Materials for Biomedical Engineering: Thermoset and Thermoplastic Polymers 0128168749, 9780128168745

Table of contents :
Cover
Thermoset and Thermoplastic Polymers
Copyright
List of Contributors
Series Preface
Preface
1 Introduction in thermoplastic and thermosetting polymers
1.1 Introduction
1.2 Biomedical Polymers
1.3 Thermoplastic and Thermosetting Polymers
1.3.1 Thermoplastics
1.3.2 Thermosets
1.4 Biomedical Thermoplastic and Thermosetting Polymers
1.4.1 Polyethylene Glycol
1.4.2 Polyvinyl Alcohol
1.4.3 Chitosan
1.4.4 Shape Memory Polymers
1.5 Conclusion
References
2 Laser surface texturing of thermoplastics to improve biological performance
2.1 Introduction
2.2 Impact of Roughness and Wettability on Biocompatibility
2.3 Surface Engineering Processes
2.3.1 Surface Roughening
2.3.2 Surface Chemical Modification
2.4 Basics of Laser Surface Texturing
2.4.1 Introduction
2.4.2 Process Fundamentals
2.4.3 Components of a Laser Texturing Setup
2.4.4 Processing Parameters
2.4.4.1 Wavelength of the laser radiation
2.4.4.2 Beam mode (continuous-wave versus pulsed lasers)
2.4.4.3 Pulse length
2.4.4.4 Pulse energy
2.5 Laser Surface Texturing of Thermoplastic Polymers
2.5.1 Poly(Etheretherketone)
2.5.2 Polycarbonate
2.5.3 Polypropylene
2.5.4 Polyethylene
2.5.5 Poly(Ethylene Terephthalate)
2.5.6 Ultra-High-Molecular-Weight Polyethylene
2.5.7 Poly(Methyl Methacrylate)
2.5.8 Other Thermoplastic Polymers
2.6 Challenges and Future Trends
2.7 Conclusions
Acknowledgments
References
3 Light-mediated thermoset polymers
3.1 Introduction
3.2 Types of Light-Sensitive Polymers
3.2.1 Polyurethanes
3.2.1.1 Synthetic routes to photocrosslinkable polyurethane polymers
3.2.1.2 Radiation curing in surface modification
3.2.1.3 Shape memory polymers
3.2.2 Poly-acrylates/methacrylates
3.2.2.1 As dental materials
3.2.2.2 As hydrogel biomaterials
3.2.3 Light-Sensitive Vinyl Monomers
3.2.3.1 N-Vinyl pyrrrolidone
3.2.3.2 Vinyl carbonate
3.2.3.3 Vinyl esters
3.3 Photoinitiators
3.3.1 Irgacure 2959
3.3.2 2,4,6-Trimethyl benzoyl-Diphenyl phosphine Oxide
3.3.3 Camphorquinone with Amine Photoinitiator System
3.4 Mechanisms of Light Sensitization
3.5 Polyacrylates for Biopolymer Applications
3.5.1 Polycaprolactone
3.5.2 Starch
3.5.3 Dextran
3.5.4 Gelatin
3.5.5 Chitosan
3.5.6 Hyaluronic Acid
3.6 Recent Advancements and Trends in Light-Mediated Polymerizations
3.7 Conclusion
References
4 Thermoset, bioactive, metal–polymer composites for medical applications
4.1 Thermosetting Polymers
4.1.1 Introduction
4.1.2 Synthesis of Thermoset Polymers
4.1.2.1 Synthesis of thermosetting polymers by polymerization
4.1.2.2 Synthesis of thermosetting polymers by crosslinking or curing
4.1.3 Properties of Thermosetting Polymers
4.1.3.1 Formulations
4.1.3.2 Solvent resistant
4.1.3.3 Melt viscosity
4.1.3.4 Mechanical properties
4.1.3.5 Fiber impregnation
4.1.3.6 Processing cycle
4.1.4 Characterization of Thermoset Polymers
4.1.4.1 Fourier transform infrared spectroscopy
4.1.4.2 Nuclear magnetic resonance spectroscopy
4.1.4.3 Differential scanning colorimetry
4.1.4.4 Thermogravimetric analysis
4.1.4.5 Dynamic mechanical thermal analysis
4.1.4.6 X-ray fluorescence spectroscopy
4.1.5 Applications of Thermoset Polymers
4.1.5.1 Urea–formaldehyde resin
4.1.5.2 Melamine–formaldehyde resins
4.1.5.3 Phenol–formaldehyde resin
4.1.5.4 Polyelectrolytes
4.1.5.5 Polyurethane
4.1.5.6 Epoxy resins
4.1.5.7 Unsaturated polyester resin
4.2 Thermoset Metal–Polymer Composites
4.2.1 Introduction
4.2.2 Synthesis of Thermoset Composites
4.2.3 Properties of Thermoset Polymer Composites
4.2.3.1 Tensile strength
4.2.3.2 Fracture surface
4.2.3.3 Stress–strain behavior
4.2.3.4 Dynamic mechanical properties
4.2.3.5 Wear performance
4.2.4 Characterization of Thermoset Polymer Composite
4.2.5 Applications of Thermoset Polymer Composites
4.3 Applications in Biomedical Engineering
4.3.1 In Dentistry
4.3.2 In Prosthetic Heart Valves
4.3.3 In Bones
4.3.4 In Bone Grafting
4.3.5 In Prosthetic Sockets
4.3.6 In Medical Devices
References
Further Reading
5 Epoxy composites in biomedical engineering
5.1 Introduction
5.2 Artificial Implants and Bone Fixation Plates
5.2.1 Artificial Implants
5.2.2 Fixation Plates, Screws, and Intramedullary Nails
5.3 Tribological Characterization of Green Composites for Biomedical Applications
5.3.1 Bulk Composites
5.3.2 Composite Coatings
5.4 Dental Applications
5.5 Research Works Based on Bioepoxy Resins
5.6 Composite Shape Memory Polymers for Biomedical Applications
5.7 General Biomedical Applications
5.8 Summary of Research Works in Epoxy Composites for Biomedical Applications
5.9 Conclusions
References
6 Polyethylene and polypropylene matrix composites for biomedical applications
6.1 Introduction
6.2 Polyolefin Composites
6.3 Biomedical Engineering
6.4 Biocompatibility Evaluation of Polyolefin-Based Biocomposites
6.4.1 Tests for Biocompatibility
6.5 Fabrication Techniques for Polyolefin Biomedical Composites
6.5.1 Molding
6.5.2 Extrusion
6.5.3 Melt Electrospinning
6.5.4 Filament Winding
6.5.5 Thermoplastic Pultrusion
6.6 Polyethylene Matrix
6.6.1 HDPE-Based Biomedical Composites
6.6.2 UHMWPE-Based Biomedical Composites
6.7 Polypropylene Matrix
6.7.1 Finger Joint Implants
6.7.2 Bone Cement
6.7.3 Scaffolds
6.7.4 Antimicrobial Applications
6.7.5 Sutures
6.8 Conclusions
References
7 Polymethacrylates
7.1 Material Selection for Medical Applications: Requirements for Several Kinds of Medical Applications
7.2 Chemistry of Polymethacrylates and Their Composites
7.2.1 Monomers
7.2.1.1 Methyl methacrylate
7.2.1.2 Other methacrylates for dental applications
7.2.1.3 Composition of the matrix
Monomers
Activators and polymerization initiators
Polymerization inhibitors
Coupling agents
7.2.2 Dental Composites
7.2.2.1 Particle size and distribution of fillers
7.2.2.2 Viscosity
7.2.2.3 Polymerization mode
7.2.3 Challenges in Improving Properties
7.3 Methods for Material Synthesis
7.3.1 Radical Polymerization Reaction of PMMA (Difunctionnal Monomer)
7.3.1.1 Mechanistic aspects
7.3.1.2 Kinetic aspects
7.3.2 Polymerization of Methacrylate Networks
7.3.2.1 Mechanistic aspects
7.3.2.2 Polymerization kinetics
7.3.3 Parameters Influencing Polymerization
7.3.3.1 Intrinsic factors
7.3.3.2 Extrinsic factors
7.3.4 Polymerization Shrinkage and its Consequences
7.4 Physicochemical, Biological and Mechanical Properties
7.4.1 Structure–Properties Relationships and Link With Clinical Applications
7.4.1.1 Glass transition temperature and other transitions
7.4.1.2 Short deformation properties
7.4.1.3 Ultimate properties
7.4.2 Biocompatibility
7.5 Long-Term Behavior
7.5.1 Aging by Physical Relaxation
7.5.2 Humid Ageing
7.5.2.1 Water solubility
7.5.2.2 Water diffusion
7.5.2.3 Consequences of physical ageing on mechanical properties
7.5.2.4 Role of the interface
7.5.2.5 Effect of penetrant composition mixture
7.5.3 Chemical Ageing by Hydrolysis
7.5.4 Chemical Ageing by Radiolysis
7.5.5 Creep and Fatigue
7.6 Conclusion and Prospects for the Future of These Materials
References
8 Thermoset polymethacrylate-based materials for dental applications
8.1 Introduction
8.1.1 Gold
8.1.2 Porcelain
8.1.3 Vulcanite
8.1.4 Aluminum
8.1.5 Celluloid
8.1.6 Bakelite
8.1.7 Polyvinyl Chloride
8.1.8 Base Metal Alloys
8.2 Poly(Methyl Methacrylate) as a Denture Base
8.2.1 Classification of PMMA Resins
8.2.1.1 According to the ISO standards
8.2.1.2 According to method of polymerization
Heat cured PMMA
Polymerization stages of heat cured PMMA
Initiation and activation
Propagation
Termination
The sandy stage
The stringy stage
The doughy stage
The rubbery stage
The stiff stage
The compression molding technique
The injection molding technique
Polymerization cycles
Chemically cured PMMA
Light cured PMMA
Microwave curing PMMA
8.3 Properties of PMMA Denture Base Resins
8.3.1 Flexural Strength
8.3.2 Fracture Toughness
8.3.3 Impact Strength
8.3.4 Crosslinking
8.3.5 Sorption and Solubility
8.3.6 Thermal Conductivity
8.3.7 Residual Monomer
8.3.8 Color Stability
8.3.9 Radiopacity
8.3.10 Biocompatibility and Cytotoxicity
8.4 Contemporary Denture Base Materials and Modifications of PMMA
8.4.1 Polyamides
8.4.2 Epoxy Resins
8.4.3 Polycarbonates
8.5 Chemical Modification of PMMA
8.6 Reinforcement of PMMA Denture Base Materials
8.6.1 Reinforcement With Metal Wires or Mesh
8.6.2 Fiber Reinforcement
8.6.2.1 Effects of fiber length on properties of fiber reinforced denture base resins
8.6.2.2 Effect of fiber orientation
8.6.2.3 Effects of resin impregnation on PMMA resin-based materials
8.6.2.4 The effect of silane treatment on properties of PMMA denture base resins
8.6.3 Different Types of Fibers Used in Dentistry
8.6.3.1 Carbon fibers
8.6.3.2 Aramid fibers
8.6.3.3 Polyethylene (UHMWPE) fibers
8.6.3.4 Glass fibers
Types and composition of glass fibers
Properties of glass fiber reinforced denture base resins
8.7 Conclusion
List of Abbreviations
References
Further Reading
9 Maleic anhydride copolymers as a base for neoglycoconjugate synthesis for lectin binding
9.1 Introduction
9.2 Experimental
9.2.1 Materials
9.2.2 Instrumentation
9.2.3 Methods
9.2.3.1 Synthesis of N-glycyl-β-glycopyranosylamines
Synthesis of N-glycyl-2-actamido-2-deoxy-β-d-glucopyranosylamine (N-Gly-GlcNAc)
Synthesis of N-glycyl-4-O-β-d-galactopyranosyl-β-d-glucopyranosylamine (N-Gly-lactose)
9.2.3.2 Synthesis of neoglycoconjugates and glyconanoparticles
Carbohydrate-polymer ester bond formation: general procedure
Carbohydrate-polymer amide bond formation: general procedure
Synthesis of crosslinked glycoconjugates (CLGC-1 and -2, Scheme 9.3B): general procedure
Synthesis of silver, or gold glyconanoparticles (Scheme 9.1)
9.2.3.3 Lectins binding assays
Dot-blotting
UV-visible absorbance measurements
Binding properties of crosslinked neoglycoconjugate sorbents
9.3 Results and Discussion
9.3.1 Synthesis of Neoglycoconjugates and Metal-Labeled Glyconanoparticles
9.3.2 Characterization of Colloidal Neoglycoconjugates and Glyconanoparticles
9.3.3 Silver (or Gold)-Labeled Neoglycoconjugate: Lectin Interactions Study
9.3.3.1 Development of lectin sensors
9.3.3.2 UV-visible absorbance spectroscopy
9.3.4 Crosslinked Lectin Sorbents
9.4 Conclusions
Acknowledgment
References
Further Reading
10 Particulate systems of PLA and its copolymers
10.1 Introduction
10.2 Properties of Poly(Lactic Acid)
10.2.1 Production of Poly(Lactic Acid)
10.2.2 Unique Properties of Poly(Lactic Acid) and its Copolymers
10.2.3 Biocompatibility and Safety of Poly(Lactic Acid)
10.3 Micro- and Nanoparticulate Systems of Poly(Lactic Acid)
10.3.1 Preparation Methods of Poly(Lactic Acid) Micro- and Nanoparticles
10.3.1.1 Emulsion-based methods
Emulsification–solvent evaporation
Emulsification–solvent diffusion
Emulsification–reverse salting out
10.3.1.2 Nanoprecipitation method
10.3.1.3 Dialysis
10.3.1.4 Spray drying
10.3.1.5 In situ method for particle formation
10.3.1.6 Supercritical fluids technique
10.3.1.7 Particle formation using template/mold
10.3.1.8 Microfluidic technique
10.3.2 Challenges With Particulate System
10.4 Products Under Preclinical and Clinical Trial
10.5 Products Under Clinical Use
10.6 Advancements
10.6.1 Vaccination
10.6.2 Super Paramagnetic Iron Oxide Nanoparticles (SPIONS)
10.6.3 Cellular Interaction
10.6.4 Gene Transfection and Tissue Engineering
10.6.5 Dental Engineering
10.6.6 Active Targeting
10.6.7 Pheroid System
10.7 Conclusions
10.8 Future Perspectives
References
11 Polylactide: the polymer revolutionizing the biomedical field
11.1 Introduction
11.2 Polylactic Acid Synthesis
11.2.1 Precursors
11.2.1.1 Lactic acid
11.2.1.2 Lactide
11.2.2 Polylactic Acid Polymerization
11.2.2.1 Condensation and coupling of lactic acid
11.3 Polylactic Acid Modification
11.3.1 Modification by High Energy Radiations and Peroxides
11.3.2 Graft Copolymerization
11.4 Physicochemical Properties of Polylactic Acid
11.4.1 Rheological Properties
11.4.2 Mechanical Properties
11.4.3 Thermal Properties
11.4.4 Biodegradation Properties
11.5 Biomedical Applications of Polylactic Acid
11.5.1 Tissue Engineering
11.5.2 Drug Delivery With Polylactic Acid Particles
11.5.3 Vaccine Delivery
11.5.4 Tumor Treatment
11.5.5 Immunization With Polylactic Acid Particles
11.5.6 DNA and Gene Delivery
11.5.7 Antigen Loading
11.5.8 Protein Delivery
11.5.9 Imaging and Diagnosis
11.6 Conclusion
References
Further Reading
12 Poly(propylene fumarate)-based biocomposites for tissue engineering applications
12.1 Introduction
12.2 Poly(Propylene Fumarate): Synthesis, Properties, and Applications
12.2.1 Synthesis
12.2.2 Properties
12.2.3 Applications
12.3 Graphene Oxide: Structure, Synthesis, and Properties
12.3.1 Structure
12.3.2 Synthesis
12.3.3 Properties
12.4 Boron Nitride Nanotubes: Structure, Synthesis, and Properties
12.4.1 Structure
12.4.2 Synthesis
12.4.3 Properties
12.5 Preparation of PPF-Based Biocomposites
12.6 Characterization of PPF-Based Bionanocomposites
12.6.1 Morphology and Structure
12.6.2 Hydrophilicity, Biodegradability, and Protein Adsorption
12.6.3 Thermal Properties
12.6.4 Mechanical Properties
12.6.5 Antibacterial Properties
12.6.6 Cytotoxicity
12.6.7 Tribological Properties
12.7 Conclusion and Future Perspectives
Acknowledgement
References
13 Diblock and triblock copolymers of polylactide and polyglycolide
13.1 Introduction
13.1.1 History of Polylactide
13.1.2 History of Polyglycolide
13.1.3 Synthesis of Diblock and Triblock Copolymers of Polylactide and Polyglycolide
13.1.4 Characterization of Copolymers of Polylactide and Polyglycolide
13.1.4.1 Structural composition analysis
13.1.4.2 Aqueous solubility and injectability
13.1.4.3 Phase transition
13.1.4.4 Thermal properties
13.1.4.5 Crystallization behavior
13.1.4.6 Biocompatibility, cytotoxicity, and biodegradability
13.2 Resorbable Thermosensitive Polymers
13.2.1 Thermosensitive Polymer-Based Drug Delivery Systems
13.2.2 Commercial and Investigational Examples
13.2.3 Limitations of Thermosensitive Polymers
13.3 Resorbable Nanoparticles
13.3.1 Nanoparticle Preparation and Characterization Techniques
13.3.2 Resorbable Nanoparticles-Based Drug Delivery Systems
13.3.3 Commercial and Investigational Examples
13.3.4 Limitations of Resorbable Polymeric Nanoparticles
13.4 Conclusions and Future Perspectives
References
14 Characteristics of polymeric materials used in medicine
14.1 Introduction
14.2 Applications of Biomaterials
14.3 UHMWPE Behavior Under the Action of External Factors
14.4 Behavior of Medical Grade UHMWPE in Living Tissue
14.5 UHMWPE Versus Other Biomaterials
14.6 Background on Biopolymers in Living Tissue
14.7 Present and Future of Biopolymers, Bioplastics, and Nanobiomaterials
14.8 Conclusions
References
Further Reading
15 Application of polymethylmethacrylate, acrylic, and silicone in ophthalmology
15.1 Introduction
15.1.1 Silicone
15.1.2 Polymethylmethacrylate
15.1.2.1 Properties and advantages of polymethylmethacrylate
15.2 Application of Biomaterials in Intraocular Lenses
15.2.1 Lenses Used in Cataract Surgery
15.2.2 Phakic Lenses
15.2.3 Intraocular Lens Structure
15.2.4 Implant Positions for Intraocular Lenses
15.2.5 The Properties of Intraocular Lens Materials
15.2.5.1 Intraocular lens materials
Acrylic
Poly methyl methacrylate
Foldable hydrophobic acrylic
Foldable hydrophilic acrylic (hydrogel)
Silicone
Collamer
15.2.6 The Effect of Different Intraocular Lens Materials on Postoperative Complications
15.2.6.1 Posterior capsular opacification
15.2.6.2 Glistenings
15.2.6.3 Calcification
15.2.7 The Effect of Different Intraocular Lens Materials on the Quality of Vision in Pseudophakic Eyes
15.2.8 Different Intraocular Lenses Materials in Congenital Cataract Surgery in Children
15.2.9 Elimination of UV and Blue Rays From Intraocular Lenses
15.3 Artificial Cornea
15.3.1 History and Development of Keratoprosthesis
15.3.2 Boston Keratoprosthesis
15.3.2.1 Improvements over time
15.3.2.2 Outcomes of boston type-1 KPro
15.3.2.3 B-KPro type II
15.3.3 Osteo-Odonto-Keratoprosthesis
15.3.4 Cardona Keratoprosthesis
15.3.5 Pintucci Biointegrable Keratoprosthesis
15.3.6 KeraKlear (Keramed)
15.3.7 Moscow Eye Microsurgery Complex in Russia
15.3.8 What Next?
15.3.9 Recent Trends
15.4 Glaucoma Drainage Devices
15.4.1 Historical Perspective
15.4.2 Fundamental Principles of Glaucoma Drainage Devices
15.4.3 Types of Glaucoma Drainage Devices
15.4.3.1 Ahmed glaucoma valve
15.4.3.2 Baerveldt glaucoma implants
15.4.3.3 Molteno
15.4.3.4 Krupin slit valve
15.4.3.5 Ex-PRESS mini glaucoma shunt: Ex-PRESS glaucoma filtration device
15.5 Intracorneal Rings
15.5.1 Intacs Segments
15.5.2 Ferrara Ring Segments
15.5.3 Bisantis Intrastromal Segmented Perioptic Implants
15.5.4 MyoRing
15.5.5 KeraRing
References
Index
Back Cover

Citation preview

Thermoset and Thermoplastic Polymers

Materials for Biomedical Engineering

Thermoset and Thermoplastic Polymers

Edited by

Valentina Grumezescu Laser Department, National Institute for Laser Plasma & Radiation Physics, Romania

Alexandru Mihai Grumezescu Faculty of Applied Chemistry and Materials Science, University Politehnica of Bucharest, Bucharest, Romania

Elsevier Radarweg 29, PO Box 211, 1000 AE Amsterdam, Netherlands The Boulevard, Langford Lane, Kidlington, Oxford OX5 1GB, United Kingdom 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States Copyright © 2019 Elsevier Inc. All rights reserved. No part of this publication may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, recording, or any information storage and retrieval system, without permission in writing from the publisher. Details on how to seek permission, further information about the Publisher’s permissions policies and our arrangements with organizations such as the Copyright Clearance Center and the Copyright Licensing Agency, can be found at our website: www.elsevier.com/permissions. This book and the individual contributions contained in it are protected under copyright by the Publisher (other than as may be noted herein). Notices Knowledge and best practice in this field are constantly changing. As new research and experience broaden our understanding, changes in research methods, professional practices, or medical treatment may become necessary. Practitioners and researchers must always rely on their own experience and knowledge in evaluating and using any information, methods, compounds, or experiments described herein. In using such information or methods they should be mindful of their own safety and the safety of others, including parties for whom they have a professional responsibility. To the fullest extent of the law, neither the Publisher nor the authors, contributors, or editors, assume any liability for any injury and/or damage to persons or property as a matter of products liability, negligence or otherwise, or from any use or operation of any methods, products, instructions, or ideas contained in the material herein. British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress ISBN: 978-0-12-816874-5 For Information on all Elsevier publications visit our website at https://www.elsevier.com/books-and-journals

Publisher: Matthew Deans Acquisition Editor: Gwen Jones Editorial Project Manager: Emma Hayes Production Project Manager: Debasish Ghosh Cover Designer: Greg Harris Typeset by MPS Limited, Chennai, India

List of Contributors Hossein Aghamollaei Chemical Injuries Research Center, Systems biology and Poisonings Institute, Baqiyatallah University of Medical Sciences, Tehran, Iran Naveed Ahmed Department of Pharmacy, Quaid-i-Azam University, Islamabad, Pakistan Felipe Arias-Gonza´lez Applied Physics Department, University of Vigo, EEI, Lagoas-Marcosende, Vigo, Spain Sanjay Arora Department of Pharmaceutical Sciences, College of Health Professions, North Dakota State University, Fargo, ND, United States Muhammad Imran Asad Department of Pharmacy, Quaid-i-Azam University, Islamabad, Pakistan Mehmood Asghar National University of Medical Sciences (NUMS), The Mall, Rawalpindi, Pakistan Alexandra Bıˆrca˘ Department of Science and Engineering of Oxide Materials and Nanomaterials, Faculty of Applied Chemistry and Materials Science, University Politehnica of Bucharest, Bucharest, Romania; Faculty of Engineering in Foreign Languages, University Politehnica of Bucharest, Bucharest, Romania Satheesan Bobby Mechanical Engineering Department, King Fahd University of Petroleum and Minerals, Dhahran, Saudi Arabia Mohamed Boutinguiza Applied Physics Department, University of Vigo, EEI, Lagoas-Marcosende, Vigo, Spain Adolfo Chantada Applied Physics Department, University of Vigo, EEI, Lagoas-Marcosende, Vigo, Spain Murthy Chavali Shree Velagapudi Ramakrishna Memorial College, Acharya Nagarjuna University, Guntur, Andhra Pradesh, India; MCETRC, Tenali, Guntur, Andhra Pradesh, India Rafael Comesan˜a Materials Engineering, Applied Mechanics and Construction Department, University of Vigo, EEI, Lagoas-Marcosende, Vigo, Spain

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Ernesto David Davidson Hernandez Tecnicatura Superior Universitaria en Palenteologia, Universidad del Chubut, Rawson, Repu´blica Argentina Jesu´s del Val Applied Physics Department, University of Vigo, EEI, Lagoas-Marcosende, Vigo, Spain Ana M. Dı´ez-Pascual Analytical Chemistry, Physical Chemistry and Chemical Engineering Department, Faculty of Biology, Environmental Sciences and Chemistry, Alcala´ University, Madrid, Spain Shahab Ud Din Shaheed Zulfiqar Ali Bhutto Medical University/Pakistan Institute of Medical Sciences, Islamabad, Pakistan Oana Gherasim Department of Science and Engineering of Oxide Materials and Nanomaterials, Faculty of Applied Chemistry and Materials Science, University Politehnica of Bucharest, Bucharest, Romania; Lasers Department, National Institute for Lasers, Plasma and Radiation Physics, Magurele, Romania Vahabodin Goodarzi Applied Biotechnology Research Center, Baqiyatallah University of Medical Sciences, Tehran, Iran Aravinthan Gopanna Advanced Materials Laboratory, Yanbu Research Center, Royal Commission for Yanbu-Colleges and Institutes, Yanbu Industrial City, Kingdom of Saudi Arabia; School of Chemical Engineering, Vignan’s Foundation for Science, Technology and Research University (VFSTRU; Vignan’s University), Guntur, India Alexandru Mihai Grumezescu Department of Science and Engineering of Oxide Materials and Nanomaterials, Faculty of Applied Chemistry and Materials Science, University Politehnica of Bucharest, Bucharest, Romania Valentina Grumezescu Department of Science and Engineering of Oxide Materials and Nanomaterials, Faculty of Applied Chemistry and Materials Science, University Politehnica of Bucharest, Bucharest, Romania; Lasers Department, National Institute for Lasers, Plasma and Radiation Physics, Magurele, Romania Muhammad Hassan College of Dentistry, University of Lahore, Lahore, Pakistan Khosrow Jadidi Vision Health Research Center, Semnan University of Medical sciences, Semnan, Iran

List of Contributors

Anjali Jain Department of Pharmaceutics, National Institute of Pharmaceutical Education & Research (NIPER), Hyderabad, India Gautam Jaiswar Department of Chemistry, Dr. Bhimrao Ambedkar University, Agra, India Gul Majid Khan Department of Pharmacy, Quaid-i-Azam University, Islamabad, Pakistan Wahid Khan Department of Pharmaceutics, National Institute of Pharmaceutical Education & Research (NIPER), Hyderabad, India Maria A. Krayukhina A. N. Nesmeyanov Institute of Organoelement Compounds, Russian Academy of Sciences, Moscow, Russian Federation Agnieszka Kyzioł Faculty of Chemistry, Jagiellonian University, Krako´w, Poland Leonid M. Likhosherstov N. D. Zelinsky Institute of Organic Chemistry, Russian Academy of Sciences, Moscow, Russian Federation Lindsey Lipp Department of Pharmaceutical Sciences, College of Health Professions, North Dakota State University, Fargo, ND, United States Fernando Lusquin˜os Applied Physics Department, University of Vigo, EEI, Lagoas-Marcosende, Vigo, Spain Hari Madhav Drug Design and Synthesis Laboratory, Department of Chemistry, Jamia Millia Islamia (A Central University), New Delhi, India Alok Mittal Department of Chemistry, Maulana Azad National Institute of Technology, Bhopal, India Jean-Franc¸ois Nguyen UFR d’Odontologie, Universite´ Paris Diderot, Paris, France; Service d’Odontologie Groupe Hospitalier Pitie´ Salpeˆtrie`re, Paris, France; PSL Research University, Chimie ParisTech CNRS, Institut de Recherche de Chimie Paris, Paris, France Olga S. Novikova N. D. Zelinsky Institute of Organic Chemistry, Russian Academy of Sciences, Moscow, Russian Federation

