This book comprehensively explores the basic concepts and applications of biomaterials in tissue engineering and regener
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English Pages 598 [587] Year 2021
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
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)
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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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13
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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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)
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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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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
2
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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 (PO34), metaphosphate (PO3), or pyrophosphate (P2O47) 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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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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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
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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
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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)
• Difficul