (SMCs), 98, 101, 386–389, 403, 574 Poly (L-lactic acid)-poly(ethylene glycol)-poly (L-lactic acid) (PLLA-PEG-PLLA), 331 Poly(L-lactide-co-ɛ-caprolactone) (PLLA-CL), 431 Poly(L-lysine)(PLL), 493 Poly(2-methoxy-5-(2-ethylhexyloxy)-1)4phenylenevinylene) (MEH-PPV), 314, 315 Polymethyl methacrylate/poly(methyl methacrylate) (PMMA), 139, 140, 145, 222, 508 Poly(N-isopropylacrylamide) (PNiPAAm), 331, 540 Poly N-Isopropylacrylamide-coGlycidylmethacrylate (NGMA), 541, 542 Poly norepinephrine (pNE), 410 Poly(N-vinylcaprolactam) (PNVC), 331 Poly(1)8-octanediol citrate) (POC), 250 Poly(p-phenylene vinylene) (PPV), 449 Polypropylene (PP), 135, 202, 250, 323 Polypropylene sulphide/poly(propylene sulfide) (PPS), 250, 552 Polypyrrole (PPy), 102, 132, 144, 314–317, 442, 445, 446, 449, 450, 550 Poly ((R)-3-hydroxybutyrate) (PHB), 191, 192, 194–196, 200, 274, 277–279, 282, 288, 292–294, 476, 493, 499, 508, 515, 519–521, 523, 524, 569 Poly(styrenesulfonate)/poly(4-styrene sulfonate) (PSS), 443 Polytetrafluoroethylene (PTFE), 123, 124, 139, 143, 227, 322 Polythiophene (PT), 144, 314, 442, 449 Poly urethane (PU), 7, 132, 133, 143–145, 223, 227, 247, 251, 254, 274, 279, 283, 284, 286, 292, 298, 369, 437, 443, 508, 538, 573 Poly(vinyl alcohol)/Polyvinyl alcohol (PVA), 76, 95, 99, 103–105, 126, 128, 131, 132, 135, 141, 166, 231, 254, 329, 393, 410, 431, 436, 437, 443, 447–449, 453, 508 Polyvinylidene fluoride (PVDF), 319, 448, 449

588 Polyvinylidene fluoride-trifluoroethylene (PVDF-TrFE), 369 (POxs) Poly(2-oxazoline)s, 331 Proteoglycans (PGs), 90, 109, 278, 315, 386, 394, 395, 519, 523, 524 Proximal tubule cells (PCT), 240 R Rat aortic smooth muscle cells (RASMC), 96 Reactive Oxygen Species(ROS), 26, 445, 551, 552 Recellularization, 143, 169, 170, 173, 175, 234 Recombinant DNA technology (RDT), 168, 492 Recombinant human bone morphogenetic protein 2 (rhBMP-2), 194, 519 Reduced graphene oxide (rGO), 329 Regenerative Medicine (RM), 4, 7, 9–16, 99, 103, 166, 167, 169, 185, 189, 191, 196, 219, 247, 251, 253, 257, 271–299, 309, 365, 427, 454, 456, 470, 472–478, 486, 488–490, 497–500, 506, 508–516, 525, 544, 549, 561–576 Rejuvenation, 176, 486 Resorption, 44, 63, 68, 71, 73, 74, 98, 124, 236, 321, 336, 360, 517 Retinal ganglion cells (RGCs), 315 Rheumatoid arthritis (RA), 491, 518 S Salidroside (SDS), 172, 439 Scaffolds, 4, 5, 9–12, 15, 16, 22, 25, 28, 31, 37, 38, 42, 43, 46–49, 62, 64–66, 69–76, 90–112, 120–122, 124, 125, 127–130, 132, 133, 136, 138, 141, 143–150, 165–169, 171, 172, 175, 177–182, 184–187, 193–198, 200, 202, 217–219, 221–223, 225–227, 229, 230, 232–241, 246–248, 250–254, 273–275, 277, 278, 280–288, 290–298, 308, 311, 315, 316, 319, 324, 327–332, 334–336, 348–353, 355–363, 365–371, 382–385, 388, 389, 391, 393, 394, 396–399, 401–407, 409–413, 426–429, 431, 433–455, 457, 458, 470, 473–478, 487, 488, 499, 500, 506, 507, 509–525, 541, 543–545, 547, 548, 552, 553, 562, 565, 567, 568, 574 Schwann cells (SC), 195, 282, 286, 427, 429, 433–435, 438, 439, 444, 445, 513 Shape memory materials (SMMs), 539 Short interfering RNA (SiRNA), 547 Silicon (Si), 124, 142, 145, 148, 246, 324, 325, 367, 429, 446

Index Silk fibroin (SF), 92, 99, 120, 138, 182, 188, 200, 225, 251, 274, 283, 284, 316, 322, 328, 329, 354, 406, 431, 434, 441–443, 448, 452, 454, 509, 510, 515, 518, 523, 572, 576 Silk sericin (SS), 182 Silver (Ag), 24, 36, 132, 239, 247, 289, 394, 515, 518 Single-walled carbon nanotube (SWNT), 437 Small intestine submucosa (SIS), 446 Sodium alginate (SA), 39, 42, 97, 367, 368, 443, 450, 572 Somatic lung progenitor cells (SLPC), 235 Spinal cord injury (SCI), 196, 282, 440, 441, 446, 448 Stem cell factor (SCF), 278 Stereolithography (SLA), 110, 111, 127, 129, 450, 564, 565 Stromal cell-derived factor-1/Stromal derived factor 1 (SDF-1), 275, 298, 409 Superparamagnetic iron oxide nanoparticles (SPIONs), 369, 450 Synovium-derived mesenchymal stem cells (SMSC), 285 T Tendon stem cells (TSCs), 405 Thermally induced phase separation (TIPS), 238, 440 Three dimensional (3D), 282 Thrombomodulin (TM), 456 Tissue engineering (TE), 4, 5, 7–16, 19–49, 61–77, 89–112, 119–151, 163–203, 215–257, 272–299, 305–337, 347–371, 381–413, 423–458, 469–478, 486–489, 497–500, 505–526, 537, 538, 540, 542–550, 552–554, 562–576 Titanium dioxide (TiO2), 202, 222, 515, 518 Tobacco mosaic virus (TMV), 443 Totipotent, 471 Transforming growth factor β3 (TGF β3), 186, 397 Tricalcium phosphate (TCP), 64, 72–75, 123, 124, 129, 140, 196, 322, 360, 362, 366, 572 Trifluoro ethylene (TrFE), 319 Tumor necrosis factor-α (TNF-α), 408 Two dimensional (2D), 284 U Ultrasound (US), 382, 436, 446, 537, 538, 553 Ultraviolet (UV), 135, 245, 247, 328, 333, 334, 405, 453, 543–545, 564, 567, 571, 573

Index

589

Umbilical cord-derived mesenchymal stromal cells (UCMSCs), 440 Upper critical solution temperature (UCST), 331, 540 Urinary bladder matrix (UBM), 446, 451 U.S. Food and Drugs Administration (US FDA), 4, 521

362, 391, 392, 429, 434, 441, 442, 451, 454, 510, 520, 572 Vitronectin, 241, 250, 277

V Vascular endothelial growth factor (VEGF), 31, 34, 181, 225, 236, 241, 248, 278, 280,

Z Zero dimensional (0D), 326

X Xenografts, 6, 360, 382, 428