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Shiva Pirhadi Department of Biomedical Engineering, Science and Research Branch, Islamic Azad University, Tehran, Iran Vladimir E. Piskarev A. N. Nesmeyanov Institute of Organoelement Compounds, Russian Academy of Sciences, Moscow, Russian Federation Benjamin Pomes UFR d’Odontologie, Universite´ Paris Diderot, Paris, France; Service d’Odontologie Groupe Hospitalier Pitie´ Salpeˆtrie`re, Paris, France; Arts et Metiers ParisTech, Laboratoire de Proce´de´s et Inge´nierie en Me´canique et Mate´riaux (PIMM), CNRS, CNAM, UMR 8006, Paris, France Juan Pou Applied Physics Department, University of Vigo, EEI, Lagoas-Marcosende, Vigo, Spain Fe´lix Quintero Applied Physics Department, University of Vigo, EEI, Lagoas-Marcosende, Vigo, Spain Krishna Prasad Rajan School of Chemical Engineering, Vignan’s Foundation for Science, Technology and Research University (VFSTRU; Vignan’s University), Guntur, India; Department of Chemical Engineering Technology, Yanbu Industrial College, Royal Commission Colleges & Institutes, Yanbu Industrial City, Kingdom of Saudi Arabia Jacobo Rafael Reyes-Romero Escuela Ba´sica, Facultad de Ingenierı´a, Universidad Central de Venezuela, Caracas, Venezuela Emmanuel Richaud Arts et Metiers ParisTech, Laboratoire de Proce´de´s et Inge´nierie en Me´canique et Mate´riaux (PIMM), CNRS, CNAM, UMR 8006, Paris, France Antonio Riveiro Applied Physics Department, University of Vigo, EEI, Lagoas-Marcosende, Vigo, Spain Mohammed Abdul Samad Mechanical Engineering Department, King Fahd University of Petroleum and Minerals, Dhahran, Saudi Arabia Nadezhda A. Samoilova A. N. Nesmeyanov Institute of Organoelement Compounds, Russian Academy of Sciences, Moscow, Russian Federation Mohammad Sehri Vision Health Research Center, Semnan University of Medical sciences, Semnan, Iran

List of Contributors

Soodabeh Shafiee Department of Biochemistry, Faculty of Biological Sciences, Tarbiat Modares University, Tehran, Iran Divya Sharma Department of Pharmaceutical Sciences, College of Health Professions, North Dakota State University, Fargo, ND, United States Jagdish Singh Department of Pharmaceutical Sciences, College of Health Professions, North Dakota State University, Fargo, ND, United States Neetika Singh Materials Research Laboratory, Department of Chemistry, Jamia Millia Islamia (A Central University), New Delhi, India Rakesh Kumar Soni Department of Chemistry, Chaudhary Charan Singh University, Meerut, India Ramo´n Soto Applied Physics Department, University of Vigo, EEI, Lagoas-Marcosende, Vigo, Spain Meenu Teotia Department of Chemistry, Chaudhary Charan Singh University, Meerut, India Selvin P. Thomas Advanced Materials Laboratory, Yanbu Research Center, Royal Commission for Yanbu-Colleges and Institutes, Yanbu Industrial City, Kingdom of Saudi Arabia; Department of Chemical Engineering Technology, Yanbu Industrial College, Royal Commission Colleges & Institutes, Yanbu Industrial City, Kingdom of Saudi Arabia Asim Ur-Rehman Department of Pharmacy, Quaid-i-Azam University, Islamabad, Pakistan Muhammad Sohail Zafar Department of Restorative Dentistry, College of Dentistry, Taibah University, Al Madinah, Saudi Arabia; Department of Dental Materials, Islamic International Dental College, Riphah International University, Islamabad, Pakistan

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Series Preface In the past few decades there has been growing interest in the design and implementation of advanced materials for new biomedical applications. The development of these materials has been facilitated by multiple factors, especially the introduction of new engineering tools and technologies, emerging biomedical needs, and socioeconomic considerations. Bioengineering is an interdisciplinary field encompassing contributions from biology, medicine, chemistry, and materials science. In this context, new materials have been developed or reinvented to fulfill the need for modern and improved engineered biodevices. A multivolume series, Materials for Biomedical Engineering highlights the most relevant findings and discusses key topics in this impressive research field. Volume 1. Bioactive Materials: Properties and Applications, offers an introduction to bioactive materials, discussing the main properties, applications, and perspectives of materials with medical applications. This volume reviews recently developed materials, highlighting their impact in tissue engineering and the detection, therapy, and prophylaxis of various diseases. Volume 2. Thermoset and Thermoplastic Polymers, analyzes the main applications of advanced functional polymers in the biomedical field. In recent years there has been a revolution in thermoplastic and thermosetting polymers with medical and biological uses, which are currently being developed for medical devices, drug delivery, tailored textiles, packaging, and tissue engineering. Volume 3. Absorbable Polymers, describes the main types of polymers of different compositions with bioabsorbable and biodegradable properties. The biomedical applications of such materials are reviewed and the most innovative findings are presented in this volume. Volume 4. Biopolymer Fibers, highlights the applications of polymeric fibers of natural biological origin in biomedical engineering. Such materials are of great utility in tissue engineering and biodegradable textiles. Volume 5. Inorganic Micro- and Nanostructures and Volume 6. Organic Micro- and Nanostructures, deal, respectively, with the preparation and properties of inorganic and organic nanostructured materials with biomedical applications. Volume 7. Hydrogels and Polymer-Based Scaffolds, discusses the recent progress made in the field of polymeric materials designed as scaffolds and tools for tissue engineering. The technological challenges and advances in their production, as well as current applications in the production of scaffolds and devices for regenerative medicine are presented. Volume 8. Bioactive Materials for Antimicrobial, Anticancer, and Gene Therapy, offers an updated perspective regarding new bioactive materials with potential in the therapy of severe diseases such as infections, cancer, and genetic disorders.

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Series Preface

Volume 9. Nanobiomaterials in Tissue Engineering, provides valuable examples of recently designed nanomaterials with powerful applications in tissue engineering and artificial organ approaches. Volume 10. Nanomaterials-Based Drug Delivery, discusses the most investigated types of nanoparticles and nanoengineered materials with an impact in drug delivery. Applications for drug-therapy, and examples of such nanoscale systems are included in this volume. This series was motivated by the need to offer a scientifically solid basis for the new findings and approaches relevant to the biomedical engineering field. This scientific resource collects new information on the preparation and analysis tools of diverse materials with biomedical applications, while also offering innovative examples of their medical uses for diagnoses and therapies of diseases. The series will be of particular interest for material scientists, engineers, researchers working in the biomedical field, clinicians, and also innovative and established pharmaceutical companies interested in the latest progress made in the field of biomaterials. Michael R. Hamblin1 and Ioannis L. Liakos2 1

Harvard Medical School, Boston, MA, United States 2 Istituto Italiano di Tecnologia, Genoa, Italy

Preface The field of biomedical engineering encompasses elements from a wide spectrum of scientific disciplines. Thermoplastic polymers are characterized by long linear or branched chains of molecules. The polymer chains associate through intermolecular forces, making these materials soften when heat is applied and they can be shaped and then retain that shape upon cooling. Thermosetting polymers play an important role in polymer science due to their characteristic properties, such as temperature stability, high mechanical strength, good viscoelastic properties, and solvent resistance. These polymers are currently investigated for the design of numerous biomedical applications, such as medical devices, coatings and surfaces, drug delivery, and tissue engineering. The purpose of this book is to compile and discuss the most innovative tools developed for the production, characterization, and performance of thermoset and thermoplastic polymers with various biomedical applications. This volume contains 15 chapters prepared by outstanding authors on the field: Chapter 1, Introduction in thermoplastic and thermosetting polymers, by Oana Gherasim et al., provides a general overview regarding the recent progress reported in thermoplastics and thermosets. Particular attention is directed toward plastic-based biomaterials and biomedical devices, as well as on the leading representatives of thermoresponsive biopolymers. Chapter 2, Laser surface texturing of thermoplastics as a method to improve their biological performance, by Antonio Riveiro et al., discusses the utilization of laser sources to control and tailor the surface characteristics of thermoplastic polymers used as matrices in composites for biomedical applications. Surface features, namely topography (at the macro-, micro-, or nanoscale), surface chemistry, and wettability are linked properties and they determine the biological performance of biomedical materials. This complexity presents both a challenge and an opportunity which can be efficiently solved with the application of a laser surface texturing technique. Chapter 3, Light-mediated thermoset polymers, by Meenu Teotia et al., highlights the recent trends in biopolymer applications, their advantages and disadvantages, 3D bioprinting, and their practical applications in fabrication of artificial human organs by providing details of systematic growth in the preparation, properties, and applications of light-sensitive materials. Chapter 4, Thermoset, bioactive, metal polymer composites for medical applications, by Hari Madhav et al., introduces the role of thermosetting polymers in polymer science. These polymers cannot be reshaped (or melt after formation) and due to these properties, they are used in different applications, such as in air craft, electronic industry, household appliances, and others. They are also used in the biomedical field, that is, in bone formation, preparation of dental materials, polymeric heart valves, hard tissue replacements, prosthetic sockets, and in medical devices such as electrocardiographs, among others.

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Chapter 5, Epoxy composites in biomedical engineering, by Satheesan Bobby and Mohammed Abdul Samad, reviews the works carried out in the field of biomedical engineering using epoxy polymeric composite systems. Epoxies belong to a class of polymers commonly referred to as thermosetting plastics. These types of polymers, once set or shaped, cannot be softened or altered in form by the application of heat. In the field of biomedical engineering, epoxy composites have been widely used to prepare components of devices for medical imaging, bone-plate applications, as a novel material for dental applications like scaffolds for tissue regeneration, and as implants for bridging large osteoperiosteal gaps. Chapter 6, Polyethylene and polypropylene matrix composites for biomedical applications, by Aravinthan Gopanna et al., presents an overview of polyolefin and its composites in the biomedical sector and the specific advantages of polyethylene and polypropylene composites for biomedical applications are emphasized. The interest in polyethylene- and polypropylene-based composites is growing as well as taking center stage in biomedical markets, and will continue to develop due to the necessity of these materials for various biomedical applications. Chapter 7, Polymethacrylates, by Benjamin Pomes et al., reviews several cases of methacrylate-based polymers used for medical applications. The main chemicals and fillers used for elaborating biomaterials are presented, together with the main syntheses reactions. Their properties are recalled and discussed using the well-established structure properties relationships of polymer physiochemistry. Chapter 8, Thermoset polymethacrylate-based materials for dental applications, by Muhammad Hassan et al., highlights the benefits of PMMA and recent advancements in PMMA-based materials for dental applications. Encouraging results have been achieved by the incorporation of glass fibers (GFs) within PMMA resins for denture strengthening. The beneficial effect of the addition of GFs on the mechanical properties of acrylic resins has been documented in several research studies. GFs possess excellent esthetics, in addition to having excellent mechanical properties. As a result, they are currently the prime focus of research aimed at reinforcing acrylic dentures. Chapter 9, Maleic anhydride copolymers as a base for neoglycoconjugate synthesis for lectin binding, by Nadezhda A. Samoilova et al., describes a simple, inexpensive, and convenient method of preparation of lectin specific cross-linked, and silver (nanogold)-labeled colloidal neoglycoconjugates based on maleic anhydride copolymers. PLA and its copolymer-based particulate system are well known for controlled delivery of various therapeutic molecules including drugs, proteins, genes, and vaccines. They have established their presence in the market for the treatment of various diseases. Chapter 10, Particulate systems of PLA and its copolymers, by Anjali Jain et al., describes formulation aspects of poly(lactic acid) and its copolymer-based particulate system, ongoing research in this area, and clinical products available in the market. It also provides updated information regarding poly(lactic acid) and its copolymers used in biomedical applications due to their excellent biocompatibility and tunable physicochemical properties.

Preface

Chapter 11, Polylactide: the polymer revolutionizing the biomedical field, by Muhammad Imran Asad et al., talks about PLA synthesis and modifications to alter its properties for various applications. The applications focused on the research done over the past two decades referring to tissue engineering, tumor treatment, imaging and diagnosis, immunization, and DNA, gene, and drug delivery systems. Chapter 12, Poly(propylene fumarate)-based biocomposites for tissue engineering applications, by Ana M. Dı´ez-Pascual, presents the most relevant aspects and concerns of poly(propylene fumarate) (PPF), a linear biodegradable and biocompatible copolyester based on fumaric acid which has recently attracted much attention for biomedical applications. The nanocomposites based on PPF filled with various contents of polyethylene glycol-modified graphene oxide (PEG-GO) or polyethylene glycol-grafted boron nitride nanotubes (PEG-g-BNNTs) showed sufficient stiffness in a physiological environment to be effective as a template for bone tissue regeneration. Besides these, they also have antibacterial activity and are noncytotoxic. These new nanocomposites are perfect candidates for use in the field of bone-tissue engineering. Chapter 13, Diblock and triblock copolymers of polylactide and polyglycolide, by Divya Sharma et al., explores diblock and triblock copolymer based drug delivery systems that harness the diverse benefits poly(D,L-lactic acid) (PLA) and poly(D,L-lactic-co-glycolic acid) (PLGA) and can offer resorbable polymer matrices. Synthesis and characterization of these copolymers and their applications in the development of microspheres, nanoparticles, and stimuli-sensitive delivery systems using thermosensitive diblock and triblock copolymers are discussed in detail. Chapter 14, Characteristics of polymeric materials used in medicine, by E. Davidson and J. Reyes-Romero, gives a complete view of both natural and synthetic polymeric materials used mainly in the medical field. In addition, the authors explain the main failure mechanisms of ultrahigh molecular weight polyethylene (UHMWPE) and the background for nanobiomaterials which are currently used successfully in drug delivery. Chapter 15, Application of polymethylmethacrylate, acrylic, and silicone in ophthalmology, by Hossein Aghamollaei et al., reviews the history, properties, development, and recent advances in the application of these biomaterials in intracorneal rings, intraocular lenses, artificial corneas, and glaucoma drainage devices. Several biomaterials have been used for the design and optimization of eye implants; polymethylmethacrylate, acrylic, and silicone are three examples of these biocompatible materials. Valentina Grumezescu1 and Alexandru Mihai Grumezescu2 1

Laser Department, National Institute for Laser Plasma & Radiation Physics, Romania 2 Faculty of Applied Chemistry and Materials Science, University Politehnica of Bucharest, Bucharest, Romania

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CHAPTER

Introduction in thermoplastic and thermosetting polymers

1

Alexandra Bıˆrca˘ 1,2, Oana Gherasim1,3, Valentina Grumezescu1,3 and Alexandru Mihai Grumezescu1 1

Department of Science and Engineering of Oxide Materials and Nanomaterials, Faculty of Applied Chemistry and Materials Science, University Politehnica of Bucharest, Bucharest, Romania 2Faculty of Engineering in Foreign Languages, University Politehnica of Bucharest, Bucharest, Romania 3Lasers Department, National Institute for Lasers, Plasma and Radiation Physics, Magurele, Romania

1.1 INTRODUCTION The modern biomedical industry aims to go beyond the general requirements of conventional healthcare practice and focuses on the challenging aspects of patient-oriented therapy regardless of the concerned area (i.e., biodetection and diagnosis, preventive and curative medication, or restorative and regenerative medicine). Additionally, the current market trend entails specific requirements which must be considered within this particular emerging domain, such as advanced safety and indisputable quality at a reasonable price. The impressive knowledge based on, or derived from, novel medical-related technologies and the substantial scientific and financial efforts toward modern biomedicine enable the successful design and development of biocompatible, environmentally-friendly, and performance-enhanced materials intended for various medical purposes. Supplementary requirements regarding the selected materials are also considered depending on the specific biomedical-related application of the resulted medical devices or instruments (Kutz, 2002). The particular case of thermoresponsive polymers used for biomedical products and applications is challenging and controversial, therefore, their features need thorough investigation to fulfill specific requirements. The attractive and tunable characteristics related to thermoplastic and thermosetting polymers have gained impressive attention and enable the synthesis of special and unique compounds for modern biomedical applications. Being radiopaque, antistatic and wear-resistant compounds, plastics positively impacted the progress of different industries (Ling et al., 2008; Friedrich, 2018; Youssef and El-Sayed, 2018), but their unique features and technologies are now considered to be beneficial for biomedicine progress (Stewart, 2011). Materials for Biomedical Engineering: Thermoset and Thermoplastic Polymers. DOI: https://doi.org/10.1016/B978-0-12-816874-5.00001-3 © 2019 Elsevier Inc. All rights reserved.

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The impressive nanotechnology-derived opportunities to manipulate mater at molecular and atomic scale enables the improvement of the existing materials and the engineering of new ones (Schro¨fel et al., 2014; Ramsden, 2014) in order to produce biomaterials with genuine physicochemical features and enhanced functionality for unconventional biomedical applications. In the particular case of polymeric biomaterials, relevant criteria considered when selecting a candidate include structural versatility, physicochemical tunability, mechanical suitability, chemical stability, surface chemical reactivity, functionalization potential, solubility, degradability, circumstantial environment-dependent functionality, and specialized institutions’ approval (Ye et al., 2014; Khadka et al., 2014; Karolewicz, 2015). Even though a limited number of natural or synthetic polymers are currently confirmed and approved as medically safe or medical-graded biomaterials for commercial usage (Maitz, 2015; Marasini et al., 2017; Sourey and Bishop, 2018), their intrinsic versatile and tunable biofunctional potential has gained impressive attention regarding the development of novel materials and devices for modern biomedical applications (Sabu et al., 2011). Thermoplastic and thermosetting polymers present tremendous potential regarding the progress of modern industries. In general, plastics (etiologically including all macromolecular compounds which can be shaped by means of casting, pressing, or extruding procedures) can be classified depending on their chemical structure (Pascault et al., 2002). In this chapter we aim to provide a general overview regarding the current progress reported in plastics research (including both thermoplastic and thermosetting polymers) with a particular emphasis on biomedical-related materials and devices.

1.2 BIOMEDICAL POLYMERS Although considered to be primitive at their beginnings (more than 7 millennia ago), all great civilizations historically acknowledged the use of biomaterials in limb restoration or replacement (Thurston, 2007; Alexander, 2009). Ancient civilizations used natural or synthetic materials from nonhuman origins for dental restoration, animal-to-human bone transplantation, wound treatment, and surgical infection control (Ratner, 2004; Hildebrand, 2013; Spattel and White, 2011). With the passage of time, commonly accessible materials were adjusted for biomedical-related applications. Since the 1970s, thanks to the implication and efforts of Robert Langer, complex and genuine biomaterials started to be engineered and developed solely for biomedical applications, such as special metals or ceramics for orthopedic and orthodontic applications, and plastics or polymers for flexible medical devices (Ozdil and Aydin, 2014). Although polymers are naturally softer compounds (in comparison with metals and ceramics), they possess the intrinsic structure-derived ability to be modeled to result in biomaterials with mechanical behavior suitable for both hard- and

1.2 Biomedical Polymers

soft-tissue restorative and regenerative applications (Uskokovi´c, 2015; Stratton et al., 2016). In addition, the current progress derived from polymer functionalization possibilities and polymer manufacturing technologies resulted in increased interest toward the biomedical impact of macromolecular compounds. Currently, polymers represent the predominant class of materials used in the biomedical industry (Kendall and Lynch, 2016) and their attractive and successful application has been reported for a wide variety of biomedical-related domains, including detection and sensing devices (Shahrokhian and Salimian, 2018; Umrao et al., 2018; Zou et al., 2018), drug delivery systems (Szczeblinska et al., 2017; Dave et al., 2017; Yong-Gao et al., 2018), gene therapy vehicles (Olden et al., 2018; Nam et al., 2018), medical device functional coatings (Motealleh et al., 2018; Karimi et al., 2018), and restoring or replacing structures of damaged and nonfunctional tissues (Huang et al., 2017; Kim et al., 2018; Lukasiewicz et al., 2018; Montalbano et al., 2018). The processing of polymers is a remarkably widespread, specialized, and complex domain which considers complementary knowledge from polymer chemistry, rheology, and heat and mass transfer phenomena. In the particular case of biomedical polymers, the current challenge is to adjust the synthesis and processing procedures in order to accordingly manage their distinctive characteristics (especially degradability) and to handle the presence of any residual material which can be accomplished by avoiding processing aids and solvents that possess toxic potential. Over the past three decades, several experimental techniques were tested and selected for the development of plastic-based biomedical materials and devices (Damodaran et al., 2016) such as injection molding (Heidari et al., 2017, 2018), solution processing (Rethinam et al., 2018; Clavijo-Chaparro et al., 2018), spin and dip coating (Kim et al., 2017; Maver et al., 2018; Chartpitak et al., 2018), electrospraying (Li et al., 2017; Moreno et al., 2018) electrospinning (Cho et al., 2015; Zhang et al., 2018), and laser-assisted techniques (Yu et al., 2015; Grumezescu et al., 2017). Biocompatible polymers represent the ultimate and most challenging class of biomaterials and have tremendous potential regarding modern biomedical applications which include, but are not limited to, surgical sutures, tissue engineering scaffolds, medical implants, and drug-eluting devices. In terms of biomedical materials and devices production, polymers possess some major advantages over metallic and ceramic biomaterials due to their intrinsic synthesis and processing simplicity and biodegradability. The adaptability of the synthesis protocols and the tunability of functionalization strategies provide some indisputable advantages of polymers as an attractive and suitable choice for biomedical applications in terms of physicochemical and mechanical properties, biocompatibility, solubility and degradability, encapsulation, and the release profiles of various biologic molecules (Ramakrishna et al., 2001). Depending on the concerned biomedical application, the use of biodegradable polymeric compounds is preferred over nondegradable ones since their enzymatic or nonenzymatic degradation could result in biocompatible and nontoxic products

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which could be further metabolically eliminated without removal procedures being required (Ulery et al., 2011; Kluin et al., 2013). The use of naturally derived biodegradable polymers in medical applications was firstly mentioned in early 3rd century AD, through the successful use of collagen-based catgut sutures (Nutton, 2012; Gomes and Reis, 2004). The implementation of biodegradable synthetic polymers in healthcare practice, however, was only established in the latter half of the 1960s (Nair and Laurencin, 2007). Poly(glycolic acid) or polyglycolide, PGA, represented one of the very openly degradable synthetic polymers for biomedical purposes and was initially used in the 1970s for the fabrication of sutures and later for bone pin applications (Ulery et al., 2011). Given the impactful progress reported in the attractive fields of polymer chemistry and macromolecular processing technologies, several natural, synthetic, or hybrid polymers are currently recommended as safe biomaterials for modern biomedicine. Moreover, when taking into consideration the complex implications of specific biomedical applications and the intrinsic challenging complexity of biologic structures, the quest toward research and engineering of novel and performance-enhanced biomedical polymers multiplied exponentially during the past few years. Undeniably, regenerative medicine and therapeutic medical devices represent the most advanced fields of modern biomedicine, especially due to the versatile and tunable characteristics of macromolecular compounds. Therefore, polymerbased biomaterials dominate the soft tissue engineering and drug delivery industries (Zhang et al., 2018; Ahsan et al., 2018; Yegappan et al., 2018; Santos et al., 2018). Thanks to their close similarity with natural organic tissue components, polymers have been recently introduced in orthopedic clinical practice for hip socket replacement applications. However, due to their intrinsic degradation susceptibility under physiological environments (as a result of biochemical and mechanical factors), the use of polymers for hard tissue replacement or restoration procedures encountered several drawbacks such as tissue irritation and inflammation, corrosion, and wear misfit (Hyman and Privitera, 2005). However, further progress reported in polymer science and nanotechnology led to a tremendous extension of polymer applications so that polymer-based biomaterials and biomedical devices are slowly replacing metals and ceramics in the hard tissue engineering field as well (MarketsandMarkets, 2017; Bhattacharjee et al., 2017; Torgbo and Sukyai, 2018; Soundarya et al., 2018). Following specific medical procedure, bidirectional phenomena occur between the human tissue and the considered biomedical device, so the contact period between these elements is an essential aspect which must be considered when selecting the polymer type (Table 1.1). One of the first and most essential requirements which must be considered when selecting a polymer for biomedical use refers to polymer intrinsic biocompatibility and behavior under biological-like environmental exposure. Regardless of the contact period with body simulated fluids and tissues, the degradation of polymer-based materials or devices should result in biocompatible and

1.3 Thermoplastic and Thermosetting Polymers

Table 1.1 Polymers for Biomedical Applications, Selected Depending on the Contact Period (Hyman and Privitera, 2005) No.

Application of Medical Devices

Polymer Selection

1

Noncontact with human body, e.g., syringes, blood storage bags, glucose drip bags Short-term contact with human body, e.g., catheters, feeding tubes, drainage tubes, surgical instruments Medium-term contact with human body, e.g., cultures, ligatures Long-term contact with human body, e.g., implants, drug delivery devices

PVC, PA, PE, PS, Epoxy resins

2

3 4

PVC, PA, PE, PP, PU, PEEK, Teflon, Silicone rubber, Natural rubber, Polyester, Polyphenylsulfone PA, PP, Polyester PE, UHMWPE, PET, PU, PMMA, Silicone rubber, Polysulphones, Hydrogels, Polyphosphazenes, Thermoplastic elastomers, Polydimethylsiloxane

PA, polyamide (nylon); PE, polyethylene; PEEK, polyether ether ketone; PET, polyethylene terephthalate; PMMA, poly(methyl methacrylate); PP, polypropylene; PS, polystyrene; PU, polyurethane; PVC, polyvinyl chloride; UHMWPE, ultra-high molecular weight polyethylene.

metabolically susceptible by-products. Consequently, the relevant biological assessment of polymer-based materials and devices should consider the bidirectional phenomena occurred after the interaction of the selected biomaterial with the specific biological environment, cell, or tissue.

1.3 THERMOPLASTIC AND THERMOSETTING POLYMERS In general, the properties of macromolecular compounds are strongly related to temperature variation. At adequately low temperatures, amorphous polymers are hard and glass-like, while they adopt a softer and more malleable form when a critical temperature is reached—normally, the glass transition temperature (Tg) (Tiwari et al., 2016). At temperatures above the Tg value, the polymer chains show increased changeableness which allows the bulk material to flow. At temperatures below Tg, the flexibility is restricted and the polymer acts properly like a glassy, elastic solid. The glass transition temperature of a polymer is conditioned by the chemical structure, degree of chemical or physical crosslinking, and molecular weight (Ward and Sweeney, 2012). The Tg value represents the essential indicator regarding the classification of a plastic compound either as a thermoplastic or as a thermosetting polymer, which further determines the potential biomedical use of the considered material.

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The intrinsic material properties of synthetic and semisynthetic plastics are influenced by the values of Tg parameters which fluctuate with molecular weight, as described by Fox and Flory with the relation (Fox and Flory, 1954): Tg 5 Tga 2 K=Mn

where Tga is the highest glass transition temperature that can be reached at a theoretical infinite molecular weight, K is an empirical parameter associated with the free volume of polymer sample, and Mn number is the average molecular weight (Fox and Flory, 1954).

1.3.1 THERMOPLASTICS After reaching the glass transition temperature or melting temperature, thermoplastic polymers soften and adopt a viscous liquid state, followed by their subsequent transformation into glassy or semicrystalline hard solids after the cooling process (Massy, 2017). The thermal melting—solidification behavior of thermoplastics is a reversible and limited process, meaning that a restricted number of heating and cooling cycles can be performed without any structural or functional effects such as color and shape modification, microstructural alteration, and mechanical dysfunction (Radlmaier et al., 2017; Oliveira et al., 2017). If a higher value than the melting temperature is applied to a thermoplastic polymer, then its entire crystalline structure is modified, the linear macromolecular backbone chain becomes randomly scattered, and the specific physicochemical properties can be altered. The relevant characteristics of thermoplastic polymers which recommend them in various research and industry applications include their superior mechanical behavior (improved strength values, high toughness, and hardness), indisputable chemical stability, optical transparency, durability, thermal and electrical behavior, self-lubrication ability, and hydrophobic or waterproofing potential (Ginjupalli et al., 2017; Butt et al., 2018). In order to develop novel mixtures and composite or hybrid materials based on thermoplastics and to further extend their applicability in the biomedical filed, several techniques were successfully used for joining thermoplastic polymers by direct bonding means (Amanat et al., 2010), namely thermal (Yu et al., 2017; Shaegh et al., 2018), friction (Altmeyer et al., 2015; Liang et al., 2018), and electromagnetic bonding (Amanat et al., 2010; Mian et al., 2005).

1.3.2 THERMOSETS The liquid solid state transition that occurs after a thermosetting polymer is heated above its melting temperature and then cooled down is an irreversible solidification process. During the so-called curing process, small molecules are chemically linked and create complex interconnected networks, thus, resulting in

1.4 Biomedical Thermoplastic and Thermosetting Polymers

a permanent hard and rigid product (Massy, 2017). Therefore, further heating of thermosets will result in chemical decomposition and severe structural alteration. In comparison with thermoplastics, the mechanical properties (tensile strength, compressive strength, and hardness) of thermosetting polymers are not temperature-dependent. The current progress in polymer science enables tremendous compositional and structural possibilities regarding the development of novel thermosetting resins for functional and structural practice with properties similar to those of thermoplastics, but with better solvent resistance and improved high-temperature dimensional stability (Meyer et al., 1995; Raquez et al., 2010). Thermoplastics and thermosets have been extensively investigated and explored for high-performance composite applications. One major advantage of thermoplastic polymers is related to their minimal chemical change during and after processing, as well as their ductility and recycling potential. Still, some amorphous thermoplastics possess a crucial deficiency in solvent resistance. By comparison, thermosetting polymers present excellent solvent resistance and a substantial specific database. Nevertheless, thermosetting resins may be too fragile for special applications and their recycling is drastically limited by the crosslinking occurrence (Meyer et al., 1995). Given the sustained demand for modern and performance-enhanced biomaterials and biomedical devices, both thermoplastic and thermosetting polymers represent suitable candidates for such specific applications thanks to their specific versatile processability and attractive physicochemical and biofunctional features.

1.4 BIOMEDICAL THERMOPLASTIC AND THERMOSETTING POLYMERS Novel engineered materials and devices based on thermoplastics and thermosets were successfully assessed as suitable candidates for various biomedical applications. When it comes to the use of plastics in biomaterials and medical devices fabrication, several specific requirements must be considered. The proper selection of a thermoplastic or thermosetting polymer for biomedical devices should undergo an initial evaluation based on the contact type and the contact period with biological fluids and structures (cells and tissues) (Joseph et al., 2017; Rokaya et al., 2018; Xu et al., 2018). Depending on the direct exposure time of plastic-based devices to biological environments, these biocompatible compounds can be selected for short-term use (contact below 1 day), mid-term or prolonged utilization (contact lasting between 1 and 30 days), and long-term or permanent application (contact over 30 days) (Maitz, 2015; Albert, 2002). In the particular case of contact between plastic-based devices and biological media, drug delivery formulations require additional concerns regarding plasticdrug interactions. For instance, the selected thermoplastic or thermosetting

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polymer should not interfere with the drug-related therapeutic effects which is possible by the extraction of additives and monomers or by the chemical reactivity between the polymer and the drug (Guo et al., 1998). In contrast, the plasticbased system should provide suitable solubility and degradation to maximize the drug’s therapeutic efficiency and to minimize or eliminate the side and unwanted effects (De Jong and Borm, 2008; Safari and Zarnegar, 2014). On the other hand, it is essential for the concerned drug not to cause or allow the degradation of the selected plastic (Khadka et al., 2014; Guo et al., 1998). Because thermoplastic polymers are used in medical devices that entail intimate contact with the human body they must undergo a sterilization process which should not affect material properties. The sterilization of plastics which results in the complete removal of all viable microorganisms (Briggs et al., 2009; Govindaraj and Muthuraman, 2015) can be achieved by using distinctive methods such as radiation sterilization (performed by using gamma or electron beam radiation), chemical sterilization (acquired by using ethylene oxide, ozone and hydrogen peroxide), and thermal sterilization (attained by means of autoclaving and dry heat exposure) (Wilson and Nayak, 2013; Rogers, 2012). A large number of thermoplastic polymers can withstand ethylene oxide (ETO) exposure without notable modifications of their properties or color changing. Radiation sterilization can result in chain separation or crosslinking of thermoplastics, thus, significantly affecting their mechanical properties (tensile strength, impact strength, and elongation). It is worth noting that multiple exposures of plastics to radiation sterilization result in cumulative degradation phenomena (Klein, 1987). Several thermoplastics are resistant to radiation sterilization protocols, including acrylonitrile butadiene styrene (ABS), polyarylamide (PARA), polyether ether ketone (PEEK), polyethylenimine (PEI), polyether sulfone (PES), polysulfone (PSU), polyphenyl sulfone (PPSU), and thermoplastic polyurethane (TPU). The autoclaving method, also called steam sterilization, enables the terminal sterilization of medical devices after repeated exposure to high temperature saturated steam under vacuum conditions. There are some thermoplastic polymers (such as PPSU and PEEK) which resist very well during autoclaving and preserve their structural and physicochemical integrity even after thousands of autoclaving cycles. On the other hand, some polyethylenes can be autoclaved at lower temperatures due to their reduced temperature resistance. Because of their intrinsic poor resistance to heat or moisture environment, polymers like styrenes (ABS, PS) and polyesters (polybutylene terephthalate—PBT or polyethylene terephthalate—PET) are not good candidates for autoclaving. Plenty thermoplastic-based medical products are designed and produced for multiple use, so they require several autoclaving treatments. For this purpose, plastics like PEEK, PEI, PSU, PPSU and polycarbonate (PC) which possess higher toughness and a minor tendency to discolor during steam sterilization represent better choices for autoclaving treatments (Leonard, 1994).

1.4 Biomedical Thermoplastic and Thermosetting Polymers

Another aspect related to the medical devices containing thermoplastic and thermosetting polymers refers to their potential degradation after using chemical cleaners. Regular hospital cleaning procedures frequently entail using a large variety of disinfecting chemicals which can damage plastic and impair the medical device’s performance. Some cleaning substances (such as isopropyl alcohol, peroxides, and bleaches) may join to the polymer chains of some thermoplastics and cause stress cracking or even breakage (McKeen, 2014). An important aspect that must be considered when selecting a plastic (above other polymers) for biomedical applications is related to its superior chemical resistance and to the fact that no additives or supplementary improvements will remarkably change the resistance of the polymer to a selected chemical. Usually, semicrystalline polymers (like polypropylene—PP, PE and polyamides—PAs) have higher chemical resistance than amorphous polymers (such as ABS and PC). Still, exceptions are possible as has been reported by subjecting polycarbonate and nylon to hydrogen peroxide which resulted in more stable and mechanically modified compounds, respectively (McKeen, 2014). When selecting a plastic polymer for biomedical applications, it must fulfill specific mechanical requirements such as strength, impact resistance, or stiffness. There are some thermoplastics which possess suitable mechanical properties both under high and low temperature conditions, such as PC, PEEK, PPSU, nylon, and polyoxymethylene (POM). Such specific behavior is essential when considering diverse climate conditions occurring during transportation, for example. During transportation, the effects related to environmental temperature variation should be minimal or even absent since device integrity can be affected (Damodaran et al., 2016). In order to increase patients’ quality of life and to enhance the general healthcare practice, significant efforts have been oriented toward the diminution of traumas related to conventional surgery by improving the safety and efficiency of the surgical procedures. Some medical instruments used during medical procedures underwent almost total design changes and resulted in minimally invasive action on patients. The materials used to handle larger component loads are designed to provide high strength and rigidity, as in the case of polyether amide. The proper selection of a plastic polymer for wear- and environment-dependent conditions includes wear pairs (when a material is in contact with another), wear configurations (like sliding, rotary) and lubricity, respectively humidity and dryness (Damodaran et al., 2016). Another relevant example of plastics used for medical purposes is represented by the fabrication of temperature-resistant and dielectric coatings for medical devices. For thermoplastic-based dielectric coatings, it is essential that they maintain their electrical properties at different temperatures. Several polymers showed remarkable results in this respect, such as polystyrene blends, PC and PC blends, and polyphenylene, etc. (Damodaran et al., 2016). The proper selection of plastic polymers for biomedical applications should also consider the miniaturization tendency. Thanks to their intrinsic facile

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processability and tunable functionality, plastics were successfully used for the fabrication of small-sized devices intended for daily healthcare-related activities by facilitating the qualitative identification and/or quantitative estimation of the biological structures and moieties (Lindstro¨m and Andersson-Svahn, 2011; Rı´os and Zougagh, 2015). In this respect, the dimensional stability of plastic-based biomaterials is another fundamental material-dependent factor which must be considered since firm tolerance property is essential for miniaturized internal constituents. The overall functionality of the device is related to the adequate and accurate interaction of its components. Choosing a material with superior dimensional stability for biomedical devices is not a facile process since some predictions regarding interactions with various chemicals and thermal behavior must be stated. Among thermoplastics, PEI, PC, PES, PSU, and PPSU are acknowledged as dimensionally stable compounds because they possess low and uniform shrinkage and low warpage allowing them to retain tight tolerance in complex components. Semicrystalline plastics own higher and less uniform shrinkage properties, but additive fillers enable shrinkage delay (Fink, 2010). Although the main desideratum in restorative and regenerative medical practice is to restore the structural and functional integrity of the damaged tissue, particular attention must be oriented toward the durability of the material and its related aesthetic aspect. The performance of some equipment and devices designed for clinical use may be strongly impacted by the durability of the plastics used during fabrication. For example, if transportation of certain equipment or devices is required, ABS, PC, and PC-ABS blends are often used for enclosing components (Fink, 2010). In most cases, the plastic material integrated within a medical device undergoes X-ray or fluoroscope imaging which requires accurate discrimination between device-related and healthcare-related aspects. In such situations, the use of certain additives is preferred in order to induce radiopacity to thermoplastic polymers. Particularly, this strategy to generate additive thermoplastics is mandatory for thermoplastics used for surgical implants and catheters that must to be properly monitored during surgery. Usually, thermoplastics attain radiopaque characteristics by using barium sulfate or tungsten-bismuth-based minerals as additives. The use of bismuth subcarbonate and bismuth trioxide is less common since these additives are more expensive and are only used in special cases (Kikuchi and Okano, 2002). Other specific biomedical applications, such as active substance delivery, require appropriate conductive plastics. In some aerosol and dry powder drug delivery applications thermoplastics may encourage the in situ modification of the surface, resulting in improper drug immobilization and subsequent incorrect drug dosages. The solution to remove this effect is to use permanently antistatic compounds which reduce or eliminate static build-up, thus, allowing the accurate dosing of drugs in pharmaceutical applications (Kikuchi and Okano, 2002). Plenty medical devices that have any plastic-on-plastic or plastic-on-metal sliding parts (equipment elements, implants, sliding covers) need to possess

1.4 Biomedical Thermoplastic and Thermosetting Polymers

superior wear resistance; the proper lubrication of these device parts is, thus, required. A thermoplastic material that shows superior wear resistance is ultrahigh molecular weight polyethylene; its mechanical properties can also be improved by using a crosslinking process (the formation of stronger bonding between the molecular chains which constitute the polyethylene). Another strategy often used to improve the lubricity and wear resistance of a plastic material is to use additives like silicone and polytetraflouroethylene (Teflon, PTFE) (Kikuchi and Okano, 2002). Manufacturing feasibility is another important factor to consider when selecting a plastic material for specific biomedical products fabrication, especially given the fact that thermoplastic and thermosetting polymers are obtained and processed by using individual processes (Kikuchi and Okano, 2002).

1.4.1 POLYETHYLENE GLYCOL In terms of intrinsic physicochemical characteristics and biological behavior, the most versatile representatives of polyether compounds with multifunctional biomedical-related applications are polyethylene glycol (PEG) polymers. The molecular weight of PEGs, given by the terminal hydroxyl groups and the backbone constituent ether groups, significantly impact the overall properties of PEGs (Gullapalli and Mazzitelli, 2015; Badi, 2017). Depending on the research field and based on their molecular weights, these polymers are generally known as polyethylene glycols and oxides (PEGs and PEOs) corresponding to low and high molecular weights, respectively. Polyethylene glycols are aliphatic hydrophilic compounds, which are highly soluble in water and most organic solvents and possess excellent intrinsic biocompatibility and biodegradation. By increasing the molecular weight of PEGs, their physical form changes from liquid to solid and several physicochemical modifications occur, such as crystalline nature transformation from amorphous to crystalline (Majumdar et al., 2010; Tai et al., 2017), increasing phase transition temperature (Ye et al., 2013; Feng et al., 2015), hydrophilicity and swelling (Verhoeven et al., 2009; Plisko et al., 2018), decreasing gas permeability (Lin and Freeman, 2004; Trupej et al., 2015), and thermodynamic stability and dissolution rate (Maggi et al., 2002; Windbergs et al., 2009; Wu et al., 2016). In the particular case of transplant surgery—which involves complex procedures for harvesting, storage, and delivery of vital organs—there are several factors which must be considered in order to provide the structural and functional integrity of the organ. The programmatic functional loss of the harvested organ may occur even during conventional hypothermic preservation (Fuller and Lee, 2007; Nicholson and Hosgood, 2017). Transplant surgical procedures must be performed as soon as possible after harvesting, so the time the organ does not have sustained blood flow should be minimized. During the 1980s, a solution based on metabolically inert substances (i.e., ions to initiate osmotic pressure mechanism, various moieties for prevention of cell swelling and ischemia, and inhibition of intracellular edema and hydroxyethyl

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starch for fluid retention in the endovascular space and prevention of extracellular space swelling) was proposed at the University of Wisconsin for organ preservation (Fink, 2011; Karam et al., 2005; Ostro´z˙ ka-Cie´slik et al., 2018). To assess the performance of this storage solution, liver transplants were initially done in dogs. Afterwards, the so-called Wisconsin University (WU) solution became the gold standard for cold preservation of abdominal organs assigned for transplant surgery (Karam et al., 2005; Guibert et al., 2011). In order to enhance the functional performance or to extend the clinical use of a WU preservation solution, notable results were reported by replacing the organic constituent (hydroxyethyl starch—HES) with PEGs or by simply combining WU with PEGcontaining solutions (Guibert et al., 2011; Wicomb et al., 2001). The replacement of HES with polyethylene glycol (8,000 Da molecular weight) in the WU solution was reported on rabbit hearts and resulted in beneficial organ preservation for up to 24 hours under low oxygen perfusion and subsequent excellent cardiac function (Wicomb and Collins, 1989). The superior results obtained in a pig model by replacing HES with PEG 35,000 (Badet et al., 2005) further enabled the clinical transfer of this PEG-containing preservation solution for kidney transplantation (Codas et al., 2009). This solution also proved its beneficial utilization in rat liver transplant surgery (Franco-Gou et al., 2007). The addition of PEG 8,000 to sucrose-based balanced salt solutions enabled proper liver and kidney storage and functions in rodents (Semenchenko et al., 2006) and rabbits (Fuller et al., 2006), respectively. However, there are still some aspects which require thorough clarification regarding the precise action of PEGs in organ preservation solutions. Polyethylene glycols are known to reduce ischemia injury by blocking cell swelling, improve tissue viability, remove free radicals, and interact with cellular membrane lipids (Collins and Wicomb, 1992). PEGs and PEOs are additive constituents of drug formulations designed for oral and parenteral administration. Generally, these polymers are suitable elements for performance-enhanced drug delivery systems. In particular: (1) low molecular weight PEGs are more suitable for drug-related tunability, which is known as the PEGylation process and is achieved by increasing drug stability, solubility and bioactivity, decreasing drug immunogenicity (stealth effect) and controlling pharmacodynamics mechanism (Katre, 1993; Guichard et al., 2017; Khan et al., 2018); and (2) high molecular weight PEGs are more suitable for system-related tunability, which outlines the excipient role of PEOs and is mainly achieved by providing drug system bioavailability and biodistribution, adjusting delivery or release mechanisms, and controlling pharmacokinetics (Sahoo et al., 2015; Khatri et al., 2018; Palcso´ and Zelko´, 2018). For the particular case of pharmaceutical tablets, one of the methods used to improve the activity of the pharmaceutical system is to fabricate thin coatings of PEGs, which can either result in esthetic function (by modifying the appearance, the drug formulation is more easily accepted by the patient) or in structural function (by providing the dimensional integrity of the compressed tablet). For instance, the drug release profile can be adjusted and adapted to the targeted

1.4 Biomedical Thermoplastic and Thermosetting Polymers

tissue or organ by using PEO as the pharmaceutical system coating. The PEO layer can absorb water molecules and form a hydrogel, which further undergoes slower or faster dissolving processes (depending on the molecular weight) and enables drug release. In addition, the drug release kinetics has a reverse dependence with the PEO concentration as was evidenced for distinctive active biosubstances (Seth and Stamm, 2000). When selecting the proper PEG for pharmaceutical formulations, it is mandatory to thoroughly evaluate the effects of the polymer on the drug itself (Gullapalli and Mazzitelli, 2015; Kumar et al., 2018). It is well-known that the current treatment against diabetes consists of insulin injection—this practice has been the reference diabetes treatment option since the 1920s. A medical research study demonstrated that an active pancreas extract had curative results in diabetic dogs. At the same time, the same active extract proved to represent a suitable treatment option for diabetic patients since its administration was related to considerable lifesaving clinical improvement. Due to the negative effects that the conventional insulin administration method has on patients, substantial knowledge and efforts have been directed toward engineering novel insulin-based formulations in an attempt to improve the methods for insulin administration and bioassimilation (Banting et al., 1922). Animal extracts provide all the insulin used for diabetes management. However, the occurrence of recombinant technology enabled the commercial scale production of human insulin. Currently, a mixture consisting of insulin, PEG, and oleic acid can be orally administered in diabetic patients. Besides avoiding the traditional unpleasant administration route, this pharmaceutical dosage can also enhance the activity of insulin via PEGylation means. In such an insulin-based PEG-containing lipophilic conjugate, the bonds between polymer and fatty acid are hydrolysable. In contrast, these bonds are not affected during blood circulation, allowing the proper bloodstream pathway for the highly active insulin-based system (Ekwuribe et al., 2006).

1.4.2 POLYVINYL ALCOHOL Polyvinyl alcohol (PVA) is a synthetic linear and crystalline polymer, highly soluble in water and polar solvents, with attractive and tunable mechanical and physicochemical properties, and excellent biodegradation. By increasing the molecular weight and, consequently, the hydrolysis degree of PVA, its physicochemical characteristics can be modified either proportionally (tensile strength, adhesive strength, water and solvent resistance, and film-forming ability) or inversely proportionally (solubility, flexibility, and water sensitivity) (Tang and Alavi, 2011; Muppalaneni and Omidian, 2013; Rafique et al., 2016). Due to its intrinsic structural and functional versatility, PVA was successfully introduced in safe pharmaceutical formulations, but its biofunctional potential enables its extensive use in various research and clinical domains of modern biomedicine (Kadajji and Betageri, 2011; Wan et al., 2014).

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Hydrogels that contain PVA as the basic substance fulfill the biomechanical, biochemical, and biofunctional requirements encountered in modern biomedical applications (Pasut et al., 2004). Hydrogels based on PVA have been successfully used as corneal prostheses (Veronese, 2001; Bertens et al., 2018), optical lenses (Delgado et al., 1992; Al-Kinani et al., 2018), catheter coatings and catheterization (Ahire et al., 2017; Lin et al., 2018; Tian et al., 2018), artificial organ engineering and regenerative medicine (Khandare and Minko, 2006; Wang et al., 2018; Timofejeva et al., 2017), thin films for wound dressings (Simo˜es et al., 2018; Miguel et al., 2018; Kamoun et al., 2017), and subcutaneous drug delivery devices (Kamble et al., 2017; Ye et al., 2018). More than that, the treatment of acne and pimples has been reported by using PVA-based hydrogels (Yamaoka et al., 1994). Given their easy tunable composition and microstructure, but also their biomechanical versatility (mainly related to its viscoelastic nature, high hydration, and swelling potential), PVA hydrogels were successfully assessed for artificial cartilage applications (Banting et al., 1922; Yang et al., 2017). Another potential application of PVA composites is represented by the development of nondegradable meniscal replacements (Pasut et al., 2004; Hayes et al., 2016). The creep endurance of PVA-based hydrogels can be increased by high temperature annealing. However, such processing may result in pore collapse and subsequent water content and surface lubricity diminish. An attractive approach to avoid such unwanted material-related effects is to improve PVA hydrogels by adding acrylamide (AAm) which enables lubrication enhancement and high-creep resistance maintenance (Banting et al., 1922). A hydrogel membrane consisting of a mixture of sodium alginate and PVA showed suitable controlled release of prazosin hydrochloride (antihypertensive drug) during transdermal delivery evaluation (Suzuki et al., 1984). At the same time, pure PVA hydrogels proved enhanced efficiency for the delivery of atenolol which acts like an antagonist for blood hypertension (Katre et al., 1987). Contact lenses based on PVA can be obtained by a very simple method using repeated freezing defrosting stages of the PVA solution under relatively high temperature conditions in order to obtain transparent gels. Moreover, no addition of chemical crosslinking agents is required, so these materials could represent the higher-level strategy for contact lenses fabrication (when compared to the conventional use of 2-hydroxyethyl methacrylate) (Roffler et al., 2008). Thermogelling polymers, pertaining to the so-called smart biomaterials class, represent the absolute desideratum in specific, selective, and controlled topical or systemic therapies since they adopt the liquid state under room temperature conditions followed by the solid-state transformation under physiological environments. These types of hydrogels imply the regional inoculation of the liquid solution at a temperature below the mammalian body. Subsequently, a thermal phase transition occurs once the hydrogel reaches the host’s body temperature, resulting in the in situ gelation of the hydrogel and the formation of a solid hydrogel-based structure. The thermogelling or thermoresponsive biomaterials showed excellent results in various healthcare-related applications such as wound care, replacement

1.4 Biomedical Thermoplastic and Thermosetting Polymers

of intervertebral discs’ inner core, cartilage replacement, disk replacement, joint replacement, gastrointestinal devices, surgical barriers, and cosmetic and reconstructive surgery (Zhu et al., 2001; Bajpai et al., 2008).

1.4.3 CHITOSAN Chitosan (CS) is a natural cationic polysaccharide obtained by the deacetylation of chitin derived from microorganisms or animal sources. The physicochemical characteristics of CS-based biomaterials (such as solubility, hydration, swelling, and degradation mechanisms) can be properly adjusted by modifying the molecular weight of chitosan, which is also transposed into the modification of degree of removed and substituted acetyl functions (Campana et al., 2017; Islam et al., 2017). The use of chitosan and chitosan derivatives in novel biomaterials for modern healthcare therapeutic strategies relies on their versatile biological properties (biocompatibility, biodegradability, and nonimmunogenicity, as well as their intrinsic hemostatic, antioxidant, antimicrobial, analgesic, and mucoadhesive action) and is an emerging research activity worldwide (Bernkop-Schnu¨rch and Du¨nnhaupt, 2012; Wang et al., 2017). Chitosan is a nontoxic biopolymer with weak intrinsic antimicrobial activity. However, even if the use of CS to maintain pharmaceutical compositions has been limited by its insolubility at pH levels above 6, it is used as an active element in contact lens storage solutions (Samuelsson, 2002; Makkliang et al., 2017). Until recently, some issues were related to the intranasal administration of proteins having a molecular weight of at least 10 kDa. Currently, this limitation can be exceeded by using powdery formulations including the concerned protein and a chitosan derivate which can provide the productive absorption of the protein (Jamieson et al., 1988). Also, a new method for intranasal administration was proposed for vaccine formulations containing chitosan due to the adjuvant effects of CS in immune response improvement (Wicomb et al., 2001). Clinically relevant results with respect to CS-based biomaterials and medical devices were also reported in wound healing (Patrulea et al., 2016; Sandri et al., 2017), tissue engineering (Azizian et al., 2018; Jahan et al., 2019), antimicrobial therapy (Abdelkader et al., 2017; Ameeduzzafar et al., 2018), and antitumor treatment (Tang et al., 2016; Lallana et al., 2017).

1.4.4 SHAPE MEMORY POLYMERS Shape memory polymers (SMPs) is the latest and most attractive representative of intelligent polymer-based biomaterials as they possess the ability to adapt their shape and to exert their expected functionality following an external activation method (Hager et al., 2015; Pilate et al., 2016). Due to their intrinsic physicochemical and biofunctional versatility (including biocompatibility and functionalization potential) and tremendous tunability possibilities, SMPs proved to be ideal candidates for biomedical applications such as actuators, microelectromechanical

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systems, self-healing and health monitoring purposes, and biomedical devices fabrication (Liu et al., 2007). Due to their scientific and technological significance, SMPs currently represent the top choice in various biomedical-related research fields. When compared to shape memory alloys or ceramics, SMPs present deformation to a much higher degree and have a broad scope of fluctuating mechanical properties. Also, by taking into consideration the reduced cost implications and easy processability, as well as the biological characteristics (biocompatibility, biodegradability, and nontoxicity), SMPs provide indisputable advantages over other biomaterials (Anis et al., 2013; Behl and Lendlein, 2007). Biodegradable SMPs came into the spotlight of clinically relevant research after Lendlein and Langer developed a biodegradable polymer based on polycaprolactone (PCL) and proved its potential beneficial impact in medical applications (Lendlein and Kelch, 2002). The currently acknowledged modern biodegradable SMPs are based on poly(L-lactide) (PLLA), polyglycolide (PGA), or polycaprolactone (PCL) and their corresponding copolymers or composites (Jain, 2000). The advantages of biodegradable SMPs based on PCL in the soft phase relies on the recovery temperature of the polymer. PCL has the intrinsic ability to be modified at similar human physiological temperatures and transformed within its corresponding diol, the process being easily adjusted by the proper selection of the molecular weight of PCL, hard phase component selection and soft phase/hard phase configuration (Abayasinghe et al., 2004). SMPs possess enormous potential applications in biology and medicine and are particularly suitable for biomedical devices fabrication (Langer and Tirrell, 2004; Zhang et al., 2018). For example, conventional strategies for medical device implantation often requires complicated implantation surgery, while the more advanced laparoscope surgery enables small and compact device implantation with minimal invasive effects. SMPs can revolutionize the biomedical implants domain since their shape memory effect enables initial miniaturized devices to undergo expansion and shape transformation following the laparoscopic implantation and the physiological-triggered structural modification (Lendlein and Langer, 2002; Tamai et al., 2000). Lendlein and Langer described a thermoplastic elastomer-based SMP with potential application for smart sutures. One of the biggest challenges in the field of endoscopic surgery derives from the necessity of precision procedures on small areas, such as tying sutures or closing an open lumen or incision. The success of such dimensionally reduced endoscopic procedures require wound edges to be properly pressed together or subsequent hernia may develop due to the formation of mechanically impaired scar tissue caused by improper wound closure. In order to overcome such drawbacks related to laparoscopic surgery, smart surgical sutures represent the best alternative. In this case, the temporary shape of the SMPs-based suture can be achieved by controlled elongation while subsequent suture shrinkage and knot tightening occur after physiological stimuli activation (Tamai et al., 2000).

References

The SMPs have been indicated as suitable materials for stent application considering their high strain recovery (which surpasses the unfavorable strain hardening results encountered in metal stent expansion) and their potential to generate smaller and delivery-controllable drug vehicles (Vogt et al., 2004; Palmaz, 2004). The proper utilization of SMPs in the stents industry relies on several materialbased advantages, such as biodegradation capability, better conformity matching, personalized device fabrication, and molecular surface engineering (Cooke et al., 2003; Ewert et al., 2004). In pediatric surgery, the concerned materials must fulfill additional requirements (in comparison with adult surgery). In this case, SMPs may prove beneficial for pediatric stenting since they enable the implanted device to expand and adapt its shape throughout the growth of the patients (Ewert et al., 2004; Epstein, 2004). The material-related setting and programming of the SMPbased stent can be established before implantation surgery so neither auxiliary devices nor additional surgical procedures are required (Yakacki et al., 2007).

1.5 CONCLUSION Thermoplastic and thermosetting polymers represent attractive and challenging candidates in various research and industry domains. In the biomedical field, plastics have been acknowledged as revolutionizing biomaterials since their tunable physicochemical properties and attractive biological functionality outline tremendous applicative potential. By combining particular application-related requirements (toughness, hardness, durability, optical transparency, wear resistance, solvent resistance, self-lubrication ability, and hydrophobic or waterproofing potential) with general biomaterial-related characteristics (such as biocompatibility, nontoxicity, nonimmunogenicity, solubility, biodegradability), both thermoplastics and thermosets proved to represent suitable candidates for modern and performance-enhanced biomedicine.

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microvalves using laser ablation and thermal fusion bonding. Sens. Actuators B 255 (1), 100 109. Shahrokhian, S., Salimian, R., 2018. Ultrasensitive detection of cancer biomarkers using conducting polymer/electrochemically reduced graphene oxide-based biosensor: application toward BRCA1 sensing. Sens. Actuators B 266, 160 169. Simo˜es, D., Miguel, S.P., Ribeiro, M.P., Coutinho, P., Mendonc¸a, A.G., Correia, I.J., 2018. Recent advances on antimicrobial wound dressing: a review. Eur. J. Pharm. Biopharm. 127, 130 141. Soundarya, S.P., Menon, A.H., Chandran, S.V., Selvamurugan, N., 2018. Bone tissue engineering: scaffold preparation using chitosan and other biomaterials with different design and fabrication techniques, Int. J. Biol. Macromol., 119. pp. 1228 1239. Sourey, W.N., Bishop, C.J., 2018. Clinically advancing and promising polymer-based therapeutics. Acta Biomater. 67, 1 20. Spattel, S., White, M.D., 2011. The Portrayal of J. Marion Sims’ Controversial Surgical Legacy. J. Urol. 185, 2424 2427. Stewart, R., 2011. Thermoplastic composites—recyclable and fast to process. Reinf. Plast. 55, 22 28. Stratton, S., Shelke, N.B., Hoshino, K., Rudraiah, S., Kumbar, S.G., 2016. Bioactive polymeric scaffolds for tissue engineering. Bioact. Mater. 1 (2), 93 108. Suzuki, T., Kanbara, N., Tomono, T., Hayashi, N., Shinohara, I., 1984. Physicochemical and biological properties of poly (ethylene glycol)-coupled immunoglobulin G. Biochim. Biophys. Acta 788, 248 255. Szczeblinska, J., Fijalkowski, K., Kohn, J., Fray, M.E., 2017. Antibiotic loaded microspheres as antimicrobial delivery systems for medical applications. Mater. Sci. Eng. C 77, 69 75. Tai, N.L., Adhikari, R., Shanks, R., Adhikari, B., 2017. Starch-polyurethane films synthesized using polyethylene glycol-isocyanate (PEG-iso): effects of molecular weight, crystallinity, and composition of PEG-iso on physiochemical characteristics and hydrophobicity of the films. Food Pack. Shelf Life 14 (B), 116 127. Tamai, H., Igaki, K., Kyo, E., Kosuga, K., Kawashima, A., Matsui, S., et al., 2000. Initial and 6-month results of biodegradable poly-l-lactic acid coronary stents in humans. Circulation 102, 399 404. Tang, X., Alavi, S., 2011. Recent advances in starch, polyvinyl alcohol based polymer blends, nanocomposites and their biodegradability. Carbohydr. Polym. 85 (1), 7 16. Tang, X., Zeng, B., Gao, J.K., Liu, H.Q., 2016. Molecular mechanism of enhanced anticancer effect of nanoparticle formulated LY2835219 via p16-CDK4/6-pRb pathway in colorectal carcinoma cell line. J. Nanomater. 2095878. Thurston, A.J., 2007. Pare´ and prosthetics: the early history of artificial limbs. ANZ J. Surgery 77 (12), 1114 1119. Tian, L., Lu, L., Feng, J., Melancon, M.P., 2018. Radiopaque nano and polymeric materials for atherosclerosis imaging, embolization and other catheterization procedures. Acta Pharm. Sin. B 8 (3), 360 370. Timofejeva, A., D’Este, M., Loca, D., 2017. Calcium phosphate/polyvinyl alcohol composite hydrogels: a review on the freeze-thawing synthesis approach and applications in regenerative medicine. Eur. Polym. J. 95, 547 565. Tiwari, A., Bajpai, A., Bajpai, J., Saini, R.K., Agrawal, P., 2016. Smart Biomaterial Devices: Polymers in Biomedical Sciences. CRC Press.

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CHAPTER

Laser surface texturing of thermoplastics to improve biological performance

2

Antonio Riveiro1, Adolfo Chantada1, Ramo´n Soto1, Jesu´s del Val1, Felipe Arias-Gonza´lez1, Rafael Comesan˜a2, Mohamed Boutinguiza1, Fe´lix Quintero1, Fernando Lusquin˜os1 and Juan Pou1 1

Applied Physics Department, University of Vigo, EEI, Lagoas-Marcosende, Vigo, Spain 2 Materials Engineering, Applied Mechanics and Construction Department, University of Vigo, EEI, Lagoas-Marcosende, Vigo, Spain

2.1 INTRODUCTION Thermoplastic polymers are characterized by long linear or branched chains of molecules. The polymer chains associate through intermolecular forces, make these materials soften when heat is applied, and can be shaped and retain that shape upon cooling. This process is reversible, which makes these materials ideal candidates to be processed by injection molding. Therefore, this is a key property in designing and developing biomedical implants (Puskas and Chen, 2004). Typical thermoplastic polymers used in biomedical applications are ultra-highmolecular-weight polyethylene (UHMWPE), polylactic acid, polyether ether ketone (PEEK), polypropylene (PP), nylon, and various polyesters, among others. Thermoplastic polymers exhibit the excellent mechanical properties required in implants for applications such as knee or hip implants, orthopedic devices, dental implants, ligature clips, tissue staples, and skin covering devices or stents, among others. In addition to the potentially suitable mechanical properties and reduced density of thermoplastics compared to other biomaterials (namely, metals and ceramics), they do not interfere with or deteriorate the biological tissue they are in contact with. This property is essential for biomedical applications. All these features make thermoplastic polymers potentially suitable biomaterials. Hence, over the past few years substantial research has been focused on the use of these materials for biomedical applications (Puskas and Chen, 2004; Shoichet et al., 1994; Lin et al., 2005; Van et al., 2012; Ferna´ndez-D’Arlas et al., 2015). Although these materials are biocompatible and have mechanical properties similar to those of human tissues (e.g., bone), they are chemically and biologically inert, showing a minimum interrelation with the tissues or cells in contact. The advantage of this is that they do not release any toxic constituents to the human body environment. However, on the downside, the body usually responds Materials for Biomedical Engineering: Thermoset and Thermoplastic Polymers. DOI: https://doi.org/10.1016/B978-0-12-816874-5.00002-5 © 2019 Elsevier Inc. All rights reserved.

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to them by inducing the formation of a fibrous connective tissue (fibrosis) on the material surface which can result in the implant becoming totally encapsulated by fibrous tissue. The problem occurs when the fibrous layer prevents the normal operation of the implant in terms of mechanical performance or drug delivery. Numerous studies have concluded that the strength of an osseointegrated implant is far greater than that of a fibrous-encapsulated implant (Parithimarkalaignan and Padmanabhan, 2013). These issues are further aggravated if infections occur due to the colonization of bacteria, fungi, and viruses around the implant. Medications may not work due to the presence of an impervious fibrous layer and the removal of the implant may be the only possible solution for recovery (Dee et al., 2003). One option to address these complications is the coating or mixture of these polymers with bioactive materials, for example, hydroxyapatite or bioactive glass (see, e.g., Tanner, 2010; Durham et al., 2016); however, the adhesion of the coating can be poor or the mechanical properties of the composites can be compromised. One alternative to avoid these drawbacks consists in the enhancement of the biological properties of the polymer surface by applying a proper surface treatment. Then, the mechanical properties remain almost unaltered and the biological properties of the composite can be promoted. This surface treatment can consist in the modification of the surface chemistry and/or topography of the thermoplastic. By using this strategy, the fibrosis of implants can be reduced in order to promote tissue integration. Therefore, control or tailoring of thermoplastic implants (made of a composite material or only thermoplastics) can be an effective way to increase their biocompatibility.

2.2 IMPACT OF ROUGHNESS AND WETTABILITY ON BIOCOMPATIBILITY As discussed, the response of tissues to the implant is largely controlled by the nature and texture of the surface of the material. The main objective of the surface texturing or treatment of the surface of implants is to promote cellular activity. Biocompatibility is closely related to the cell response consequence of the interaction of cells with the surface and, in particular, with cell adhesion (Kaiser and Bruinink, 2004). Textured surfaces show a greater area to improve bonebonding through osseointegration mechanisms. Furthermore, textured surfaces help the ingrowth of the tissues (Alla et al., 2011). Different length scales (macro-, micro-, and nano-) in the surface of the implants are relevant. Macrosized topographies are associated with a roughness value ranging from several millimeters to microns. This scale is directly related to the implant geometry. In this case, the fixation of the implant and its long-term mechanical stability can be promoted through proper macro-roughness (Dahiya et al., 2014). Microsized topographies affect cell adhesion and proliferation.

2.2 Impact of Roughness and Wettability on Biocompatibility

Micro-roughness has been well-established in the literature to have a beneficial effect on the osseointegration of implants (Gittens et al., 2014). Nanosized topographies greatly influence the adsorption of proteins, therefore, they affect the proliferation and enhance cell adhesion. It is also known to alter the gene expression profiles of various cell types and could be a way to direct cell differentiation (Bettinger et al., 2009). Another physicochemical property of the surface implant is the surface energy. This parameter is closely related to the wettability (the ability of a fluid to spread along the surface of a solid) of the material. This property is commonly measured through the water contact angle Θ (WCA) (see Fig. 2.1A), which is the angle between the surface of the liquid and the outline of the contact surface. Materials with a high affinity to water (i.e., hydrophilic) show high surface energy and the WCA is low. On the contrary, for hydrophobic materials the surface energy is low and water forms spherical caps on the surface of the material and the WCA is large (see Fig. 2.1A). It is shown that, hydrophilic substrates improve the cell adhesion and spreading to a greater extent than hydrophobic materials. Schakenraad et al. demonstrated that poor cell spreading on hydrophobic substrata and good spreading on hydrophilic substrata is observed (even in the absence or presence of preabsorbed serum proteins) (Schakenraad et al., 1986). However, the

FIGURE 2.1 (A) Water contact angle for hydrophilic (i.e., with high surface energy) and hydrophobic materials (i.e., with low surface energy). (B) Influence of the roughness on the wettability: The Wenzels model implies that hydrophilic material becomes more hydrophilic and vice versa, while the Cassie-Baxter model implies that the material tends to be more hydrophobic independently of the initial nature (the red dotted line refers to the contact angle for a smooth surface).

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cell adhesion can be drastically reduced for highly hydrophilic materials. Hence, there exists an optimum range for the surface energy (similarly for the WCA) (Baier, 2006). Finally, we should mention the influence of the roughness on the wettability of any material. The Wenzel model was one of the first attempts to quantify and determine the influence of roughness on wettability (Wenzel, 1936). Wenzel determined that roughness promotes the intrinsic wettability of materials. Then, the hydrophilicity is increased with the roughness as this model is based on the assumption of the penetration of liquids into the roughness grooves (Fig. 2.1A). However, if liquid does not penetrate into the grooves, this model cannot be applied and a new approach to deal with the heterogeneity of the surface is required. Cassie and Baxter took this problem into account and developed a new model to address it (Fig. 2.1B) (Cassie and Baxter, 1944). In this case, air pockets trapped in the grooves increase the hydrophobicity of the material regardless of whether or not the material is hydrophobic or hydrophilic. In general, very rough surfaces tend to follow the Cassie-Baxter model, but less rough surfaces tend to the follow Wenzel model (Gennes et al., 2013). Consequently, this analysis shows us a way to manipulate the wettability of materials and, then, their biological performance through a simple modification of their surface roughness. Another possible approach to change the wettability of a material and adjust its biological performance is by varying the surface chemistry, for example, via a chemical modification of the surface. The wettability of a surface is determined by the outermost chemical groups of the solid. In this case, materials which are nonpolar, for example, polymers, show a low surface energy. Therefore, the formation of polar or charged functional groups by chemical methods is another possible approach to address the low wettability of many polymers (Siow et al., 2006).

2.3 SURFACE ENGINEERING PROCESSES The surface modification of materials for biomedical applications is addressed, in general, using two different approaches to enhance their biological characteristics: (1) deformation, removal, or controlled addition of other materials to increase the roughness; or (2) modification of the surface chemistry (Ratner, 1995).

2.3.1 SURFACE ROUGHENING There are several techniques aimed at roughening thermoplastic surfaces (Etsion, 2005) including, among others, sand blasting (Dumitrescu et al., 2011), plasmaand chemical-based etching (Huntington et al., 2013), lithography (Ressier et al., 2012), direct surface machining (micromilling) (Guckenberger et al., 2015), polymer grafting (Chung et al., 2003), and laser-based techniques.

2.3 Surface Engineering Processes

Sandblasting is based on the propulsion of a sand-air (or sand-water) jet at high velocity. The high momentum of the jet is transferred to the surface causing its roughening. The main problem of this technique is possible damage to the surfaces (Karakoca and Yilmaz, 2009). In addition, the abrasion produced during the process removes a substantial amount of material, which may affect the dimension and clinical adaptation of processed workpieces (Zhang et al., 2004). Furthermore, it can generate heat, sparking, and static electricity which can damage the workpieces. Plasma- and chemical-based etching takes place when the top layer of a material surface is degraded by its exposition to volatile etch products (from a plasma) or to corrosive chemicals which induces the dissolution of the solid material. In the case of polymers, this degradation occurs by chain scission, in which former bonds are broken and new ones are formed (Govindarajan and Shandas, 2014). This is considered a versatile technique to obtain a tailored wetting behavior, resulting in an adaptable biological response. However, it also presents some disadvantages, including, among others, the complex physical phenomena involved in the process, high system-dependency of the parameters, reduced adaptability, poor repeatability in the patterning, toxicity of some used gases, and high-cost equipment (Chan et al., 1996). Regarding the surface patterning of polymers, lithography is one of the most common techniques. It is based on the creation of nanosized patterns on a surface by scanning a focused beam of ions (ion beam lithography) or electrons (electron beam lithography). Photolithography is another lithography method using light to transfer a geometrical pattern from a photomask to a light-sensitive photoresist on a substrate. All these techniques are appropriate to accurately control the surface topography of the irradiated surface (Govindarajan and Shandas, 2014); however, the main disadvantage of these techniques are related to the high cost of the equipment. Another disadvantage is the small size of the treated areas. This precludes the implementation of these techniques in large-scale production and the potential contamination of the surface due to the incomplete decomposition of the precursors (Li et al., 2013). An alternative method to address the surface microfabrication tasks is micromilling. This is based on the material removal from a surface via accurate cutting tools to create microscale patterns. However, micromilling is far from being considered an ideal technique to create controlled rough surfaces in thermoplastics due to some drawbacks, such as the production of unwanted debris during the process and its low repeatability. Furthermore, this technique requires large equipment and technical expertise from the operator, resulting in high start-up costs (Guckenberger et al., 2015). Polymer grafting is another technique aimed at increasing the roughness of thermoplastic surfaces and then to improve their biocompatibility (Govindarajan and Shandas, 2014). The process is based on grafting polymer-chain blocks on the surface, affecting the topography at the nanometer-scale (Bhattacharya and Misra, 2004). However, the main disadvantages of this technique is the difficultly

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in controlling the parameters and the high reaction temperatures reached throughout the process (Roy et al., 2005).

2.3.2 SURFACE CHEMICAL MODIFICATION The techniques commonly applied to modify only the surface chemistry of a material are: chemical vapor deposition (CVD) (Alf et al., 2010), grafting techniques (Ku et al., 2010), and self-assembled monolayers (SAMs; Arima and Iwata, 2007). CVD is a chemical process used to produce thin films via the reaction of two or more gases initiated by thermal energy or other energetic sources such as plasma or laser. Although this method aims to alter the chemistry of a surface, it is normally used for building micro- and nanosized features into ordered structures (Guo et al., 2011). This technique has been widely used to modify the chemistry of thermoplastic surfaces with the aim to enhance their biological interactions (Alf et al., 2010). The main drawbacks of this technique are the high temperatures required (up to 900 C) and the necessity of specialized precursor materials which increases the overall costs. Moreover, some gases used in this process are toxic and corrosive (Jones and Hitchman, 2008). Chemical grafting of thermoplastics alters the surface reactivity of a material by adding polymer chains, using chemical methods, onto it. However, polymer grafting presents the disadvantage of lack of control throughout the process, given the great dependency between parameters and results (Thakur and Thakur, 2014). Another method to chemically modify thermoplastic surfaces is via the formation of SAMs (Arima and Iwata, 2007). SAMs are based on molecular assemblies created on surfaces by adsorption. These molecules are composed by head groups attached onto the substrate, and tail groups (functional groups), which provide tailored properties to the modified surfaces (Govindarajan and Shandas, 2014). The main disadvantages related to SAMs are the limited potential surfaces that can be treated, the final substrate quality, and the high costs of the technique (Mrksich and Whitesides, 1996).

2.4 BASICS OF LASER SURFACE TEXTURING 2.4.1 INTRODUCTION Laser surface texturing (LST) is one of the most promising methods to overcome most of the drawbacks presented by most of the techniques for biomedical applications (Riveiro et al., 2018). This technique is able to modify simultaneously the surface roughness and chemistry of materials, avoiding the involvement of toxic substances. In the case of the topography, this can be tailored at the macro-, micro-, and nanosize scale, simultaneously, and without using different tools. Furthermore, it presents a high spatial and temporal resolution (Lippert, 2004).

2.4 Basics of Laser Surface Texturing

Contamination of the workpiece can be avoided as this is a noncontact technique (which is a major advantage for biomedical applications). Furthermore, the high processing speed, treatment of large areas, high throughput, and easy automation should be also be noted.

2.4.2 PROCESS FUNDAMENTALS LST is a simple processing technique to simultaneously change the topography and chemistry of surfaces (Etsion, 2005). A laser beam is focused onto the surface of a material and is absorbed by the topmost layer (Fig. 2.2). This optical energy is thermalized and induces the heating of the material. During the process, melting or even vaporization temperatures are reached. Then, selective removal of material is produced. On the other hand, if ultraviolet (UV) lasers are used, the photons have enough energy to break chemical bonds and to modify the chemistry of the material. Hence, thermal or photochemical processes can accomplish the surface modification of thermoplastic parts with laser technology: •

•

•

Thermal processes: the optical energy is absorbed by the surface of the material increasing its temperature. This induces diverse phenomena, including melting or vaporization (Tan et al., 2015). In general, these phenomena modify the surface roughness. Photochemical processes: if UV lasers are used, the photon energy is enough to break molecular bonds. This mechanism is responsible for most of the chemical modifications of surfaces (e.g., oxidation, ion implantation, etc.) (Wong et al., 2001). Photophysical processes: thermal and nonthermal mechanisms are responsible for the overall processing rate. The degree of thermal and nonthermal

FIGURE 2.2 Scheme of the principle of operation of laser surface texturing.

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contributions depends on the relative yield of the respective reaction channels (Ba¨uerle, 2013). Under these processes, both surface roughness and chemistry can be modified. Laser texturing can be carried out by the production of regular or irregular patterns of dimples, bumps, or grooves (Etsion, 2005). The final topography and chemical modifications in the surface depend on the preponderance of thermal or nonthermal processes; for example, if the melting of the surface is preponderant, bumps, depressions or the foaming of materials can occur. The increment of the absorbed optical energy produces the vaporization of the surface. Then, the removal of material is mainly produced; however, some melting or thermal degradation appear as side effects, especially if long laser pulses are used. Nonthermal processing is associated with the use of ultrashort-pulse lasers (i.e., laser pulses with a duration in the order or picoseconds or less). Pits and grooves to avoid undesirable thermal effects (e.g., generation of a heat affected zone) can be produced. In this case, molecular or atomic bonds are broken and heating is avoided. Ultrafast laser processing eliminates the necessity of post-processing as no recast layers are formed (Ba¨uerle, 2013).

2.4.3 COMPONENTS OF A LASER TEXTURING SETUP LST can be performed under two main configurations: (1) using a stationary laser beam; or (2) using a nonstationary laser beam with regard to the workpiece. In the first case, texturing is performed using a mask (see Fig. 2.3A). A stencil of the desired mark is projected onto the workpiece. Then, the picture of the mask on the object is made using a lens. This requires the application of an extremely short impulse of light energy on the workpiece. Then, pulsed lasers, such as a pulsed TEA CO2 laser, excimer laser, or pulsed Nd:YAG laser (in general, with an average power less than 100 W) are normally used. In the second case, texturing is performed via the relative motion of the laser beam with respect to the surface of the workpiece. This relative movement can be performed by moving the laser, the workpiece, or both. The most common approaches for doing this is the utilization of a Cartesian system (Fig. 2.3B) or galvanometer mirrors (Fig. 2.3C). The last alternative is more practical due to higher versatility and throughput.

2.4.4 PROCESSING PARAMETERS 2.4.4.1 Wavelength of the laser radiation The optical properties of thermoplastics without any additives or reinforcements are influenced by the macromolecule chain structure as well as by the type of bonding within the macromolecule chain and between chains. The wavelength of the electromagnetic radiation strongly determines the interaction of this radiation with the materials. Furthermore, the optical response will be also be dependent on the temperature because the material structure of thermoplastics depends on the temperature.

2.4 Basics of Laser Surface Texturing

FIGURE 2.3 Laser surface texturing using (A) a stationary beam and a mask or by imposing a relative movement between the laser beam and the workpiece using (B) a Cartesian system or (C) a scanning system.

In general, thermoplastics are transparent to visible and near-infrared laser radiation (NIR), which includes the radiation produced by the Nd:YAG, fiber, or Nd:YVO4 lasers, among others (see Fig. 2.4). In this regard, amorphous plastics such as PC and PMMA practically transmit the whole laser radiation, while semicrystalline thermoplastics, such as PA6 or PP, exhibit a lower transmittance due to scattering on the crystalline structures; however, the absorption of the laser radiation in this case is mainly produced in the volume of the workpiece and the surface practically remains unaltered. Therefore, laser texturing of thermoplastics with these laser radiations requires the utilization of additives to increase the laser absorption in the surface of the thermoplastics. Carbon black pigments arise as an effective method (see Riveiro et al., 2014, 2016). In the spectral range of mid-infrared laser radiation (MIR), the interaction of the electromagnetic radiation and the thermoplastics commonly produce melting of the surface and even thermal decomposition if the deposited energy is excessive. These

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FIGURE 2.4 Transmittance of two typical thermoplastics, polypropylene (PP) and ultra-high-molecularweight polyethylene (UHMWPE), as a function of the laser wavelength.

wavelengths are able to produce grooves and pits on the surface of thermoplastics in a very efficient manner.

2.4.4.2 Beam mode (continuous-wave versus pulsed lasers) Lasers can operate under two main modes: continuous-wave (CW) or pulsed operation. CW lasers produce a continuous, uninterrupted beam of light, ideally with a very stable output power. On the other hand, the pulsed laser sources produce an output which takes the form of pulses of light—in general, the pulse duration is in the micro- or nanosecond regime, while ultrafast pulses refer to femtosecond and picosecond pulse width. Depending on the operation mode, materials respond very differently according to the laser radiation mode used. CW lasers play an important role in surface texturing based on thermal interaction mechanisms. In these cases, the focalized laser beam mainly melts the workpiece surface. Therefore, the resulting scanned surface shows protrusions formed by molten and rapidly resolidified material. Conversely, in the case of pulsed mode lasers, both thermal and photochemical interaction mechanisms may take place. In this case, if the laser-induced excitation is sufficiently high, surface molecules can be directly broken. This phenomenon happens using high repetition rate lasers; pico (ps) and femtosecond (fs) lasers, or striking surfaces with highly energetic photons using UV laser sources.

2.4.4.3 Pulse length The duration of the laser pulses during processing in the pulsed mode notably influences the response of the material to the laser radiation. This parameter

2.5 Laser Surface Texturing of Thermoplastic Polymers

allows the precise control of the energy deposition on the material. It should be emphasized that for nonmetals, in particular for polymers, the thermalization time can be as long as 1026 seconds (Ba¨uerle, 2013). Therefore, thermal effects (e.g., melting or vaporizations) are clearly dominant for pulses with a length ranging from ms, up to even ns. Nonthermal effects arise for ps and fs laser pulses. The major advantage of ultrashort laser pulses (,1 ps) is its ability to perform clean and precise patterning (Heitz et al., 1994).

2.4.4.4 Pulse energy Laser ablation of thermoplastics by means of highly energetic photons (UV laser sources) has attracted much interest for biomedical applications (Gotoh et al., 2011). By irradiating a surface using laser wavelengths beneath the visible range (λ , 400 nm), besides the surface patterning, an alteration in the surface chemistry is commonly observed. An example of this is found in research conducted by Slepiˇcka et al. in 2013. In this study, poly-L-lactic-acid was subjected to ablation using a KrF laser (λ 5 248 nm, pulse duration 5 20 40 ns) to assess the effect of the fluence and number of pulses on the final roughness and wettability. It was found that, an increase in these parameters lead to an increase in both roughness and wettability. Regarding surface chemistry, they concluded that it was strongly influenced not only by the pulse energy, but also by the number of laser pulses (Slepiˇcka et al., 2013).

2.5 LASER SURFACE TEXTURING OF THERMOPLASTIC POLYMERS 2.5.1 POLY(ETHERETHERKETONE) PEEK is an engineering thermoplastic showing superior mechanical properties, high temperature durability, and chemical resistance. In addition to that, PEEK exhibits bone-like stiffness (Abu Bakar et al., 2003) and sterilization capacity (Godara et al., 2007). These properties make it suitable for biomedical applications. This polymer is used as an intervertebral spacer, spinal cage, or cervical disk prosthesis. However, this thermoplastic also has a poor interfacial biocompatibility due to its large chemical stability which leads to poor bone-implant interactions. LST of PEEK was initially addressed by Laurens et al. (1998). PEEK surfaces were modified below the ablation threshold using an ArF laser (λ 5 193 nm and pulse duration 5 20 ns). It was observed that the influence of the processing atmosphere (i.e., a neutral or an oxidizing atmosphere) on the chemical modifications. Loss of polymer aromaticity and scission of carbonyl groups occurred using neutral atmospheres; however, the high energetic incoming photons increased the C O/C and carboxylic functions for oxidizing atmospheres. It was observed that adhesion can be controlled as the polar component was enhanced after the laser treatment.

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The influence of other laser wavelengths was also studied (Riveiro et al., 2012; Zheng et al., 2015; Wilson et al., 2015). LST was demonstrated using laser sources ranging from UV (355 nm) to middle infrared (MIR; 10.6 μm) laser wavelengths. Riveiro et al. studied the effect of 355, 532, and 1064 nm ns laser wavelengths on the surface roughness and wettability of PEEK. Response substantially varied depending on the laser wavelength (see Fig. 2.5A). Burning of the treated surface occurred for 1064 nm laser wavelength, while the surface was mainly ablated for the 532 nm laser radiation. Small grooves with a typical width of 100 μm were produced. Samples treated with UV laser radiation (355 nm) only showed slight thermal affection. This laser wavelength was considered the most suitable for biomedical purposes. The formation of polar groups, such as carboxyl (O CQO) and peroxide (O O) groups, given the exposition to UV laser wavelength, successfully reduced the WCA of PEEK after LST (see Fig. 2.5B) which, in turn, increases the cell adhesion of treated PEEK surfaces (Riveiro et al., 2012). Similar results were obtained using Q-switched Nd:YAG lasers (λ 5 1064 nm) (Wilson et al., 2015). In this case, an increment in the surface energy (from 44.9 up to 78.5 mJ m22) and in the WCA was observed. This was attributed to the increment in hydroxyl and/or carboxylic groups, along with the reduction in carbonyl groups. Then, NIR lasers were demonstrated to be suitable for LST of PEEK due to the formation of functional polar groups on the treated surfaces. The increased biocompatibility of treated surfaces was showed by Zheng et al. (2015). This work demonstrated that a combination of CO2 laser (λ 5 10600 nm) and plasma treatments increased the cell adhesion and proliferation of MC3T3-E1 preosteoblast cells on PEEK surfaces. A simultaneous increment of carboxylic groups on the treated surface was also observed. Consequently, the superior biocompatibility was attributed to the induced roughness and formation of polar groups (Zheng et al., 2015).

2.5.2 POLYCARBONATE Polycarbonate (PC) is an amorphous thermoplastic with mechanical properties and biostability suitable for biomedical applications (e.g., in cardiac surgery) (Ahad et al., 2014). LST of PC for biomedical applications was mostly performed using ns laser sources with wavelengths ranging from UV (λ 5 248 nm nm) to NIR (λ 5 1064 nm). LST of PC was firstly addressed in PC/PMMA blends by Viville et al. (1996). The main physical and chemical modifications on KrF (λ 5 248 nm and pulse duration 5 30 ns) laser-treated surfaces were evaluated. It demonstrated the possibility to induce micropatterns and chemical modifications on PC/PMMA blended surfaces. Chemical modifications after UV laser treatment was also studied in depth by Ahad et al, using a laser-plasma extreme ultraviolet source (Ahad et al., 2014). The analysis of the surfaces by XPS analysis showed

2.5 Laser Surface Texturing of Thermoplastic Polymers

FIGURE 2.5 (A) SEM micrographs of laser-treated PEEK surfaces, and (B) optical images of water drops on the surface of PEEK treated with 1064, 532 and 355 nm laser radiation (under identical processing conditions). Extracted from: Riveiro, A., Soto, R., Comesan˜a, R., Boutinguiza, M., Del, V., Quintero, F., et al., 2012. Laser surface modification of PEEK. Appl. Surf. Sci. 258 (23), 9437 9442.

a reduction in the oxygen content and an increment of the WCA in treated PC surfaces. In contrast, opposite results were obtained by Laurens using similar processing conditions (λ 5 193 and 248 nm, and pulse duration 5 20 and 30 ns) (Laurens et al., 2000).

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Surface oxidation and new polar groups on the PC surface were observed after treatment with UV laser radiation. The formation of carboxyl (O CQO) and peroxide (O O) groups was assigned to the scission of initial bonds of PC (due to the interaction with energetic UV photons) as well as to the further reactions leading to radicals due to the presence of an oxidizing atmosphere (Laurens et al., 2000). The biocompatibility of laser-treated PC surfaces was addressed by Ramazani and coworkers (2009). In this case, a pulsed Nd:YAG laser source (emitting laser radiation with λ 5 355 and 1064 nm and pulse duration of 10 ns) was utilized. Treatment with 1064 nm laser radiation led to the thermal degradation of the material; however, the material was photodegraded during 355 nm laser radiation. Roughness increased under both treatments, but WCA was only decreased (from 70 degrees for the base material up to 40 degrees for the treated surfaces) for the 1064 nm laser treatment (WCA remained similar for 355 nm laser treatment). Under both treatments, polar groups (especially, carbonyl groups) were formed. Cell viability tests using mouse L929 fibroblast cells confirmed the improvement in cell growth using 1064 nm laser radiation (Fig. 2.6).

2.5.3 POLYPROPYLENE PP shows good thermal and chemical stability as well as suitable mechanical properties (Himma et al., 2016). However, the integration of PP with surroundings tissues when implanted is poor due to its excellent chemical stability (consequence of its low surface energy). Different treatments were devised to address this problem, for example, coating deposition (Erbil et al., 2003), plasma treatment (Cui and Brown, 2002; Lai et al., 2006), graft polymerization (Shim et al., 2001), or injection molding (Myllymaa et al., 2009), among others. However, precise control in the process and subsequent surface modification remained a major challenge. Production of micro- and nanosized patterns on PP surfaces were produced by Belaud et al. using a fs laser source (Ti:Sapphire laser, λ 5 800 nm, pulse duration 130 fs) (Belaud et al., 2014, 2015). Ablation was identified as the main removal mechanism. This process was promoted by coating the PP surface with a 2 3 μm silver layer. LST of PP surfaces using ns lasers with different laser wavelength (Nd:YVO4 laser, λ 5 1064, 532, and 355 nm, pulse duration , 100 ns) was studied by Riveiro et al. (Riveiro et al., 2016). Surfaces were carbon-coated prior to any laser treatment to address the problem of the large transparency (see Fig. 2.4) of PP to laser wavelengths in the range of λ 5 400 1600 nm (Wissemborski and Klein, 2010). Melting was identified as the main physical mechanism in LST of PP. Moreover, adhesion of carbon particles to treated surfaces was observed after processing. Treated surfaces showed a roughness larger than Ra 5 1 μm, the value considered to be the minimum required for the improvement of bone to implant bonding (Stanford, 2008). Treated surfaces showed carbonyl (CQO) and

2.5 Laser Surface Texturing of Thermoplastic Polymers

FIGURE 2.6 SEM micrographs of cell growth (L929 fibroblast cells) on: (A) untreated PC sample molded on a steel substrate; (B) untreated PC sample molded on a glass substrate; (C) PC sample treated with 1064 nm laser radiation (beam diameter 0.8 mm, 5 pulses); (D) PC sample treated with 1064 nm laser radiation (beam diameter 0.8 mm, 1 pulse); and (E) PC sample treated with 1064 nm laser radiation (beam diameter 3 mm, 5 pulses). Extracted from: Ramazani, S.A.A., Mousavi, S.A., Seyedjafari, E., Poursalehi, R., Sareh, S., Silakhori, K., et al., 2009. Polycarbonate surface cell’s adhesion examination after Nd:YAG laser irradiation. Mater. Sci. Eng. C 29 (4), 1491 1497.

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hydroxyl (C O) groups. The formation of these polar groups combined with a suitable topography suggested that treated surfaces should show a suitable biological response; however, no biological characterization was performed. Buchman et al. obtained similar results after treatment with a Nd:YAG laser (λ 5 1064, 535, 350 and 266 nm) (Buchman et al., 2014). A laser wavelength of 1064 nm was identified as the most suitable for the production of a microtopography suitable to promote cell adhesion. Oxidation was observed in treated surfaces. On the other hand, Murahara and Okoshi studied the ArF laser treatment of PP surfaces in the presence of tap water (Murahara and Okoshi, 1995). C H and C H3 bonds are responsible for the good chemical stability of PP. In this case, laser treatment creates OH functional groups on the surface. The authors identified the optimum laser fluence and shot number (laser fluence of 12.5 mJ cm22 and a shot number of 10,000) to improve wettability. The reduction in the WCA after the excimer laser treatment (WCA for base material: 93 degrees, minimum WCA after laser treatment 65 degrees) was attributed to the formation of OH groups and not to the modification of the surface topography.

2.5.4 POLYETHYLENE Polyethylene (PE) is bioinert and nonbiodegradable when in contact with corporal fluids. This thermoplastic is used for implants in multiple applications—for example, jaw reconstruction (Deshpande and Munoli, 2010). LST was explored to improve the biological response of PE (Okoshi and Inoue, 2003, 2004; Blanchemain et al., 2007). Fs-laser treatment (Ti:Sapphire laser, λ 5 790 nm, and 395 nm) was applied by Okoshi and Inoue on PE surfaces (Okoshi and Inoue, 2003, 2004). Ablation was observed under both laser wavelengths; however, the ablation threshold was higher for 790 nm laser treatment (ablation threshold for the 790 nm wavelength was around 50 mJ cm22, while for 395 nm was 17 mJ cm22) (Okoshi and Inoue, 2003). This was attributed to the higher energy of photons for the 395 nm laser radiation. In both cases, the formation of carbonyl (CQO) polar groups was identified in the treated surfaces (Okoshi and Inoue, 2004). Laser treatment with a second harmonics 532 nm Nd:YAG laser (pulse duration 5 10 ns) was studied by Yi and Feng (1999). The ablated surface showed isolated particles at the bottom of the spots and no rims were detected. The biological characterization of laser-treated surfaces was addressed by Blanchemain et al. (2007). PE surfaces showed a higher roughness (Ra 5 0.2 μm) after Nd:YAG (λ 5 1064 nm) laser treatment. Formation of new polar groups (e.g., hydroxyl groups) and oxidation of the surface were detected after treatment with XPS. LST did not significantly improve the WCA; however, the cell adhesion and proliferation (determined with in vitro tests using human embryonic epithelial cells L132) were promoted. Cells showed higher spreading with large

2.5 Laser Surface Texturing of Thermoplastic Polymers

lamellipodia. This indicates active cell migration. This better biological performance was attributed to the chemical and topographical modifications after LST.

2.5.5 POLY(ETHYLENE TEREPHTHALATE) Poly(ethylene terephthalate) (PET) is a thermoplastic polymer based on modified ethylene glycol and purified terephthalic acid. Its use for biomedical applications is focused on membranes, vascular grafts, ligaments, and tendon repairs (Kannan et al., 2005; Maitz, 2015). However, this polymer is also bioinert. Wong et al. evaluated the effect of KrF excimer laser (λ 5 248 nm) radiation on the chemical properties of PET (Wong et al., 2001). XPS analyses revealed that after treatments using high fluence, the O:C ratio decreased, whereas the contrary effect was observed after high fluence treatments. This effect was attributed to CO and CO2 desorption. In addition, it also revealed an increase in the oxygen content with the increasing number of pulses. New polar groups (e.g., carboxyl) were induced, as commonly happens with most of the polymers. The increase in oxygen content resulted in a decrease in the WCA. Deeper analyses, including biological tests, were also performed by Mayer and coworkers (2006). This research used the same laser equipment to perform LST on PET. Regarding chemical composition, similar results as the research conducted by Wong and coworkers were revealed. These results showed to be potentially favorable for cell attachment. In vitro tests were also carried out using human embryonic epithelial cells L132 and it was demonstrated that LST effectively enhanced cell behavior (in particular cell proliferation and adhesion) along with cell morphology modification. LST on PET was also performed using CO2-pulsed laser sources. The effect of this laser wavelength on the surface modification of PET in terms of chemical composition, and wettability was assessed by Dadsetan et al. (1999). In this study, the laser treatment resulted in an increase in the surface wettability of PET because of the formation of oxidized groups, such as hydroperoxides, hydroxyl, and additional carbonyl functions. These results could indicate an enhancement of the biological interactions of irradiated surfaces (Dadsetan et al., 1999). In order to prove this, the same authors published complementary studies incorporating biological tests to the research (Dadsetan et al., 2001). The hemocompatibility of irradiated surfaces was examined in vitro to assess the platelet adhesion compared to the pristine surface. In this way, the thrombogenicity of PET surfaces was evaluated by the data obtained from the number of platelets adhered to the surface. This means that the lower the number of adhered platelets resulted in a lower risk of formation of procoagulant sites. It was found that the number of platelets adhering to PET surface was successfully reduced after LST. This fact demonstrated the ability of MIR laser irradiation to enhance the biomedical applications of PET. Mirzadeh and Dadsetan compared the process of texturing using a KrF excimer laser (λ 5 248 nm) and line tunable pulsed TEA CO2 laser emitting at

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λ 5 9.25 μm or λ 5 10.28 μm (Mirzadeh and Dadsetan, 2003). They found that different textures were formed as a function of the laser wavelength used in the experiments. The microstructures on the KrF irradiated PET surface were in the range of 1 μm, whereas the microstructures on the CO2 irradiated PET were around 10 20 μm. It also was noticed that the typical dimensions of the textures on the PET surfaces grew in size with increasing laser wavelength. In all the cases, the laser-treated surfaces exhibited larger wettability. The authors hypothesized that the formation of various kinds of oxidized groups (such as carboxylic acids, hydroxyl, and aldehyde) was responsible for the larger wettability of the CO2 laser treatment. In the case of the KrF laser treatment, the authors suggest that laser irradiation produces radicals on the surface that may be converted to hydrophilic groups in contact with air. Fibroblast cell adhesion and spreading was evaluated in the laser-irradiated surfaces with both surfaces. This shows that in the case of KrF laser, wettability is the main parameter affecting cell spreading, whereas both surface morphology and wettability are important in cell attachment and spreading on the CO2 laser-treated surfaces.

2.5.6 ULTRA-HIGH-MOLECULAR-WEIGHT POLYETHYLENE UHMWPE is a thermoplastic made up by long chains of polyethylene aligned in the same direction. This polymer shows high wear resistance, good chemical stability, and does not elicit a toxic response when in contact with corporal fluids. UHMWPE is highly used in biomedical applications, for example, in hip, knee, patella, or spine implants (Riveiro et al., 2014; Lorusso et al., 2008). However, this material is also bioinert as a consequence of its good chemical stability. LST of UHMWPE was demonstrated using different laser sources, from fs- up to ns-laser sources (Riveiro et al., 2014; Torrisi et al., 2010; Ferna´ndez-Pradas et al., 2012). Ferna´ndez-Pradas et al. used a femtosecond laser (Yb:KYW fs laser, λ 5 1027 nm, pulse duration 5 450 fs) to produce craters on UHMWPE (Ferna´ndez-Pradas et al., 2012). The utilization of this laser source modified the surface topography, but the surface chemistry remained unaltered. The highest ablation efficiency was obtained for pulse energies above 6 μJ. In contrast, the utilization of longer pulses, but with a shorter laser wavelength, demonstrated the possibility to modify the surface topography and the chemistry in the ablated areas. Torrisi et al. showed the breaking of C H and C C chemical bonds during ablation with an UV ps-laser (iodine PALS; Prague Asterix Laser System laser) (Torrisi et al., 2004). Moreover, the increment in carbon content for the treated areas was also confirmed. Riveiro et al. studied the influence of three ns laser wavelengths (1064, 532, and 355 nm) (Riveiro et al., 2014). Samples were carbon-coated prior to processing to increase the absorption of laser radiation. Part of the carbon particles remained attached to the treated areas due to the melting of the surfaces. These particles were not considered harmful as they do not induce toxic effects on tissues in contact or severe inflammatory reactions. Minor chemical modifications were observed using these laser wavelengths; however,

2.5 Laser Surface Texturing of Thermoplastic Polymers

the topography and the WCA was modified compared with the base material. The 532 and 355 nm laser radiations produced surfaces with an average roughness close to Ra 5 1 μm, and the WCA was substantially reduced. Cell viability tests were not performed, but the increment in the roughness and the formation of polar carbon clusters were noticed to enhance the biocompatibility of the treated areas (Ramakrishna et al., 2001; Grubova et al., 2016).

2.5.7 POLY(METHYL METHACRYLATE) Poly(methyl methacrylate) (PMMA) is an amorphous thermoplastic that, as in the case of PC, presents high transmittance to visible light. This fact makes it an alternative to glass for applications needing transparency. Moreover, in the past few years the use of this material for lab-on-a-chip applications has substantially increased (Van et al., 2012; Wang et al., 2011; Tsougeni et al., 2012; Marco et al., 2012) because of its low cost, transparency, and appropriate mechanical and chemical properties. In order to enhance the surface activity of PMMA, a great range of laser wavelengths were successfully evaluated. These included laser sources, such as CO2 (λ 5 10600 nm), Nd:YAG (λ 5 1064 nm), high power diode (λ 5 810 nm), and excimer (λ 5 248 nm) lasers. All laser treatments led to an increase on the wettability of the textured PMMA. However, excimer laser resulted as the most efficient treatment to reduce the WCA. This phenomenon was attributed to the increase of the O2 content of the treated surface due to the photo-oxidation mechanism involved during the process. An increase in the polar groups on the surface was also observed, which subsequently increases the wettability of the treated PMMA (Lawrence and Li, 2001). Utilization of fs lasers (Ti:Sapphire laser, λ 5 800 nm, pulse duration of 150 fs) was explored by De Marco et al. (2010) to produce microchannels for microfluidic applications. An increment in the WCA (from around 20 degrees for the base material, up to 94 degrees) was observed after processing, cleaning, and drying the surface. The biocompatibility of these surfaces is believed to be unsuitable, but they can find applications as antibacterial treatments due to their hydrophobic character. The increment in wettability is attributed to the formation of a double-scale roughness, which turned the surface into an intermediate state between those explained by Wenzel (1936) and Cassie-Baxter models (Cassie and Baxter, 1944). No chemical changes were observed in the laser-treated areas. On the contrary, Wang et al. demonstrated that more hydrophilic surfaces can be obtained if the laser fluence is adjusted during the treatment of PMMA surfaces with fs lasers (Wang et al., 2011). WCA was tailored from 0 to 125 degrees by only increasing the laser fluence during the treatment. Hydrophobic surfaces were produced with laser fluences ranging from 0.4 up to 2.5 J cm22; however, hydrophilic surfaces were obtained for higher fluences. This was attributed to the formation of polar groups, such as O C, and OQC. These results suggest that fs

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laser radiation can be suitable for the increment of the biocompatibility of PMMS surfaces; however, this should be confirmed with in vitro tests.

2.5.8 OTHER THERMOPLASTIC POLYMERS The utilization of LST to modify the surface of other thermoplastics for biomedical applications was also addressed. In this sense, polytetrafluoroethylene (PTFE) was treated with UV laser radiation by Ahad et al. in a nitrogen-rich atmosphere (Ahad et al., 2015). The surface pattern and chemical modification of the surface was observed after treatment. An increment in the O:C ratio and a reduction in the F:C ratio was observed with the number of laser pulses. An improvement in cell adhesion was observed for the laser-treated surfaces. Moreover, cells exhibited a suitable cell morphology. Waugh et al. studied the LST of nylon 6,6 using a CO2 laser (Waugh et al., 2009). An effect on the cell viability (using osteoblast cells) in laser treated surfaces was observed. Cells covered a larger area in those surfaces treated with this laser radiation. Polyimide (PI) was also textured using different laser radiations (Gu¨nther et al., 2016). Nd:YAG laser radiation (λ 5 1064 nm) as well as doubling (λ 5 532 nm), tripling (λ 5 355 nm), and quadrupling (λ 5 266 nm) fundamental laser frequency were used to create different patterning topographies. The main aim of these topographical features was to prevent bacterial adhesion and biofilm formation on the surface of indwelling medical devices. Different topographies with different periodicities were produced and the adhesion of Staphylococcus aureus (one of the most relevant gram-positive bacterial pathogens that can form biofilms on medical devices such as catheters, valves, prostheses, and implantable venous access systems) was determined. It was observed that lineand pillar-like patterns promoted S. aureus adhesion, whereas complex lamella micro-topography reduced the adhesion in static and continuous flow culture conditions. Furthermore, lamella-like textured surfaces retained the capacity to inhibit S. aureus adhesion both when the surface was coated with human serum proteins and when the material was implanted subcutaneously in a foreign-body associated infection model.

2.6 CHALLENGES AND FUTURE TRENDS Results presented in this chapter demonstrates the potential of LST to increase the biological performance of thermoplastics and composites. However, as can be deduced from data previously exposed, several issues remain unanswered. Most of the research found in the literature did not perform in vivo tests to corroborate the superior biological performance of laser-treated thermoplastics. A lot of work has been done in vitro and only showing the better cell responses, but no studies

2.7 Conclusions

in animals or clinical trials were performed. Another important issue is not completely related to this laser technique. Most of the experiments only irradiate a thermoplastic with a certain laser radiation to create a pattern; then, the surface topography, chemistry, and cell response are analyzed. However, the optimum pattern for each biomedical application was not studied. This question is related to how certain cell responses can be elicited as a function of the surface of the material (Bettinger et al., 2009). Another connected issue to be investigated in future works is that most of the research only analyzed the influence of the micro-roughness produced by LST. This technique is also able to modify the topography and chemistry at the nanoscale. Therefore, it should be addressed if the surface chemistry is modified at the nanoscale after the laser treatments, and if its heterogeneous modification can be interesting for biomedical applications (e.g., cell guiding or aligning). Currently, only common polymers, but not composites, have been lasertreated. Studies in more thermoplastics, and also in composites, with the available laser sources in the market should be performed. Then general conclusions—for example, on the most suitable laser wavelengths to modify the surface roughness, chemistry, or even both—could be drawn. Works found in the literature also showed that no studies on the simultaneous treatment with different laser wavelengths (e.g., one suitable for the modification of the surface topography and another suitable for the modification of the surface chemistry) were performed. Future works can also be focused on the modification of the surfaces of thermoplastics not only for enhancing the cell response, but also to prevent bacteria adhesion. Finally, a very interesting research line could be the computer design of surfaces to elicit determined cell or bacteria responses for any thermoplastic or composite. The output of this design would be the processing parameters required to apply to that material for any selected biomedical application. This task needs more simulation work on the LST of thermoplastics—this has been mainly addressed by Conde et al. using FEM simulations (Conde et al., 2011). In this regard, both the surface roughness and chemistry should be properly addressed by these kind of simulations.

2.7 CONCLUSIONS This chapter shows the capabilities of LST to enhance the biological performance of thermoplastics, which could potentially be applied to composites. Surface topography and chemistry of thermoplastics can be simultaneously modified for biological purposes. Moreover, this technique shows great processing advantages, such as, versatility, high throughput, excellent resolution, or the nonutilization of toxic chemicals, among others. This technique makes it possible to deal with the intrinsic bio-inertness presented in these materials and functionalizing their surfaces by inducing physicochemical modifications. Results demonstrate the great

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potential of LST to address some drawbacks exhibited by thermoplastics or composites used within the field of biomedical implants; however, there is still a long way to go to introduce this technique into the biomedical industry.

ACKNOWLEDGMENTS This work was partially supported by the EU research project MARMED (Marmed 2011-1/ 164), Government of Spain (MAT2015-71459-C2-P, FPU13/02944, and FPU15/04745 grants), and by Xunta de Galicia—ED431B 2016/042 (GPC) and Plan I2C Grant Program POS-A/2013/161, ED481B 2016/047-0, ED481D 2017/010.

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CHAPTER

Light-mediated thermoset polymers

3

Meenu Teotia1, Alok Mittal2 and Rakesh Kumar Soni1 1

Department of Chemistry, Chaudhary Charan Singh University, Meerut, India 2Department of Chemistry, Maulana Azad National Institute of Technology, Bhopal, India

3.1 INTRODUCTION Light-mediated thermoset polymers are crosslinked insoluble polymer networks which are obtained through the photopolymerization of viscous, liquid-state compositions on exposure to suitable radiation. Irreversible chemical bonds are formed and provide excellent mechanical properties, making the cured polymers ideal for high-performance applications. Radiation curing technology is an environmental friendly technique which transforms monomers and oligomers into macromolecules through a light-induced chain reaction and possess faster curing rates (Soni et al., 2010; Teotia et al., 2018). One of the key factors associated with radiation curing is that there are no volatile emissions during polymerization and, hence, stringent pollution controls are not required. Light-mediated thermosets can be readily obtained using room temperature operations, so eliminating the need of high temperature and pressure conditions (Chandra and Soni, 1994). Acrylate and methacrylate groups have been widely used in photopolymerization processes which are extensively employed in various industrial applications (Chandra and Soni, 1993). The practice of using photopolymerizable materials in biomedical applications is not that new and continuing development in the direction of new and improved biomaterials and smart fabrication techniques over the years have established a need for radiation processes in the biomedical field. The pioneering work by Bany et al. in 1985 reported UV curable compositions comprising monourethane acrylates of polysiloxane alcohols and copolymerizable ethylenically-unsaturated monomers for the fabrication of ophthalmic devices. The compositions were found suitable for the fabrication of contact lenses with improved oxygen transmissivity and hydrolytically stability. In 1986, Rosiak et al. (1989) developed the first radiation-cured, hydrogel-based, wound-dressing material. Hydron, Vigilon, Intrasite, Gelperm, and Biolex are some of the materials used in wound dressings containing a hydroxyethyl methacrylate

Materials for Biomedical Engineering: Thermoset and Thermoplastic Polymers. DOI: https://doi.org/10.1016/B978-0-12-816874-5.00003-7 © 2019 Elsevier Inc. All rights reserved.

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monomer with polyethyleneglycol, agar, or acrylamide. The use of radiationcurable, polyurethane (PU)-based, bioabsorbable and biocompatible polymers in the fabrication of surgical devices (particularly wound closure devices such as surgical staples and clips) in a stereolithography apparatus is well-documented (Regula et al., 1997). Suturing, stapling, and adhesive strips have been the common methods for wound closing, but they cause pain to the patients and also need to be removed when the materials used are not biodegradable. Photocurable bioadhesives are able to overcome these disadvantages and have emerged as efficient biomaterials with fast curing rates, controlled polymerizations, and can be applied to even weak and diseased tissues (Benson, 2002). The use of N-vinylpyrrolidone in the preparation of UV-curable bioadhesives having suitable adhesive strength has been reported by Kao et al. (1997). Polycaprolactone-based, UV-curable, and biodegradable adhesives have also been prepared with improved surgical applications (Gila et al., 2011). Biomedical sensors have been developed using photosensitive polyurethane encapsulant formulations microstructured by photolithography for biomedical and clinical applications (Munoz et al., 1997). Reports are also available on the use of photopolymerized acrylic polymers for the fabrication of foldable, highly refractive index ophthalmic device materials such as contact lenses, intracorneal lenses, and intraocular lenses (LeBoeuf and Karakelle, 1999). UV-curing technology has been successfully applied to modify the polymeric surfaces with acrylic coatings, providing lubricious and antimicrobial surfaces to fabricate clinically advantageous medical devices (DiTizio and DiCosmo, 2004). Hydrophilic macromers have been modified through free radical polymerization in presence of longer wavelength UV or visible light for the encapsulation of cells, as tissue adhesives, as barriers to prevent the interaction of one cell tissue, as plugs, and in drug delivery devices (Hubbell et al., 2002). Strong and nonimmunogenic thin films that are permeable to gases and nutrients can be obtained through photopolymerization, which can be applied to a variety of surfaces with distinct geometries (Hubbell et al., 2003). Photocurable endoprosthesis systems have also been developed for the fabrication of prosthetic devices to restore or enhance the flow of fluids through body lumen or ducts (Williams et al., 2006). The modification of biopolymers with photocrosslinkable functional groups has attracted much attention from research teams all over the world over the past decade. The modified biopolymers are being used extensively in the synthesis of hydrogels, which exhibit enormous possibilities in tissue engineering and drug delivery systems (Calo and Khutoryanskiy, 2015). Hydrogels with improved control on properties, flexibility, and purity can be produced readily through the radiation photopolymerization of natural and synthetic hydrophilic polymers (Lugao and Malmonge, 2001; Lugao et al., 2002). Natural polymers such as gelatin, starch, dextran, chitosan, haluronic acid, alginates, gellan gum (Pacelli et al., 2015) etc. are being modified into photocrosslinkable materials and used in hydrogels synthesis. Their biocompatibility, flexibility, high permeability to water, and resemblance to living tissues makes them highly applicable in

3.2 Types of Light-Sensitive Polymers

biomedical applications. With the advent of stereolithography in clinical applications such as patient-specific models and functional parts implantable devices can be tailored specifically for the patient which reduces the risk involved in surgical operations (Melchels et al., 2010). New developments in resin chemistry with highly desirable properties such as biodegradability, noncytotoxicity, and favorable mechanical properties, etc., have promoted the use of photofabrication techniques in the direct fabrication of implants. The past two decades have witnessed light-mediated thermoset polymers as potential applicants in biomedical applications. Various reviews summarized the promising applicability of these radiation curing processes in drug delivery systems, tissue engineering, and 3D bioprinting (Ifkovits and Burdick, 2007; Li et al., 2012; Pereira and Bartolo, 2015). In this chapter an attempt has been made to present a comprehensive account of the properties and applications of some light-sensitive polymers employed in biomedical field. The chapter is divided into six sections. The first section introduces the subject matter to the readers and second section details the systematic evolvement of photopolymerizable polyurethanes, compositions based on polyacrylates-methacrylates, and a few light-sensitive vinyl monomers such as N-vinylpyrrolidone, vinyl carbonates, and vinyl esters. Photoinitiators used to initiate photopolymerization and the involved mechanisms have been described in the third and fourth sections. Properties of photoinitiators mainly used in biomedically advantageous reactions have been discussed and compared for the benefit of the readers. In section five, modified biopolymers and their synthetic routes along with their applications have been highlighted. In the past few years, the focus has been shifted to the use of longerwavelength UV or visible-light radiations in light-mediated reactions. Some recent photopolymerizations carried out in the presence of longer wavelength radiations are discussed in section six.

3.2 TYPES OF LIGHT-SENSITIVE POLYMERS Light-sensitive polymers contain functional groups which can be photocrosslinked on exposure to suitable radiations to obtain cured polymer networks. Generally, carbon
carbon double bonds are involved to set up chain reactions in free radical polymerizations. Photocurable acrylated polyurethanes, acrylates/methacrylates, and vinyl monomers are discussed next.

3.2.1 POLYURETHANES Polyurethanes are a popular choice of materials to be used in the fabrication of medical devices such as pacemaker lead insulators, heart valves, cardiac assist devices, artificial hearts, lungs, pancreas, blood tubing, filters, vascular prostheses, many types of catheters in various clinical applications, and tissue

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replacement and augmentation (Lamba et al., 1998). They are very promising synthetic biomaterials because of their excellent mechanical properties and biocompatibility. The reaction of an isocyanate containing two or more isocyanate groups with a polyol with a minimum of two or more hydroxyl groups in the presence of UV radiation or a suitable catalyst produces polyurethanes. They are heterogeneous polymers containing urethane linkages in which other functional groups can also be incorporated to achieve the desirable mechanical and biological properties. Polyurethanes offer a wide spectrum of applications based on the selection of isocyanates, polyols, and chain extenders used in their synthesis. Isocyanates contributing to hard domains and soft domains are due to polyols in the segmented structure of polyurethanes. The functionality of the polyol determines their mechanical properties, ranging from flexible polyurethanes to rigid polymers. Both aliphatic and aromatic isocyanates can be used to synthesize polyurethanes. Hexamethylene diisocyanate, toluene diisocyanate, methylene diphenyl diisocyanate, and its polymeric derivatives etc., are the commonly used diisocyanates in their synthesis. Polyols may be polyesters, polyethers, polycarbonates, hydrocarbons, polycaprolactones, polysulphides, polydimethyl siloxanes and polyols containing fluorine and phosphorous (Soto et al., 2014a), and polyols recovered from renewable resources, etc. (Petrovic, 2008). The low functionality and high molecular weight polyols are suitable for flexible applications; however, the high functionality and low molecular weight of polyol results in rigid polymers. Urethanes are mostly introduced in radiation curable formulations as urethane acrylates. Generally, a hydroxyl functional acrylate is used to introduce acrylic moieties in the polyurethane backbone which is responsible for light-mediated polymerization reactions. UV curable polyurethanes are extensively applied in coating applications, wood coatings, paper coating, architectural glass lamination, packaging, machinery, furniture, bedding, and carpets, etc. (Romaskevic et al., 2006; Teotia and Soni, 2018). In this chapter, we limit our discussion to the UVcurable compositions containing urethane acrylates widely used in biomedical applications. The next section covers the systematic growth in photopolymerizable compositions of urethane acrylates with different modifications in the selection of reactants and methods to get improved biocompatibility of the UV cured polymer.

3.2.1.1 Synthetic routes to photocrosslinkable polyurethane polymers Polyurethanes are applied extensively as biomaterials and, hence, the properties of the cured polymer like tensile strength, elongation, blood compatibility, surface chemistry, roughness, hydrophilicity, etc. influence their biocompatible nature. UV-curable urethane acrylates possess lower ultimate extensions than thermoplastic polyurethanes because of their high crosslink density. The flexibility is an important characteristic of polyurethanes for their wide spectrum biomedical applications. Reports are available on the development of high-performance urethane acrylates with superior elongations and mechanical properties through

3.2 Types of Light-Sensitive Polymers

deblocking chemistry (Velankar et al., 1996). Urethane acrylate was UV-cured and the acrylate tipping agent was removed through annealing leading to lowdensity, acrylic polymers with incorporated thermoplastic polyurethanes in the network. Urethane acrylate oligomers can be synthesized either by direct addition of diisocyanate-terminated polyol on each end with a hydroxy functional acrylate or by reverse addition in which the product of the reaction of diisocyanate and hydroxy-functional acrylate can be used to terminate a polyol (Fig. 3.1) (Swiderski and Khudyakov, 2004). Epoxy urethane acrylates can be prepared from modified epoxy resins and diisocyanates in two-stage polymerization. Initially the epoxy resin was modified with acrylic acid to get the acrylic moieties at both ends and, further, the pendant hydroxyl groups were reacted with diisocyanate to obtain urethane acrylate (Oprea et al., 2000). The literature reports the

FIGURE 3.1 Steps involved in photocuring urethane acrylate.

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synthesis of hyperbranched urethane-acrylates based on aliphatic-hyperbranched polyesters and polyethyleneglycol acrylate with improved mechanical properties and solvent resistance (Tasic et al., 2004). The prepared hyperbranched-urethane acrylates (HB-UAs) are free from oxygen inhibition and were found to be very reactive when evaluated for use in UV-curable coatings. Further, HB-UA oligomers containing soybean fatty acids were also synthesized from 2,2-bis (hydroxymethyl) propionic acid, and di-trimethylol propane and a certain number of hydroxyl end groups were modified with soybean fatty acids (Dzunuzovic et al., 2005). These partially modified hyperbranched polyesters (HBP) were reacted with diisocyanate adduct with 2-hydroxyethyl acrylate in different ratios to get urethane acrylates. UV-curable, waterborne, hyperbranched polyurethaneacrylate dispersions of varying generation number have been developed and investigated for particle size, photopolymerization kinetics, and thermal stability (Asif et al., 2005). Concentrations of acidic groups, acrylate groups, and the functionality of the hyperbranched polyesters were optimized for these properties. The cured polymers were found to possess good thermal stability as no weight loss was observed until 200  C. The water resistance of the membranes fabricated from UV-curable waterborne polyurethane dispersions can be improved by using polyester polyols, minimizing the COOH content, and increasing the gel content by increasing a higher CQC level. Polyurethane with hydroxyl-terminated polybutadiene or modified with dihydroxybutyl-terminated polydimethylsiloxane offers superior water resistance to developed membranes (Bai et al., 2007). Yang et al. (2009) developed UV-curable PU aqueous dispersions by blending multifunctional thiol and ene terminated polyurethane aqueous dispersions with a superior shelf life than nonaqueous PU dispersions. Low-oxygen inhibition was observed while curing and the cured PU coatings showed improved physical properties, solution stability, and photopolymerization behavior. UV-curable urethane prepolymer and hydrophilic monomers such as 2-hydroxyethyl methacrylate, N-vinyl pyrrolidone, and glycerol methacrylate have been used in the preparation of hydrogels to get higher oxygen permeability and tear strength than polyacrylates due to the presence of urethane linkages (Lai and Baccei, 1991). Polyurethane-based light curable elastic hydrogels can be synthesized from polycaprolactone diol, polyethylene glycol, lysine diisocyanate, and 2-hydroxyethyl methacrylate through a UV light initiated polymerization reaction (Zhang et al., 2007). Polyurethanes with different soft segmental structures, hydrophilicity, and cytophilicity can be obtained by varying the polycaprolactone to polyethylene glycol ratio during prepolymer synthesis. Long-term contact of polyurethanes with blood leads to the absorption of plasma protein which results in platelet adhesion and clot formation. To check this thrombogenicity, researchers have focused on modifications in PU compositions. Extensive work has been conducted on the incorporation of fluorine atoms in PU networks to improve surface properties enhancing their biocompatibility to obtain more thromboresistant materials (Chen and Kuo, 2000; Wang and Wei, 2005). Lin et al. (2008) developed UV-curable fluorinated poly(urethane-acrylate)

3.2 Types of Light-Sensitive Polymers

oligomers from 1H,1H,12H,12H-perfluoro-1,12-dodecanediol, 1,6-hexamethylene diisocyanate/4,40 -diphenylmethane diisocyanate, and 2-hydroxyethyl methacrylate for end-capping with photocrosslinkable methacrylate groups. The authors determined their relative index of platelet adhesion and found superior blood compatibility. The hydrophobic fluorocarbon chains incorporated in the PU network led to a reduction in the adhesion of blood platelets onto the materials due to phase separation and a low total-surface energy. Easily injectable photopolymerizable urethane acrylates have been synthesized from poly(propylene glycol) or poly(caprolactone diol) and hydroxyethyl methacrylate which can be molded easily into any desired shape for biomedical applications (Pereira et al., 2010). The authors performed in vitro as well as in vivo experiments in order to investigate the biocompatibility of the cured polymer. The mechanical properties of cytocompatible PUAs were found comparable to soft tissues and no significant inflammatory reaction was observed through histology cross sections. Efforts have been made to enhance the thermomechanical properties of the polyurethane acrylates by incorporating polyhedral oligomeric silesquioxane through free radical photopolymerization (Kim et al., 2009). Low surface energy films were obtained with increased flexibility, Tg, thermal stability, and dimensional stability. Photocurable, waterborne, polyurethane acrylate has also been modified with octavinyl polyhedral oligomeric silsesquioxane to improve water resistance and thermal oxidative stability (Wang et al., 2011). The modified polymer was obtained through photopolymerization between the two materials having acrylate functionalities. This work demonstrated the potential use of lightmediated polymers in biomedical applications. Waterborne polyurethane acrylates based on isophorone diisocyanate, polyether polyol, dimethylol propionic acid, and hydroxyethyl methyl acrylate with good solvent resistance characteristics can also be prepared in situ through a self-emulsifying method (Xu et al., 2012). In recent research, a novel, hyperbranched polyurethane acrylate (HBPUA) has been developed to be used as a crosslinker in radiation-curable compositions (Zhang et al., 2016b). The addition of HBPUA induced manifold improvements in mechanical properties of the cured polymer and the coatings showed excellent adhesion onto polycarbonate and polyvinylchloride sheets.

3.2.1.2 Radiation curing in surface modification The degradation of segmented polyurethanes can be induced on contact with activated macrophages when implanted as medical devices in living organisms over a period of time. Macrophases produce oxygen radicals which attack the soft segments in the PU network and cracking is initiated. Reduction in chemical stability and biocompatibility is a matter of great concern for the long-term usage of polymeric materials. Researchers have focused on the surface modifications of the segmented polyurethanes in order to present a solution to this problem. Graft photopolymerization involves coating UVcurable monomers or oligomers onto a polyurethane surface, which inhibits

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platelet activation. N-vinyl pyrrolidone, polyvinylpyrrollidone, polyhydroxyethyl methacrylate, and polyacrylamide have been grafted successfully on PU surfaces to enhance their biocompatibility (Lamba et al., 1998). N,N0 dimethyl aminoethyl methacrylate has also been used in photografting polyurethane surfaces to promote cell adhesion and growth and, hence, increasing their cytocompatibility (Guan et al., 2001). Feng et al. (2013) grafted poly(ethylene glycol) monoacrylates with molecular weights between 400 and 1000 g mol21 on the surface of polycarbonate urethane for increasing its hydrophilicity and improving its hemocompatibility through UV-initiated polymerization. PEGMAs grafting decreased platelet adhesion onto the PU surface significantly and a remarkable increase in hydrophilicity was observed which was further increased with the increasing molecular weight of the PEGMAs.

3.2.1.3 Shape memory polymers Shape memory polymers are extensively used in implantable biomedical devices, especially for vascular stents, tissue scaffolds, and clot-removal devices (Volk et al., 2011). Tunable material properties, excellent flexibility, and the biocompatibility of polyurethanes makes them interesting partners in SMPs biomedical applications (Small et al., 2010; Lendlein et al., 2010). Shape memory materials are elastic polymer networks able to recover their initial shape after deformation upon exposure to suitable stimuli sensitive switches incorporated in the polymer network. Hard domains in the polymeric network act as net points which determine their permanent shape and can be chemical or physical in nature; however, soft segments act as molecular switches. The molecular switches may be thermally induced or light induced, the later stimulation is independent of temperature effects and can be operated in light-sensitive polymers (Sobczak, 2015). Generally, cinnamic acid or cinnamylidene acetic acid moieties are introduced either through grafting or creating interpenetrating networks to work as light-stimulated switches. The polymer is first stretched and irradiated with a particular wavelength leading to formation of covalent links (photoreversible) upon [2 1 2] cycloaddition reaction forming a cyclobutane ring and further irradiation with different wavelengths results in ring cleavage to recover the permanent shape (Lendlein et al., 2010). Thermal stimulus is widely used in the fabrication of shape-memory biomedical devices in which additional crosslinks are cleaved through differences in temperature. Nair et al. (2010) developed photopolymerized thiol-ene and urethane thiol-ene systems to be used as shape-memory polymers for the fabrication of medical devices. These systems have homogeneous networks, involve low volume shrinkage, and are free of oxygen inhibition. Glass transition temperatures of the polymer systems was in the range of 30
40  C with enhanced toughness making them suitable for a broader range of biomedical, shape-memory applications involving thermal stimulation.

3.2 Types of Light-Sensitive Polymers

3.2.2 POLY-ACRYLATES/METHACRYLATES The term “poly-acrylates/methacrylates” means photopolymerized homopolymers, copolymers, or terpolymers of ester derivatives of acrylic acid and methacrylic acid (Fig. 3.2). The photopolymerization of bifunctional/multifunctional, acrylate/ methacrylate monomers and oligomers generates highly crosslinked networks with excellent physical, mechanical, and thermal properties, which make them suitable for a variety of applications (Teotia and Soni, 2016). Such UV-cured polymers find huge applications as abrasion-resistant coatings for eye glass lenses and optical fibers, in aspheric lenses, dental restorative materials, second harmonic generating materials, and hydrogels, etc. (Anseth et al., 1994). These are the major partners in fabrication of a variety of medical devices like medical spikes, cassettes, catheter accessories, urological accessories, adapters, pump housing, filter housings, check valves, chest drainage units, and blood-handling

FIGURE 3.2 General structures of: (A) acrylate; (B) methacrylate; (C) diacrylate; (D) dimethacrylate; and (E) acrylate-methacrylate.

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components. Due to their high mechanical performance, UV-polymerized polyacrylates are also used in pressure-sensitive adhesives in medical tapes which are latex and solvent free with minimized migrating ingredients, and antiallergic, that is, they do not cause skin irritation and are also noncytotoxic. Generally, pure acrylate compositions possess high-crosslink density, potential cytotoxicity, and skin irritancy. When we talk about photopolymerized biomaterials, their biocompatibility, hemocompatibility, and surface properties are of prime importance for their potential use in biomedical applications. Researchers have focused on developing modified compositions by incorporating biologically advantageous functionalized monomers and oligomers to utilize the advantages of photoclick reactions and, at the same time, offering biomaterials with potential end uses. In the next section, we will discuss acrylated/methacrylated photopolymerizable compositions including homopolymers and optimized mixtures of different functionalized monomers and oligomers.

3.2.2.1 As dental materials Methacrylate based photopolymers are used extensively in dentistry due to their excellent mechanical properties. They are the major partners among various commercially available dental resins for direct restoration (Leprince et al., 2013). Researchers have investigated different photocurable, commercial, dental methacrylate-based composites with the help of photocalorimetry and thermogravimetry (Malhotra and Mala, 2010; Gorsche et al., 2014; Strassler, 2011). The final properties of the cured material depend on the curing rate, heat of reaction, and filler content, which varies from one composition to other. One important desirable property of a material to be used in dental applications is its “low shrinkage” during and after photocrosslinking to avoid any harm while protecting the surrounding tissues. Bland and Peppas (1996) developed photopolymerized multifunctional (meth)acrylates as model polymers for dental applications. Different acrylates such as 1,1,1-trimethylolethane trimethacrylate 1,1,1-trimethylolpropane triacrylate, 1,1,1-trimethylolpropane trimethacrylate, ethylene glycol dimethacrylate, and triethylene glycol dimethacrylate were photopolymerized with an epoxy diacrylate in the presence of dimethoxy-2-phenyl-acetophenone photoinitiator. The shrinkage volume of the cured polymer decreased with an increase in the size of the pendant groups and it was less for methacrylates than acrylates, making the former superior for dental applications. Polydimethacrylate resins used in dental composites can be prepared by the photopolymerization of neat monomers and mixtures of them with varying ratios in the presence of a camphorquinone/N,N-dimethylaminoethyl methacrylate system as a photoinitiator with very low water-uptake values (Sideridou et al., 2004). Triethyleneglycol dimethacrylate and bisphenol-A-glycidyl-dimethacrylate (Bis GMA) are widely used in dental restorative resins with high conversions (65%), but suffer from polymerization shrinkage. Lu et al. (2005) developed Bis GMA/monomethacrylate with secondary functionalities systems and reduced shrinkage, rapid curing, and higher conversions. The authors reported Bis

3.2 Types of Light-Sensitive Polymers

GMA-morpholine carbamate methacrylate system to develop superior dental resins with 86% final conversion, 3.5 times faster polymerization rate, and also polymerization volumetric shrinkage was reduced to 30%. Gatti et al. (2007) developed copolymers of Bis GMA, triethyleneglycol dimethacrylate, and urethane dimethacrylate to prepare dentistry resins. In an another work, a visible light curable urethane acrylate/tripropylene glycol diacrylate oligomer was synthesized to be used as a root canal sealer (Hsieh et al., 2008). It was prepared by reacting 2-hydroxyethylmethacrylate with isophorone diisocyanate in a 1:1 mole ratio and found to be compatible with ZnO/thermoplastic polyurethane filling material due to the enhanced interaction between urethane linkages. The literature reports the synthesis of a dimethacrylate monomer 5,50 -bis[4-(20 -hydroxy-30 methacryloyloxy-propoxy)-phenyl]-hexahydro-4,7-methan-oindan and its composition with tri(ethylene glycol) dimethacrylate which was investigated for contact angle, diffusion coefficient, water sorption and solubility, degree of double bond conversion, shrinkage, flexural strength, and modulus (He et al., 2010). The authors compared this system with the Bis GMA/TEGDMA system and found lower volume shrinkage and higher flexural strength and modulus of the photocured resins than the Bis GMA reference. Another effort to replace Bis GMA in dental compositions has been put forth by Liu et al. (2013a) by introducing a novel tertiary amine-containing urethane dimethacrylate monomer as one component of dental materials. The incorporation of dimethacrylate monomer (25% of Bis GMA) raised the double-bond conversion and reduced polymerization shrinkage when compared to Bis-GMA/TEGDMA resin. The new copolymer possesses superior mechanical properties and higher water solubility and sorption. Also in 2013, Acosta Ortiz et al., developed a new composition to reduce the photopolymerization shrinkage of a dental resin based on dimethacrylates. The authors incorporated five and six-membered ring spiroorthocarbonates functionalized with allylic groups as antishrinkage additives which increased the crosslink density via ring opening polymerization. The shrinkage was decreased to 53% with improved double bond conversion and compressive and flexural strength. New dimethacrylates containing urethane linkages with quaternary alkyl ammonium and polyethylene glycol short sequences have also been prepared and investigated for their applicability as dental materials. A composition involving cationic dimethacrylate, Bis GMA, and TEGDMA showed high polymerization rates; however, the polymerization shrinkage increased slightly within acceptable limits (Buruiana et al., 2014).

3.2.2.2 As hydrogel biomaterials 2-Hydroxyethyl methacrylate is one of the oldest UV curable monomer which is widely used in synthesis of biomaterials (Gibas and Janik, 2010). The properties of poly(hydroxyethyl methacrylate) depends upon the synthesis methods, degree of crosslinking, and temperature conditions and may be modified with other natural and synthetic materials. It is a well-known hydrogel material used in biomedical applications (Trimaille et al., 2016). Poly (hydroxyethyl methacrylate)

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can be obtained through photopolymerization in the presence of UV radiation by reacting 2-hydroxyethyl methacrylate with polyethylene glycol dimethacrylate and benzoin isobutyl ether as a crosslinking agent and photoinitiator, respectively (Young et al., 1998). 1,1,1-trimethylol propane trimethacrylate and diethylene glycol dimethacylate (Li and Lee, 2005) have also been used as crosslinkers for the synthesis of PHEMA. Main applications of PHEMA include wound dressings, contact lenses, artificial skin, drug delivery systems, cell regeneration, tissue scaffolds, and artificial cartilage, etc. The effect of the chain length of the crosslinker has also been investigated by using ethylene glycol dimethacrylate, diethylene glycol dimethacylate, triethylene glycol dimethacrylate, tetraethylene glycol dimethacrylate, and neopentyl glycol dimethacrylate as crosslinkers in UV photopolymerization of HEMA in the presence of Benacure 1173 as photoinitiator (Ananthoji, 2012). The water absorption capability of hydrogels depends on the pore size in the crosslinked network which increases with an increase in the chain length of the crosslinker. PHEMA can be modified in numerous ways to obtain choice-based biomaterials. A new magnetic rotational spectroscopy method has been developed for estimating the conversion of the CQC bond during photopolymerization of HEMA, which will be highly beneficial for the in vivo mapping of the rheological properties of biofluids/ biopolymers (Tokarev et al., 2012). Hydrogels based on polyethylene glycol diacrylates and dimethacrylates are used extensively in drug delivery systems and tissue scaffolds owing to their excellent biocompatibility, water solubility, and nontoxicity (Underhill et al., 2007). Lin-Gibson et al. (2004), presented synthesis methods and detailed characterizations of biocompatible and photopolymerizable poly(ethylene glycol) urethane-dimethacrylates and poly(ethylene glycol) dimethacrylates. Hydrogels of these dimethacrylates were prepared by exposing their aqueous solutions to suitable radiation in the presence of I2929 photoinitiator and seeded with bovine chondrocytes to assess their biocompatibility. The cells were completely viable even after 2 weeks. Murtezi et al. (2015) have used poly(ethylene glycol) methyl ether methacrylate and poly(ethylene glycol) diacrylate in the synthesis of clickable hydrogels. Poly(methyl methacrylate) with alkyne terminal groups was prepared by photoinitiated free radical polymerization of methyl methacrylate using camphorquinone as the photoinitiator and 3-(trimethylsilyl)propargyl alcohol as the hydrogen donor. Poly(ethylene glycol) methacrylate has also been used to improve the biocompatibility of UV-crosslinked, electrospun, hybrid nanofibers used for the fabrication of artificial vascular scaffolds (Wang et al., 2012). The monomer was incorporated in a polyurethane solution and crosslinked in the presence of N,N0 -methylene bisacrylamide and benzophenone. PEGMA-PU scaffolds showed better cell morphology and proliferation than pure PU scaffolds. Recent research reports the use of photopolymerized methacrylate platforms patterned with microtopographical features to precisely guide Schwann cell alignment and neurite growth (Li et al., 2015). Photocrosslinkable poly(ethylene glycol) acrylate, poly(ethylene glycol) diacrylate, and gelatin methacrylate-based hydrogel

3.2 Types of Light-Sensitive Polymers

formulations have also been used as coating materials for electrospun polycaprolactone (PCL) fibrous mats. Photopolymerization was carried out in the presence of Irgacure 2959 to get coated fibrous matrices for fabrication of vascular implants (Correia et al., 2016). More recently, Schmocker et al. (2016) reported a photopolymerizable poly(ethylene glycol)dimethacrylate nano-fibrillated cellulose composite hydrogel to be used as functional orthopedic implants for the replacement of nucleus pulposus. A customized, minimally invasive, medical device was used for injection and in situ photopolymerization of hydrogels into an intervertebral disc of a bovine organ model to evaluate its mechanical resistance, which was found suitable for long-term usage.

3.2.3 LIGHT-SENSITIVE VINYL MONOMERS Another class of light-sensitive materials with potential applicability in biomedical field is of vinyl monomers conjugated to different functional groups making them accessible to absorb in UV-Visible region. Vinyl monomers are generally less cytotoxic than pure acrylates/methacrylates (Liska and Husar, 2012) and various reports are available that focus on the use of these monomers in photopolymerizable formulations as alternatives to toxic and irritant acrylate monomers. Some of the green monomers with low toxicity, such as N-vinyl pyrrolidone, vinyl carbonates, and vinyl esters, have been discussed in this section.

3.2.3.1 N-Vinyl pyrrrolidone Polyvinyl pyrrolidone is a hydrosoluble, bioinert, and synthetic polymer usually obtained through radical polymerization of N-vinyl pyrrolidone which found numerous applications in the synthesis of hydrogels for biomedical applications. PVP hydrogels are advantageous because of their simple formation and lowproduction costs, softness and flexibility, and their ability to absorb large amounts of liquid while retaining good mechanical properties. Kao et al.(1997) prepared a series of UV-curable bioadhesives from copolymers of N-vinyl pyrrolidone with 2-acrylamido methyl 1-propane sulfonic acid, glycidyl acrylate, vinyl succinimide, and 2-isocyanatoethyl methacrylate. These formulations were cured using UV radiation with a set time of about 3 minutes. The authors determined the bond strength of the copolymers in a porcine intestine specimen and found adhesion values up to 4.6 N m21 and that they could be applied as single-layered, wound dressings and bioadhesives. N-vinyl pyrrolidone can also be photocrosslinked to provide low-friction, hydrophilic hydrogel coatings on medical devices such as catheters and guide wires, etc. (Madsen, 1998; Madsen et al., 2010). Photocrosslinking of hydrogels through UV radiation is more advantageous than the use of high energy radiation such as gamma radiation or electron beam (Razzak et al., 1999; Can, 2005). Lopergolo et al. (2003) presented an option of UV radiations to direct photocrosslinking in aqueous solutions using a lowpressure mercury lamp. The polymer obtained through UV curing possessed similar micro- and macroscopic properties when compared to high-energy radiation

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hydrogels and presented no cytotoxicity. This approach was further improved by Fechine et al. (2004), by using photo-Fenton reactions and hydrogen peroxide to accelerate the UV crosslinking process of aqueous solutions of PVP. Polyvinyl pyrrolidone is generally blended with other polymers such as polyethylene glycols, agar, cellulose, and glycerol etc., in order to enhance their biocompatibility and mechanical properties and to widen the application spectrum. PVP and PEG hydrogels are elastic, sterile, and noncytotoxic and are widely used in wound dressings (Lugao et al., 2002). The use of agar in these hydrogel formulations can facilitate the growth of microorganisms; however, the ratio of PVP and PEG can be tuned to obtain microbe-resistant hydrogels (Ajji et al., 2005). Wang et al. (2007), incorporated cellulose or carboxymethyl cellulose to have superior flexibility than pure PVP hydrogels. Copolymers of methacrylic acid and N-vinyl pyrrolidone hydrogels have also been synthesized and evaluated for their use as carriers for oral protein delivery (Carr and Peppas, 2010). Nearly 90% loading efficiencies were found for ethylene glycol dimethacrylate crosslinked carriers. These systems were found ideal for oral delivery of therapeutic agents that involves transcellular mechanisms for transportation. N-Vinyl pyrrolidone has also been used in grafting polyhydroxy backbones, such as dextran or polyvinyl alcohol, to obtain materials which can be used for controlled drug release and plasma substitution or expansion (Brunius et al., 2002). PVP has been connected to the polyol chain through carbonate or ester linkage which is hydrolytically labile and hence the copolymers are potentially degradable. Liu et al. (2013b) have reviewed PVP-modified surfaces, their design approaches, preparation, and potential biomedical applications. The authors reported PVP as a potential antifouling surface modifier which has comparable properties to poly(ethylene glycol). A glucocorticoid prodrug based on poly (N-vinyl pyrrolidone) has also been synthesized as a neural interface and its bioactivity and release of free drugs have been investigated with RAW264.7 macrophages in vitro (Cao and He, 2010). The prodrug is nontoxic to neurons, able to maintain its ability to manage neurite extensions, and can be applied further to reduce inflammation at the neural interface. Jansen et al. (2009), photocrosslinked functionalized, three-armed poly(D,L-lactide) oligomers, fumaric acid monoethyl ester, and N-vinylpyrrolidone in the presence of Irgacure 2959, exposing it to radiation of 365 nm wavelength, and intensity 3
5 mW cm22 for 15 minutes in a nitrogen atmosphere. The cured polymer networks were used in stereolithography to prepare tissue engineering scaffolds with optimized pore architecture and tunable material properties. Mixtures of acid aqueous cosolutions of PVP and chitosan can also be used to obtain hydrogels with UV radiation of 254 nm (Barros et al., 2011).

3.2.3.2 Vinyl carbonate Earlier reports demonstrate the use of vinyl carbonates in the fabrication of soft contact lenses, as drug carriers to release the active substances, for the

3.2 Types of Light-Sensitive Polymers

development of inflammatory drugs like dexamethasone and hydrocortisone through copolymerization with N-vinyl pyrrolidine (Nguyen and Galin, 1985), and for the modification of pharmaceutical drugs, etc. (Brosse et al., 1984). Boutevin et al. (1986) polymerized deuterated, fluorinated, and chlorinated vinyl carbonate monomers in the presence of Darocur 1173 to transparent and hard polymers. They possess glass transition temperatures approximately same to polyacrylates and polymethacrylates (Freire et al., 1988). The photoreactivity of these vinyl carbonates is comparable to acrylates and methacrylates and these are suitable for UV-visible curing. Over the years, researchers have emphasized the synthesis of functionalized vinyl carbonates with improved photoreactivities and desirable mechanical properties. Lai (1997), reported methacryloyloxyethyl vinyl carbonate as a crosslinker and incorporated hydroxyethyl methacrylate and N-vinyl pyrrolidine formulations to improve the copolymerization of UV-cured hydrogels. Difunctional vinyl carbonates have also been investigated for photoreactivities and double-bond conversions in the presence of Irgacure 819 photoinitiator. The monomers were found suitable for 3D printing parts using digital light processing methods at room temperature (Heller et al., 2011). Significant osseointegration of the 3D cellular scaffolds was indicated through in vivo tests making them suitable for the fabrication of implants. Vinyl carbonates upon irradiation generate free radicals which are more highly reactive than the initial monomer and readily abstract hydrogen to give low-reactivity radicals (Fig. 3.3). This step is responsible for their low photoreactivity which can be overcome by adding thiol-based monomers (Liska et al., 2012; Mautner et al., 2011). Efficient curing of vinyl carbonates has been performed using thiol-ene polymerization (Mautner et al., 2012). Thiyl radicals formed after donating hydrogen atoms have sufficient reactivity towards vinyl carbonates and boosts their photopolymerization. This research group developed vinyl carbonate-based photopolymerizable compositions which can be crosslinked to obtain biocompatible and biodegradable polymers with lower toxicity than methacrylates. Their degradation involves a surface erosion mechanism with the loss of carbon dioxide which is easily removed from the implantation site (Fig. 3.4). These properties of vinyl carbonates qualify them for biomedical applications. Russmueller et al. (2015) implanted in vivo polymerized vinyl carbonate 3D cellular structures into the distal femoral condyle of white rabbits. The authors found superior rates of newly formed bone (P , .001) and bone to implant contact (P , .001) making them promising biophotopolymers to replace methacrylates and polylactic acid-based materials. Researchers have also designed several vinyl carbonates by incorporating suitable functional groups to further enhance their mechanical properties. In recent research, Mautner et al. (2016) incorporated cyclic structures or urethane groups in vinyl carbonate structures and utilized the thiol ene concept to enhance the mechanical properties to the level of polylactic acid and with photoreactivity comparable to acrylates.

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FIGURE 3.3 Schematic representation of low photoreactivity of vinyl monomers.

FIGURE 3.4 Mechanism of degradation of polyvinyl carbonates.

3.2.3.3 Vinyl esters Vinyl esters are among the least explored vinyl monomers and limited monomers are available commercially because of the sensitivity of their synthetic routes toward the incorporation of different functional groups (Lee et al., 2005). Only few research reports have studied the photopolymerization behavior of vinyl

3.3 Photoinitiators

esters, such as divinyl fumarate (Wei et al., 2007), thioether linkages containing multivinyl ester monomers (Lee et al., 2004), and divinyl tri(ethylene glycol)bis (ophthalate) (Mironovich et al., 2005). Vinyl ester monomers have been used to replace acrylate and methacrylates in photopolymerization based additive manufacturing technologies for tissue engineering because of their exceptional low cytotoxicity. The final degradation product of photopolymerized vinyl esters upon hydrolysis is polyvinyl alcohol, which is nontoxic and a well-known pharmaceutical additive. These features make them more demanding in the fabrication of 3D medical implants. Heller et al. (2009) reports the synthesis of a series of low-toxicity, vinyl ester monomers from vinyl acetate using mercury acetate as a catalyst. They investigated in vitro cytotoxicity of these monomers with osteoblast cells and found them less cytotoxic than acrylates and methacrylates. These monomers were used for the fabrication of 3D cellular scaffolds for regenerative medicines and tissue engineering. Excellent biocompatibility was obtained through in vivo testing of stereolithographically fabricated 3D scaffolds by implanting them into the distal femoral bone of adult rabbits. Dworak et al. (2010) reported the synthesis of degradable and biocompatible phosphoruscontaining vinyl esters and investigated their hydrolytic erosion behavior, cytotoxicity, and mechanical stability for their potential use in biomedical applications. On incorporation of appropriate fillers such as hydroxylapatite with mechanical properties like natural bone can be obtained (Husar et al., 2011). The authors optimized the ratio of hydrophilic and hydrophobic monomers to tune the degradation behavior of photocured polymers. Low irritancy and cytotoxicity of vinyl ester resins makes them excellent alternatives to methacrylates in biomedical applications, but they possess less photoreactivity than methacrylates in the presence of abstractable hydrogens. Mautner et al. (2013), improved the photoreactivity of vinyl ester resins through the thiol-ene concept and used osteoblast cell culture experiments to confirm their low toxicity. Photocrosslinked vinyl ester and vinyl carbonate monomers have been further investigated in vitro on MC3T3-E1 cells and in vivo in a small animal model for white rabbits (Russmueller et al., 2015). The biochemical properties of these monomers were compared to standard acrylates and found to be suitable to overcome the shortcomings of conventionally used polymers, such as polylactic acid and methacrylates in 3D cellular structures. Threefold higher alkaline phosphatase activity and tenfold less toxicity was observed for these vinyl monomers, promoting their use in biomedical applications. In vivo polymerized vinyl esters showed quite good rates of new bone formation; however, less bone to implant contact was observed which suggests scope for further improvements in vinyl ester formulations and their adaptation (Qin et al., 2013).

3.3 PHOTOINITIATORS The role of photoinitiators is crucial to make radiation curing technologies more practical and applicable. Photoinitiators are the compounds which possess chromophoric groups capable of absorbing electromagnetic irradiation and have high

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extinction coefficient values in the desired region of absorption. They are the basic link between the radiation source and radiation curable systems. Photopolymerizable curing systems requires photoinitiators to absorb radiations of suitable wavelength and generate reactive species that must be capable of initiating the curing process. They can be used at ambient conditions without the need of specialized ovens or equipment. The curing process of UV-vis curable systems may involve a free radical mechanism or cationic mechanism. Photocuring is mainly based on free radical polymerization which involves three types of photoinitiators (Takimoto, 1993): (1) unimolecular, involving photofragmentation which generates radical pairs through a photoscission process; (2) bimolecular, involving hydrogen abstraction from donor molecules; and (3) electron transfer reaction. Photoinitiators involving unimolecular and bimolecular mechanism are also known as type I and type II photoinitiators, respectively, as the former involves α-cleavage and the latter abstracts hydrogen to initiate the curing process (Fig. 3.5). Benzoin ethers and Irgacures (CIBA) are examples of a type I photoinitiator; however, diketones such as camphorquinone, acenaphthenequinone, benzyl, and biacetyl etc., in conjunction with amines, are examples of type II photoinitiator systems. Generally, amines are highly toxic and mutagenic, hence, monomers containing amine groups such as N,N-dimethyl-p-toluidine, 4,40 -bis(N,N-dialkylamino) benzophenone (also known as Michelers’s ketone) and N,N-dimethylamino methacrylate, etc., are preferred. The photoinitiator system for a particular photopolymerizable formulation is selected by tuning its absorption spectrum and quantum yield efficiency with the radiation source. When we talk about photoinitiators

FIGURE 3.5 Photolysis modes of: (A) type I photoinitiator; and (B) type II photoinitiator.

3.3 Photoinitiators

in the context of photopolymers for biomedical applications, their toxicity is a major concern. Very careful monitoring of properties such as operational wavelengths, mutagenicity, and cytoxicity, etc., is essentially required. Many efficient photoinitiators have been developed and efforts are continuing to obtain more efficient photoinitiator systems which are no toxic and exhibit faster curing speeds and better photosensitivity for their use in biomedical applications. Some common photoinitiators which are being widely used in designing biomaterials will be discussed in the next section.

3.3.1 IRGACURE 2959 1-[4-(2-Hydroxyethoxy)-phenyl]-2-hydroxy-2-methyl-1-propanone (Irgacure 2959) is the most commonly used photoinitiator in hydrogel synthesis, cell encapsulations, and tissue engineering applications (Bryant et al., 2000; Fedorovich et al., 2009). It is a type I photoinitiator and its active hydroxyl group is able to react with the unsaturated sites of the UV curable systems. Low volatility, noncytotoxicity, low odor, and water solubility are some of the important features of I2959 which make it a potential photoinitiator for biomedical use. We have already discussed several reports in the previous section detailing photocrosslinking in presence of I2959. Its maximum wavelength is around 270 nm; however, researchers have used higher wavelengths (365 nm) to avoid damage to DNA, proteins, and encapsulated cells, etc. Williams et al. (2005) investigated the cellular toxicity of I2959 on six different cell populations that are generally used in tissue engineering and compared these with other photoinitiators. The authors concluded that I2959 was well-tolerated by different cell types in a range of mammalian species and proposed it as a suitable candidate for cell encapsulation techniques. Lower values of molar absorptivity at 365 nm and moderate water solubility (Table 3.1) limit the use of this photoinitiator where these parameters are essentially required. Ma et al. (2014) have modified I2959 into I2959-AA through an esterification reaction with acrylic acid. The authors further copolymerized I2959-AA with acrylic monomers to obtain a novel macrophotoinitiator. I2959-AA possesses higher chemical reactivity and solubility than I2959 and, therefore, a higher polymerization rate and monomer conversion can be obtained.

3.3.2 2,4,6-TRIMETHYL BENZOYL-DIPHENYL PHOSPHINE OXIDE As photocrosslinking has become an important tool for the encapsulation of living cells in 3D, hydrated, and biomimetic materials, more efficient photoinitiators absorbing longer wavelengths with high initiation rates and good water solubility are required. Commercially available triphenyl 2,4,6-trimethylbenzoyl-diphenylphosphine oxide (TPO) photoinitiator exhibits absorption spectra at longer wavelengths which is suitable for cell encapsulations, but have low water solubility. It absorbs in the region 350
380 nm and 420
440 nm and has high molar

75

Table 3.1 Properties of some Common Photoinitiators used in Biomedical Applications Photoinitiator Properties

I 2959

TPO

LAP

Nano TPO

CQ

Wavelength of absorption (nm)

,370

350
380; 420
440

.365

385
420

Molar extinction coefficient (M21 cm21) Water solubility (wt%)

4365nm ,2

B300
800365nm 0.03

218365nm 8.5

B680365nm 3

200
300; 469 B46469nm Poor (mixed with DMSO/DMF)

3.3 Photoinitiators

extinction coefficients values ranging from B300 to 800 M21 cm21 and is suitable for nonaqueous media only, because of its low water solubility (3.13 mg L21 at 25 C) (Kolczak et al., 1996). TPO generates highly reactive free radicals which react with unsaturated monomers and oligomers with enhanced polymerization rates and thorough curing. TPO also possess good thermal stability and produces no odor or color and is well-suited for commercially available 3D printers used in stereolithography. To improve water solubility, a lithium derivative of TPO, namely lithium phenyl-2,4,6-trimethylbenzoylphosphinate (LAP), has been synthesized and evaluated for photopolymerization efficiency and cell viability during encapsulation (Fairbanks et al., 2009a). This is also a type I photoinitiator (Fig. 3.6) and possesses superior cytocompatibility and water solubility than I2959 and can be used at higher wavelengths (Table 3.1). The authors polymerized diacrylated poly (ethylene glycol) monomers and encapsulated human neonatal fibroblasts in the presence of LAP. Excellent results were obtained with improved polymerization kinetics at low initiator concentrations and light intensities relative to I2959.

FIGURE 3.6 Cleavage of type I photoinitiators; I2959, TPO, and LAP into substituent radicals after excitation.

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Quite recently Pawar et al. (2016) developed water-dispersible nanoparticles of 2,4,6-trimethylbenzoyl-diphenylphosphine oxide with a much higher extinction coefficient and water solubility, absorbing in the range 385
420 nm. The authors used spray drying of oil-in-water microemulsions containing TPO at different concentrations to obtain free-flowing, dry powders. As no chemical modification was made, the high molar extinction coefficient of TPO was retained in the water-dispersed nanoparticles. This approach has opened new opportunities in the field of solvent-free, 3D bioprinting utilizing the advantages of photocurable systems.

3.3.3 CAMPHORQUINONE WITH AMINE PHOTOINITIATOR SYSTEM Camphorquinone (1,7,7-trimethylbicyclo [2 2 1] heptane-2,3-dione) is an aliphatic α-diketone and is utilized as an effective type II photoinitiator in the presence of amines as electron or proton donors for visible light photocrosslinking (Kamoun and Menzel, 2010). Its absorption range is 200
300 nm in the UV region and in the visible region it absorbs at 468 nm. Santini et al. (2013) have reported camphorquinone as the most commonly used photoinitiator in dentistry and reviewed different factors affecting the photopolymerization of dental materials. A camphorquinone/N,N-dimethyl aminoethyl methacrylate system has been used to prepare polydimethacrylate resins which are used in dental composites (Sideridou et al., 2004). CQ is also known as a blue light photoinitiator with a quantum yield of 0.07 6 0.01 per absorbed photon (Chena et al., 2007). Fig. 3.7 shows the

FIGURE 3.7 Generation of free radicals in a camphorquinone-amine photoinitiator system (type II).

3.4 Mechanisms of Light Sensitization

mechanism of CQ-amine photoinitiator systems. CQ absorbs light and is activated into a singlet excited state, undergoes intersystem crossing to triplet state and then forms an “exciplex”—an excited state complex with the amine coinitiator. This exciplex undergoes an electron proton transfer reaction to get CQ radical and aminoalkyl radical (Pyszka et al., 2004). Kamoun and Menzel (2012) reported that with the addition of diphenyliodonium chloride, CQ is further regenerated which accelerates the photopolymerization process. CQ is yellowish in color which may impart intense color to polymerized composites when used in higher concentrations. It also possesses poor water solubility, therefore, polar solvents like DMF or DMSO are generally mixed with water to make it soluble. This property limits its use in photocrosslinking hydrogels. Researchers have attempted the modification of CQ to improve its water solubility in order to widen its application spectrum. The authors modified CQ into carboxylated camphorquinone with improved water solubility and photoreactivity. Ketopinic acid was reacted with SeO2 in glacial acetic acid to obtain 7,7-dimethyl-2,3-dioxobicyclo[2 2 1] heptane-1-carboxylic acid (CQCOOH). The modified CQ can be used efficiently to obtain hydrogels with improved mechanical properties and cell viability (Kamoun et al., 2016).

3.4 MECHANISMS OF LIGHT SENSITIZATION The basic components of light-curable compositions are functionalized oligomers, monomers, and photoinitiators. Other additives and fillers may also be incorporated depending on the required properties of the final product. The major component is oligomer, which is used in the range 25%
90% and has a major impact on the performance of the final product. It constitutes the backbone of the cured product and is responsible for many properties such as chemical, scratch and abrasion resistance, mechanical properties, reactivity, gloss, adhesion, and nonyellowing, etc. Oligomers can be selected on the basis of molecular weights, functionality, and their chemical family such as urethane acrylates, epoxy acrylates, acrylic acylates, and other modified acrylates. Monomers are reactive diluents, also known as viscosity control agents, and can be used in the range of 15%
60%. They possess unsaturated groups, generally carbon-carbon double bonds, which are crosslinked through photocuring and affect the photopolymerization rate and film properties such as hardness, adhesion, and flexibility, etc. (Chandra et al., 1992, 1993). Monomers may be volatile and toxic and, hence, the selection of a suitable monomer is very important for getting high performance and nontoxic compositions. Photinitiators have already been discussed in the previous section, they are light-absorbing molecules and generate reactive species to initiate the process of photopolymerization.

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Two types of mechanisms are generally followed in light-mediated polymerizations: (1) free radical; and (2) cationic polymerization (Phillips, 1984; Takimoto, 1993). Table 3.2 compares the two mechanisms of light-induced polymerizations. Photocuring is mainly based on free radical polymerizations in which light initiates the curing process by activating the PI to generate free radicals. These radicals form carbon-centered radicals by adding to the unsaturated moieties present in oligomers or monomers. After initiation, a chain is propagated involving free radical mechanism and then terminated either by radical recombination or disproportionation reaction (Fig. 3.8). These reactions require inert atmosphere to prevent oxygen inhibition; however, with the advent of novel photinitiators, oxygen inhibition can be overcome. Cationic polymerization, also known as dark reaction, involves photogeneration of cations (strong Bronsted or Lewis acids) commonly based on sulphonium or iodonium salts. The photoinitiator is activated by light and propagation further proceeds in the dark through “attacking” the proton on the electron-rich group of the oligomer or monomer such as vinyl ethers or epoxies. Cationic polymerizations continue to propagate until consumption of the oligomer or monomer in the absence of impurities such as water or bases which inhibit the reaction. The main advantage of cationic polymerizations is the absence of oxygen inhibition; however, these systems are very prone to moisture and get poisoned by high humidity. Photoinitiators used in these systems are somewhat toxic and can leave corrosive residues.

Table 3.2 Comparative account of Free Radical and Cationic Photopolymerizations Free Radical Photocuring

Cationic Photocuring

Mechanism

Free radicals are generated

Resins Curing rate Photoinitiators

Acrylated oligomers Very high Efficient PI

Examples of commonly used PI Oxygen inhibition Shelf life

Benzoin ethers and their derivatives, I2959, LPA, etc. Exist Longer

Applications

Wide spectrum

Bronsted/Lewis acids are generated Epoxy, vinyl ether, etc. High Toxic and their residues can be corrosive Arylonium salts, triarylsulphonium salts Free Easily poisoned by humidity Limited

3.5 Polyacrylates for Biopolymer Applications

FIGURE 3.8 Mechanism of free radical photopolymerization.

3.5 POLYACRYLATES FOR BIOPOLYMER APPLICATIONS In Section 3.2.2, we discussed polyacrylate-methacrylate-based photopolymerizable compositions as dental restorative materials and in hydrogels with numerous biomedical applications. Here, we will focus on synthetic or natural biophotopolymers which have been modified with acrylate or methacrylate moieties or any other light-sensitive groups to obtain photocrosslinkable biopolymers. The acrylated polymers retain their basic properties and utilize the advantages of fast UV photopolymerization.

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3.5.1 POLYCAPROLACTONE Polycaprolactone (PCL) is a synthetic biodegradable linear aliphatic ester which is widely used in several biomedical applications (Bezwada et al., 1995). Photopolymerizable polycaprolactone, that is, PCL acrylate can be obtained by reaction of acryloyl chloride with PCL diol (Kweon et al., 2003). UV-cured networks of PCL acrylate possess higher thermal stability, compressive modulus, and compressive recovery ratios than polycaprolactone because of the acrylation of terminal groups. These PCL networks have found potential applications in tissue engineering for the fabrication of scaffolds. PCL can also be made light-sensitive by introducing methacrylate moieties through methacrylation with 2-isocyanatoethyl methacrylate (Ferreira et al., 2008). The acrylated PCL was photopolymerized in the presence of Irgacure I2959, which is a well-tested photoinitiator for a wide range of cell types and chemical compositions. The membranes were photopolymerized after irradiating for 60 seconds and analyzed for hemocompatibility and thrombogenity and hemolysis values were found in acceptable limits. The developed polymer promoted adhesion between the aminated surfaces, and in vivo experiments were also conducted for investigating adhesion characteristics. Dimethacrylate of polycaprolactone can also be obtained by reacting its diol with methacrylic anhydride (Meseguer-Duenas et al., 2011).

3.5.2 STARCH Starch is the most common carbohydrate made up of glucose units joined by glycosidic bonds containing two components, amylose, a water soluble, (1
4) linked, straight chain polymer and amylopectin, a water-insoluble branched chain polymer of D-glucose units. The ratio of amylose and amylopectin determines the crystalline structure of starch. Vieira et al. (2008) modified starch having pendant hydroxyl groups with 2-isocyanatoethyl methacrylate in order to obtain a photocrosslinkable polymer containing acrylic moieties. UV-cured films of the modified starch were prepared in the presence of I2959 photoinitiator and evaluated for swelling capacity in artificial lachrymal fluid, in vitro biodegradation, and morphology of the cured networks. The authors used modified starch to develop a controlled drug delivery system for ophthalmic application. Timolol maleate and sodium flurbiprofen salts were adsorbed in a photopolymerized polymer matrix and their release profiles were investigated. Starch can also be modified with methacryloyl chloride in the presence of potassium tert-butoxide to obtain starch methacrylate (Cankaya, 2016). Graft copolymerization of starch methacrylate has been performed with other acrylates to obtain crosslinked networks having semiconducting properties.

3.5 Polyacrylates for Biopolymer Applications

3.5.3 DEXTRAN Dextran is a biodegradable bacterial polysaccharide which consists of 1
6 linked D-glucopyranose residues of different lengths and varying degrees of branches with α (1
2), α (1
3), and α (1
4) linkages. Photocrosslinkable dextran hydrogels have been widely investigated for drug delivery systems, release of proteins, and as imaging agents (Kim et al., 1999; Kim and Chu, 2000). The hydroxyl groups of dextran can also be modified with acrylic reagents to obtain acrylic moieties making it suitable for photocrosslinking (Fig. 3.9). A biodegradable and biocompatible drug delivery system has been developed by modifying dextran with 2-isocyanatoethylmethacrylate to obtain a photopolymerizable polymer. Modified dextran was photocrosslinked with N-isopropylacrylamide to obtain

FIGURE 3.9 Photocrosslinkable modified polymers, (A) PCL (synthetic polymer); (B) Starch (natural polymer); and (C) Dextran (natural polymer).

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grafted polymer hydrogels and the release profile of the drug Ondansetron entrapped in the crosslinked network was determined (Almeida et al., 2011). The authors investigated the drug delivery profile of the developed system and determined its surface energy through contact angle measurement. The controlled release of the drug and the tunability of the system to different transition temperatures widens its application range. Dextran can also be acrylated with acryloyl chloride in the presence of pyridine (Omer et al., 2015). UV photopolymerization of synthesized and characterized hydrophilic dextran acrylate and hydrophobic acrylate epoxidized soyabean oil has been conducted to obtain crosslinked, hydrogel networks. The swelling properties and release profiles of these hydrogels can be tailored by adjusting the ratios of dextran acrylate and acrylate epoxidized soyabean oil which balance hydrophilicity and hydrophobicity of the cured polymer networks as well as their crosslink density.

3.5.4 GELATIN Gelatin is a natural biopolymer which can be derived via acid or alkaline hydrolysis of collagen obtained from raw animal materials (Lee and Mooney, 2001). It is an inexpensive, biocompatible, and nontoxic material which can be methacrylated to obtain photocrosslinkable and biodegradable gelatin methacrylate or methacryloyl hydrogels. Gelatin methacrylate can be prepared via the reaction of gelatin with methacrylic anhydride (Fig. 3.10). Gel MA crosslinks, when exposed to light irradiation to form hydrogels with tunable biochemical and mechanical properties. It is a photopolymerizable, cell-responsive, microengineered hydrogel that is useful in wound dressings (Van Den Bulcke et al., 2000), drug delivery systems, and tissue engineering in creating cardiac, vascular, bone, and cartilage, etc., microtissues through photopatterning cell-laden hydrogels (Nichol et al., 2010). It can also be used in dielectropatterning because of its low viscosity and biocompatibility, which is a useful tool for the 3D microscale organization of cells and making functional tissue constructs (Ramon-Azcon et al., 2012). Gel MA-based hydrogels can be microfabricated through photomasking, micromolding, bioprinting, self-assembly and microfluidic techniques to develop constructs with controlled structures and also helpful in biosensing and fundamental cell research (Yue et al., 2015). Shirahama et al. (2016) have optimized the reaction parameters such as initial pH adjustment carbonate-bicarbonate (CB) buffer molarity, gelatin concentration, methacrylic anhydride concentration, reaction time, and reaction temperature for facile, one-pot GelMA synthesis. The degree of methacrylation can be used to modify the mechanical properties of the gel, making it suitable for various applications that demand cell-responsive platforms for creating microengineered tissues and microfluid devices. Its elastic modulus may range between 3.3 and 110 kPa with varying concentrations and degrees of methacrylation (Khademhosseini et al., 2015). Gel MA is thixotropic in nature and can encapsulate a variety of cells, such as human umbilical vein endothelial cells

3.5 Polyacrylates for Biopolymer Applications

FIGURE 3.10 Modification of gelatin, chitosan, and hyaluronic acid with methacrylic anhydride.

(Lewis et al., 2014), endothelial cells, fibroblasts, and chondrocytes (Malda et al., 2013a). It is used extensively in bioprinting through photopolymerization of GelMA using a water-soluble photoinitiator such as Irgacure 2959 or LAP (Bertassoni et al., 2014). On exposure to light, the methacrylate groups form

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covalent bonds and create a solid gel network which can be optimized for tunable mechanical properties through crosslinking time. Vascularized 3D constructs have been fabricated through bioprinting techniques by printing Gel MA with sacrificial pluronic F127 which was viable even after 7 days (Lewis et al., 2014). It is also a very competent matrix bioink owing to its high cell adhesion and biocompatibility which can be further modified by adding hyaluronic acid and carbon nanotubes to obtain improved properties. Extensive work has been conducted on reinforcement of photocrosslinked networks of Gel MA to make them excellent candidates for tissue engineering (Yue et al., 2015). Minerals such as Ca21 binding-carboxyl groups (Zhou et al., 2014) and titanium (Tan et al., 2013), silk fibroin (Xiao et al., 2011), dextran glycidyl methacrylate (Wang et al., 2014), and carbon nanotubes (Ahadian et al., 2014), etc., have been incorporated into Gel MA to enhance their compressive elastic modulus. Recently, Hassanzadeh et al. (2016) codissolved chitin nanofibers and Gel MA in a suitable solvent and exposed to UV light for 3 minutes in the presence of I2559 to get the covalently bonded matrix intertwined with the chitin nanofibers. The elastic modulus of the hybrid was increased 1000-fold and strainto-failure increased by more than 200%, hence, improving the handling of hydrogels for tissue engineering applications. Hybrid hydrogels based on poly(ethylene glycol) dimethacrylate and gelatin methacrylate composite nanostructures have been fabricated through UV-assisted, capillary force lithography (Kim et al., 2014). Hybrid constructs involving the printing of Gel MA with thermoplastics such as poly(propylene fumarate) or polycaprolactone enhance the mechanical strength of a bioprint which is beneficial in cartilage and bone tissue engineering (Wang et al., 2016). The biodegradable composite hydrogels exhibited improved cell attachment and possessed the capability to endothelialize in vivo. Carboxybetaine methacrylate has also been used to reinforce Gel MA to have better mechanical properties, a controlled drug release rate, and good cell viability. The hybrid hydrogels can easily be tuned by varying the ratio of carboxybetaine methacrylate (Lai et al., 2016). As a future prospect, convergence of biological and biofabrication approaches is required in regenerating tissues to promote the progress of clinically relevant applications. The physicochemical tailorable photocrosslinkable Gel MAs with inherent bioactivity are promising hydrogels to be used in tissue repair (Klotz et al., 2016).

3.5.5 CHITOSAN Chitosan is a natural, biodegradable, linear polysaccharide composed of β (1
4) linked D-glucosamine and N-acetyl-D-glucosamine units. It has low cytotoxicity and excellent biocompatibility and is widely used in biomedical hydrogel applications. Chitosan-based hydrogels have been employed extensively in wound healing, drug delivery systems, tissue engineering, and also in brain cancer treatment (Ishihara et al., 2002; Nair et al., 2011; Kim et al., 2010). Chitosan has poor solubility in water and cell culture mediums which limits its application in

3.5 Polyacrylates for Biopolymer Applications

hydrogels. The literature reports the modification of chitosan with photosensitive components to get hydrosoluble and UV crosslinkable chitosan derivatives. Graft copolymerization of chitosan has been performed to obtain natural and synthetic hybrid materials with modified properties (Anbinder et al., 2016). Photoreactive compounds such as 4-azidobenzoic acid (Ono et al., 2000), 2-amino ethyl methacrylate (Valmikinathan et al., 2012), and ethylene glycol diacrylates (Gao et al, 2010; Ma et al., 2011), etc., have been grafted to modify chitosan. A photopolymerizable and antimicrobial chitosan derivative has been prepared through the Michael addition reaction of chitosan with polyethylene glycol diacrylate (Ma et al., 2009). The modified chitosan derivative showed good solubility in water and was able to polymerize under UV radiation in the presence of I2959. Ethylene glycol acrylate methacrylate has also been used to prepare (methacryloyloxy) ethyl carboxyethyl chitosan derivative through Michael addition reaction (Zhou et al., 2011). This water-soluble derivative was photopolymerized using UV radiation to obtain noncytotoxic hydrogels which promote cell proliferation and attachment and are potential materials for tissue engineering scaffolds. In another work, by Qi et al. (2013), methylacroloyl glycine was used to obtain a photopolymerizable chitosan derivative. The authors optimized the photopolymerization of the chitosan derivative with different concentrations of I2959 and found it less thermally stable than chitosan. The swelling property of the hydrogel decreased with an increase in photoinitiator concentration due to the high crosslink density of the cured network. In an another strategy, a hydrosoluble, UV-crosslinkable, and injectable N-methacryloyl derivative of chitosan was prepared through the chemoselective N-acylation reaction of chitosan with methacrylic anhydride (Li et al., 2015a,b). A quick and cost-effective method has been developed for the fabrication of patterned cell-laden hydrogels with rapid transdermal in vivo curing. The authors suggested the use of these polysaccharide microgels with unique amino groups for localized and sustained protein delivery. Polymethyl methacrylate grafted chitosan was prepared recently and investigated for its potential in corneal tissue engineering. The grafted polymer networks are more thermally stable than chitosan and able to degrade in the presence of lysozyme under in vitro conditions (Hemalatha et al., 2016). These grafts supported the proliferation of the human corneal epithelial cell line, suggesting their use in tissue engineering applications.

3.5.6 HYALURONIC ACID Hyaluronic acid (HA) is another linear natural polymer consisting of β-1
4 and β-1
3 linked alternating units of D-glucuronic acid and N-acetyl-D-glucosamine and can be obtained from animal products and the fermentation of bacteria (Van Vlierberghe et al., 2011). It is a hydrophilic and biocompatible polymer that has the ability to interact with numerous cell surface receptors. These properties of HA make it an outstanding candidate for tissue engineering, drug delivery systems, and wound healing applications. HA is a very promising natural

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polysaccharide for the synthesis of photocrosslinkable hydrogels due to the presence of different functional groups such as carboxylic, primary and secondary alcoholic, and N-acetyl groups. Photopolymerized networks are obtained when irradiated with methacrylates and norbornenes in the presence of a suitable photoinitiator (Kesti et al., 2015; Gramlich et al., 2013). Earlier work by Smeds et al. (1999) reports the methacrylation of hyaluronic acid with methacrylic anhydride and photocrosslinking of modified haluronic acid. The authors investigated the bulk mechanical properties of photocrosslinked networks of HA
MA and found excellent swelling capacities of these hydrogels. Compressive modulus was found to increase with the increase in HA
MA concentration; however, the viability of fibroblast cells decreased (Burdick et al., 2005). HA
MA in combination with gelatin methacrylate has been used successfully in the fabrication of tubular constructs through extrusion bioprinting (Skardal et al., 2010). HA
MA and GA
MA were used at a 4:1 ratio, seeded with HepG2 C3A cells, and then exposed to UV radiation for 120 seconds. The bioprinting involved sequential deposition of three layers: a cell-free central blended layer; a cell-containing blended layer; and an outer cell-free HA
MA layer. These constructs were further exposed to UV radiation to strengthen the polymer network. HA
MA and GA
MA blends have also been investigated for printing heart valve conduits encapsulated with human aortic valvular interstitial cells (Duan et al., 2014). After photocrosslinking in the presence of I2959, the obtained constructs showed good structural performance and high cell viability. A bioink involving fast thermal gelation followed by photopolymerization has been developed using modified HA (2%) and nonmodified HA along with a thermoresponsive polymer, poly(N-isopropylacrylamide) (15%). LAP was used as a photoinitiator and the bioink was seeded with bovine chondrocytes (Kesti et al., 2015). The authors fabricated 3D cellular constructs using the biofactory extrusion system with high structural integrity without compromising the viability of the encapsulated cells. Modified HA
MA improves chondrogenesis and the distribution of the extracellular matrix when used in composite hydrogels for cartilage tissue engineering (Levett et al., 2014).

3.6 RECENT ADVANCEMENTS AND TRENDS IN LIGHT-MEDIATED POLYMERIZATIONS Thus far we have discussed polymers involved in light-mediated fabrication of biomaterials utilizing the ultimate advantages of this fast-curing technology. The practice of light-mediated polymerizations in hydrogel synthesis is a recent focus of researchers worldwide. Photoclick reactions are being used to develop photopolymerized hydrogels with improved biocompatibility, photoencapsulation of the cells in hydrogels, and photopatterning techniques within hydrogel matrices to get tailored materials for biomedical applications. The research is being directed to

3.6 Recent Advancements and Trends

exploring new formulations with improved properties, shifting absorption to longer wavelengths, and free of oxygen inhibition to avoid the use of inert atmosphere (Hao et al., 2014). Visible light-mediated, thiol-ene, photopolymerization systems are being developed as an alternative to existing UV-curing systems for studying cell biology. Thiol-alkyne (Fairbanks et al., 2010), azide-alkyne (Adzima et al, 2011), and thiol-norbornene (Fairbanks et al., 2009b) systems are well-investigated for radical initiated photopolymerization reactions. These systems involve the use of type II photoinitiators which absorb in the visible region and are more appropriate for specific applications where UV radiation can be harmful. Thiol-bearing molecules act as coinitiators as well as crosslinkers in the initiation step and the formation of thioether
ester bonds make the hydrogels hydrolytically degradable. Lin et al. (2015) reviewed thiol-norbornene systems and reported their crosslinking through longer wavelength UV or visible lightmediated orthogonal reactions which are free of oxygen inhibition. Photopolymerized thiol-ene hydrogels can be prepared via the reaction of peptides having thiol moieties (cysteines) with acrylates or norbornenes under cytocompatible conditions. Recently, Sawicki and Kloxin (2016) presented a method for developing photopatterning hydrogels from allyloxycarbonyl functionalized peptide crosslinks, pendant peptide moieties, and thiol-functionalized poly(ethylene glycol) through cytocompatible exposure to longer wavelength UV light using lithium acylphosphinate as a photoinitiator. The authors described the procedure for sequential photoencapsulation of the cells and photopatterning of the hydrogel network with biochemical cues. Fast photopolymerization encapsulates suspended cells within the hydrogel matrices and the gel is locked in the desired shape of the mold depending on the application. Methods for encapsulation of human mesenchymal stem cells and the determination of their viability and activity have been established. In another development, Donovan et al. (2016) investigated the photopolymerization of difunctional trifluorovinyl ether monomers with multifunctional thiols to get semifluorinated polymer networks. This reaction involves the anti-Markovnikov addition of thiyl radicals to the trifluorovinyl ether group and a semifluorinated ether-thioether linkage is introduced, which imparts superior mechanical properties such as tunable Tg, high percentage elongation, more tensile strength, and toughness than thiol-ene materials with hydrogenated ether-thioether linkages. The use of biopolymers in light-mediated polymerizations for the fabrication of biomaterials is expanding its horizon due to their biodegradability, noncytotoxicity, and excellent performance in photofabrication techniques. Photofabrication techniques play a crucial role in designing sophisticated 3D constructs for tissue engineering and delivery of bioactive compounds. Stereolithography is a widely used photofabrication technique which involves a computer-controlled, laser, layer-by-layer process to obtain 3D constructs (Rundlett, 2013). It is a free-form technique able to create small features up to 40 μm; however, through photopolymerization induced by two photon absorption, smaller features (120 nm) can be created (Belfield et al., 2007). Stereolithography

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is a biomedically advantageous technique which uses photocrosslinkable materials for the construction of tissue scaffolds and artificial joints (Vanasse et al., 2014). Generally, hydrogels obtained from photocurable materials are used in tissue engineering and the creation of nerve-implanting surfaces with high spatial resolution (Sun et al., 2012; Tuft et al., 2014; Malda et al., 2013b). It works excellently when employed in combination with medical imaging techniques (MRI and CT) and also enhances the quality of complex surgical procedures. With the advent of new biocompatible resins, radiation-grafted polymers and light-sensitive modified natural polymers, cell encapsulation techniques, and stereolithography have become more important in biomedical engineering applications (Pereira and Bartolo, 2015). The practice of photofabrication techniques to engineer synthetic environments which can closely replicate the extracellular matrix to stimulate tissue repair and regeneration as well as the spatiotemporal delivery of therapeutics is the current state-of-the-art in the biomedical field. New catalyst systems are being developed for controlled radical polymerizations of acrylates using visible light and offering the ability of spatiotemporal control in tissue engineering and drug delivery systems. Fac-[Ir(ppy)3], photoredox catalysts have been used efficiently in photopatterning techniques providing a new platform for polymer synthesis (Nicolas et al., 2014). Visible-light photoredox catalysis uses energy stored in the redox potential which enable light-mediated polymerizations to continue in the dark after very short and low-intensity exposure to light (Soto et al., 2014b). Recently, Mousawi et al. (2017) reported the synthesis of azahelicenes as visiblelight photoinitiators for free radical as well as cationic polymerizations using light-emitting diodes. The photoinitiators possess excellent initiating abilities and allow the final conversion of the polymerized networks. Polyurethanes are excellent biomaterials with huge potential in biomedical applications (already detailed in the previous section), but their synthesis involves the use of toxic isocyanates. Current research aims at developing green methods for synthesizing isocyanate-free polyurethanes based on renewable resources such as plant oil and soyabean oil, etc., to mitigate severe toxicity issues and environmental concerns (Javni et al., 2013; Zhang et al., 2016a).

3.7 CONCLUSION Radiation curing is playing a key role in the fabrication of a wide range of smart biomaterials used in dentistry, tissue engineering, and drug delivery systems. The literature review presented in this chapter discloses a range of photocrosslinkable materials with distinct properties, ranging from hydrophobic dental materials to hydrophilic hydrogels. Continuing developments in resins, monomers, photoinitiators, and light sources provide an array of opportunities to tailor even smarter biomaterials. The spatiotemporal control offered by light-mediated polymers has made them in demand for tissue engineering and drug delivery systems. With the

References

growing understanding of complex cellular biological features, radiation-cured, cell-encapsulated hydrogels are able to mimic cellular environments in living tissues. Advances made in photocrosslinkable biodegradable natural polymers have emphasized the use of radiation-curable processes in biomedical applications. Different resins based on polyurethanes, acrylates or methacrylates, vinyl monomers, polycaprolactone, gelatin, starch, dextran, chitosan, and hyaluronic acid have been prepared by making use of 2-hydroxyethyl methacrylate, acrylic acid, methacrylic acid, 2-isocyanatoethylmethacrylate, and methacrylic anhydride, etc. Polyurethane and urethane acrylates were synthesized using diisocyanates with polyols containing pendant hydroxyl groups. They offer high oxygen permeability, flexibility, and tear strength. Several researchers made efforts to improve the thermomechanical properties of polyurethanes. It is pertinent to mention here that polyurethanes have been extensively used in implantable devices, especially for vascular stents, tissue scaffolds, and clot-removal devices. The ability of these materials to offer both hard and soft segments makes them suitable for such applications. Most of the photocurable acrylate systems find applications in dental fillings and other related operations. Efforts have been made to reduce the shrinkage of materials in order to minimize damage to the surrounding tissues. Acrylates and methacrylates have been used in hydrogel synthesis which are potent biomaterials used largely in tissue engineering and drug release systems. Vinyl monomers are generally less cytotoxic than acrylates, but have poor photoreactivity and possess limited applications. N-Vinyl pyrrolidone is a hydrosoluble and bioinert monomer which finds numerous applications in the synthesis of hydrogels. Currently, efforts are being made to replace vinyl monomers by sulphur active vinyl monomers. Efficient photocrosslinkable and biodegradable biopolymers have been developed and employed largely in hydrogels synthesis owing to their noncytotoxicity and biodegradable nature. In the recent past, 3D bioprinting of such polymers has enhanced their practical applications in artificial human organs. The recent trend in photopolymerized systems is the elimination of UV light to using visible light. Development of photoinitiators which catalyze reactions by LED sources is, currently, the most modern practice in this field. To summarize, this chapter presents the systematic growth and upto-date information of the subject matter highlighting the basic components of light-mediated polymerizations.

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107950. Zhou, Y., Ma, G., Shi, S., Yang, D., Nie, J., 2011. Photopolymerized water-soluble chitosan-based hydrogel as potential use in tissue engineering. Int. J. Biol. Macromol. 48 (3), 408
413. Zhou, L., Tan, G., Tan, Y., Wang, H., Liao, J., Ning, C., 2014. Biomimetic mineralization of anionic gelatin hydrogels: effect of degree of methacrylation. RSC Adv. 2014 (4), 21997
22008.

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Thermoset, bioactive, metal
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4

Hari Madhav1, Neetika Singh2 and Gautam Jaiswar3 1

Drug Design and Synthesis Laboratory, Department of Chemistry, Jamia Millia Islamia (A Central University), New Delhi, India 2Materials Research Laboratory, Department of Chemistry, Jamia Millia Islamia (A Central University), New Delhi, India 3Department of Chemistry, Dr. Bhimrao Ambedkar University, Agra, India

4.1 THERMOSETTING POLYMERS 4.1.1 INTRODUCTION Polymers are categorized in various ways, such as by synthesis method, their solvent properties, whether they are synthetic or natural, and their chemical properties, etc. On the basis of their response to temperature, polymers may be divided into two categories, that is, thermoplastic polymers and thermoset or thermosetting polymers. Thermoplastic polymers melt on high temperature and on cooling they again converted into solid form. These polymers can be recycled and easily converted from one form to another form, but thermoset or thermosetting polymers are just the opposite to thermoplastics. They cannot be reshaped or converted into liquid form at high temperature. Some commonly known thermoplastic polymers are PMMA, PE, PVC, PS, and ABS, etc., and the heating and cooling cycle may be repeated several times for thermoplastics without obtaining any chemical changes (Madhav et al., 2017; Rathore et al., 2017; Singh et al., 2018). The identification of thermoset polymers in chemistry is represented in Fig. 4.1. Thermosets are stable against temperature, for instance, they remain unchanged in their shape and physical form, that is, they do not melt even at high temperatures. When subjected to high temperatures they undergo direct degradation and cannot be recycled or reused. These polymers undergo a curing process during heating and shaping, which will cause crosslinking in their molecular structure. In other words, it may be stated that thermosets have a constitutional repeating unit in their structure. These polymers are also referred to as thermosetting resin. The most common known examples of thermosets include phenolic resin, urea
formaldehyde resin, unsaturated polyesters (UPEs), and epoxy resin, etc. These polymers are formed by irreversible liquid or powder to solid Materials for Biomedical Engineering: Thermoset and Thermoplastic Polymers. DOI: https://doi.org/10.1016/B978-0-12-816874-5.00004-9 © 2019 Elsevier Inc. All rights reserved.

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FIGURE 4.1 Schematic representation of the identification of thermoset polymers in chemistry.

transitions which can also be produced by other techniques, such as heating, or UV or electron beam irradiation. This process is known as cure or curing of the material. In this method the primary or mother liquid/powder is converted into polymers through chemical reactions and a highly crosslinking polymeric structure is acquired. This highly crosslinked structure is mainly creditworthy for its high mechanical strength, thermal strength, and poor elasticity or elongation. The thermosetting character of polymers could be explained by the conversion of fluid state of polymers at high temperatures. The molecules which generally allow to flow at high temperature get separated, but in thermosets the highly crosslinked structure present prevents the separation of large molecules at high temperature. Thermosets are network-forming polymers which include highly reactive functional groups, such as epoxy, phenolic, UPE, polyurethane, dicyanate, bismaleimide, acrylate, and many others. Because of these reactions, the materials first show increases in viscosity and then eventually crosslink and become set, and because of this behavior the polymer becomes unable to dissolve or flow. Curing is most often achieved through thermal activation, which gives rise to the word thermoset; this is also observed in the case of light induced network forming materials or by a crosslink dual cure mechanism, for example, thermoset adhesives (Prime, 2009). The primary liquid solution of thermosets may contain various constituents, for instance, it may be a mixture of comonomers which together can react with external actions, such as heating, UV or electron beam irradiation. The constituents of the primary solution may also include nanoparticles, initiators, pigment or

4.1 Thermosetting Polymers

dyes, catalysts, and fibers, etc. The most important and essential condition to produce or for the synthesis of thermosetting polymers is that the monomer(s) present in the primary solution should contain three or more reactive groups in each molecule. These reactive groups create a highly crosslinked three-dimensional structure through irreversible chemical reaction. To convert back into liquid state, it is essential to break the synthesized covalent bonds of crosslinking, but these processes lead to degradation rather than conversion into monomers. Schematic representation of the synthesis of thermoplastics and thermosets are represented in Figs. 4.2 and 4.3. Unlike with thermoplastic polymers, during the processing of thermosets, as shown in Fig. 4.4, as the reaction proceeds, the molecular weight is increased, which begins with the growth and branching of chains, causing an increase in viscosity and a reduction in the total number of molecules. Eventually all the chains link together to form an infinite molecular weight. An important characteristic property of thermosetting polymers is their gel point, which is defined as an irreversible transformation from a viscous liquid to an elastic gel or rubber. The gel point refers to the point at which a primary liquid converts irreversibly into solid form through a curing process. Once the primary liquid crosses gel point, it converts into solid form, which cannot be reshaped, molded, or processed. At the primary stages of the curing process the thermosets can be visualized by an increase in their viscosity η. The gel point meets with the first appearance of an equilibrium (or time-independent) modulus. The polymerization reaction continues beyond the gel point to complete the crosslinked network formation, where the physical properties, that is, modulus, tensile strength, etc., build to levels characteristic of a fully developed network. In a thermoset crosslinked system, the gelation loses its ability to flow and is no longer processable above the gel point, and therefore gelation defines the upper limit of the work life. After curing, the thermoset materials can only mill or grind in micro- or nanoparticles and then be used as a filler for other applications. This allows the thermoset polymers to make components with permanent shapes and sizes. These components

FIGURE 4.2 Schematic representation of the synthesis of thermoplastic polymers in which R represents reactive sites.

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FIGURE 4.3 Schematic representation of the synthesis of thermoset polymers in which R represent reactive sites.

FIGURE 4.4 Schematic representation of thermoset curing. Cure starts with A-stage or uncured monomers and oligomers (A) proceeds via simultaneous linear growth and branching to an increasingly more viscous B-stage material below the gel point (B) continues with the formation of a gelled but incompletely crosslinked network (C) and ends with the fully cured, C-stage thermoset (D) (Prime, 1981).

4.1 Thermosetting Polymers

can be produced by curing of the primary liquid or powder within a mold and these products can then be removed from the mold without allowing time to cool. The first manmade synthetic thermosetting polymer “Bakelite” was patented on the 7th of December 1909 by Belgian chemist Leo Hendrik Baekeland. It was a phenol
formaldehyde resin, which is chemically known as polyoxybenzylmethyleneglycolanhydride, in which the phenol monomer exhibits three reactive sites at o, o0 , and p positions, the second monomer or comonomer formaldehyde exhibits two reactive sites, and the other comonomer hexamethylenetriamine, which was added in second step, shows multifunctional activity. These polymers are used in many applications, for instance, Bakelite is used in the manufacturing of parts for electrical systems due to its high heat resistance and low electrical conductivity, urea
formaldehyde polymers are in wood agglomerates, melamine
formaldehyde polymers are used in laminates, epoxies are used in electronic applications (i.e., in capacitors, transformers, circuit boards, etc.) due to their very low electrical conductivity, UPEs are in glass fiber
reinforced plastic, and polyurethane polymers are used in insulating foams, etc. The properties of these thermosetting polymers are continuously studied by various researchers for improvement and incorporation into some nanofillers, that is, nanoclay, inorganic nanoparticles, carbon nanotubes (CNTs), nanofibers, etc. These properties are also improved by modifications in monomers or polymers through functionalization. Improvements in the properties of thermosetting polymers are focused in the electronic, mechanical, automobile, storage and production, telecommunication, medical, and biomedical fields, etc.

4.1.2 SYNTHESIS OF THERMOSET POLYMERS Plastic polymers are made through the polymerization of monomers, thereby forming macromolecular chains. Besides the monomers, many other chemical substances may be needed during the polymerization process, for example, initiators, catalysts, chain transfer agents, emulsifying agents, and solvents. The chemical processes for chain formation may be divided into chain-growth polymerizations (mainly addition polymerization) and step-growth polymerization (mainly condensation polymerization, but also addition polymerization). In addition polymerization, the monomers are reacted by opening a double bond, but with no molecules being split out. In condensation polymerization, water or other simple molecules, for example, ammonia, carbon dioxide, hydrochloric acid, ethanol, and hydrogen chloride, are split out during the reaction. Vitrification is the transformation from a liquid or rubbery state to a vitreous state. It is a completely distinct phenomenon from gelation, which may or may not occur during curing depending on the cure temperature (Tcure) relative to the glass transition temperature (Tg) for full cure. Vitrification is defined as the point where Tg 5 Tcure, and the formation of glass has been observed due to the Tg increasing from below Tcure to above Tcure as a result of the cure reaction. Vitrification can occur at any stage amid the response to form either an ungelled

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Table 4.1 Glossary of Characteristic Cure Parameters Tgel Tvit Tcure Tg TgN

Time to gelation, gel time Time to vitrification Cure temperature, a process parameter Glass transition temperature, a material property Tg for fully cured thermoset with degree of chemical conversion (α) 5 1

glass or a gelled glass. It may be evaded by restoring at or above TgN, the glass progress temperature for the completely relieved system. In the shiny state, the rate of reaction will more often than not experience a critical lessening and fall underneath the compound response rate as the response ends up being controlled by the dispersion of reactants. Usually completion of vitrification is resulted in an abatement in the rate of response by 2
3 requests of size. Not at all like gelation, vitrification is reversible by warming, and substance control of the fix might be restored by warming to devitrify the somewhat relieved thermoset. Vitrification might be distinguished by a stage increment in warm limit by tweaked temperature DSC (MTDSC) and by powerful mechanical examination (DMA) as a recurrence subordinate change bringing about a glassy modulus typically .1 GPa (Prime, 2009) (Table 4.1). Thermosetting polymers might be formed in two different ways: By polymerizing (step or chain components) monomers where one of them has a usefulness higher than 2 and by synthetically making crosslinks between already shaped direct or stretched macromolecules (crosslinking of essential chains, as vulcanization improves the situation normal elastic).

4.1.2.1 Synthesis of thermosetting polymers by polymerization Step-growth polymerization follows a step-by-step formation of elementary reactions between reactive sites, which are generally functional groups, such as alcohol, acid, and isocyanate. Each independent step forms the disappearance of two coreacting sites and creates a new linking unit between a pair of molecules. To synthesized polymers, the reactants must be at least difunctional; monofunctional reactants interrupt polymer growth. The synthesis of a thermosetting polymer through addition polymerization is represented in Fig. 4.5. Some examples are seen for the synthesis of thermoset through step-growth polymerization. For example, when monomers of amide-co-imide functional benzoxazine are heated at 200 C
215 C, then they are polymerized in a crosslinked thermoset through ring opening polymerization. In this polymerization, the presence of (
NH
CO
) linkages accelerates the polymerization in comparison to ordinary benzoxazines. The intramolecular hydrogen bond between the amide linkage and the adjacent oxazine ring acts as an internal self-complementary initiator (Zhang and Ishida, 2015).

4.1 Thermosetting Polymers

FIGURE 4.5 Linear chain formation and crosslinking via addition polymerization.

In chain-growth polymerization, propagation is followed by the direct reaction of a species bearing a suitably generated active center, such as a free radical, an anion, and a cation. The monomer itself constitutes the reactive solvent and is progressively converted into the polymer. In the polymer growth mechanism, if one of the reactants has a functionality higher than two, then branched molecules of infinite structure will be obtained. The synthesis of thermosetting polymers through condensation polymerization is represented in Fig. 4.6. If glycidyl compound and melamine are mixed together at 60 C in the presence of 4-pyro(idinopyridine) as a catalyst then they will synthesize a thermoset resin N, N-diglycidyl-4-glycidyloxyaniline with melamine. Thereafter, if the temperature is increased to 80 C, it promotes the catalytic process and completely polymerizes the resin in a crosslinked thermoset (Ricciotti et al., 2013). When homopolymer poly(2,6-dimethyl-1,4-phenylene ether) (PPE), prepared from 2,6-dimethyl phenol and a copolymer of 2,6-dimethyl phenol and 2,6-di (3-methyl-2-butenyl)phenol were polymerized by oxidative coupling polymerization using copper (I) chloride-pyridine as a catalyst at 25 C, it was converted into a thermoset polymer with a highly crosslinked structure (Matsumoto et al., 2004). A new concept has also been seen in thermoset polymer science, that is, the term biodegradable or biobased thermoset polymer is used by various researchers and scientists in their work. There are many examples of biodegradable thermoset polymers seen in the literature. For example, if glycerol and citric acid are mixed and heated to between 90 C and 150 C then they are polymerized into a crosslinked polymer through condensation polymerization with water being released as a primary byproduct (Halpern et al., 2014).

4.1.2.2 Synthesis of thermosetting polymers by crosslinking or curing A linear polymer is simply a chain in which all the carbon
carbon bonds are present in a single straight line. A network polymer is synthesized due to the reaction between linear polymer chains or to the build-up of monomeric resinous reactants of a three-dimensional fish-net configuration. This process of interaction

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FIGURE 4.6 Linear chain formation and crosslinking via condensation polymerization.

is called crosslinking and is the main distinguishing feature of a thermosetting material. The term curing is most frequently used for crosslinking or gelling phenomena, while vulcanization is the industrial term for the curing of rubber. Mostly, coreactive monomers are denoted as resin and curing agents. Resin is the resinous monomer from which the family name is derived (e.g., an epoxy plastic is an epoxy resin that has been crosslinked). The term “thermo” implies that crosslinking follows the application of heat energy; most crosslinking occurs at room temperature and below. In the past few years, some researchers have used the term chemosets, in which the reaction of a thermoset takes place at room temperature. The term “setting” references the fact that an irreversible reaction has occurred on a macroscale. The network polymer formed has an “infinite” molecular weight with chemical interconnections that restrict long-chain macromovement or slippage. Molecular functionality may be defined as a number of reactive moieties available in a molecule of reactant which indicates the potential for a crosslinking reaction. A total average functionality between reactant elements greater than two suggests the potential for crosslinking, independent of mechanism. In other words, a linear polymer is formed due to the bifunctional C 5 C bond via an addition reaction, while in a condensation reaction, thermoset structure with a polyfunctional comonomer formed via tri- or polyfunctional reactant (Dodiuk and Goodman, 2014). From Fig. 4.7, a polymerizing mixture of monomers can be tracked by observing the viscosity change versus time at a given temperature. Beginning at t0, the mixture has a viscosity of η0. The heat released from the exothermic reaction decreases the viscosity (η1). As the molecular weight of the mass increases, the resultant mixed viscosity increase outpaces and quickly surpasses any reduction caused by heat. Molecular growth continues over time until a perceptible macroscopic gel-like “lump” can be sensed. This is the gel point (tgel), or more commonly, the gel time. From this point forward, the viscosity goes to infinity and the polymeric mass becomes a macroscopic plastic solid (Dodiuk and Goodman, 2014).

4.1 Thermosetting Polymers

FIGURE 4.7 Schematic representation of the change in viscosity of thermoset prepolymer during curing.

There are many examples for the synthesis of thermoset or thermosetting polymers found in the literature, in which many hardeners or curing agents were used to produce crosslinking in thermoset prepolymer resin. For this process, firstly the thermoset monomer is prepared with various functionalities and then cured with various hardeners or curing agents with heat or else it is cured only using heat by mixing two or more appropriate thermoset monomers or compounds. For example, a phenylethynylcarbonyl terminated imide compound thermoset was prepared by curing at 200 C for 3 hours followed by 220 C for 3 hours. This curing reaction involves the disappearance of CRC and formation of CQC bonds, which was confirm by Fourier transform infrared spectroscopy (FTIR). It was seen in the reaction that polyene initially formed which participated in crosslinking (Kimura et al., 2013). If epoxy monomer diglycidyl ether of bisphenol A (DGEBA) is mixed with triethylenetetramine (TETA) as a curing or crosslinking agent at room temperature than it converts into a crosslinked thermoset (Becker et al., 2011). Another example of the synthesis of a thermosetting polymer is when 2,5-bis[(2-oxiranylmethoxy)methyl]-furan (BOF) and 1,4-bis[(2-oxiranylmethoxy)methyl]-benzene (BOB) are mixed with DGEBA using diethyl toluene diamine and 4,40 -methylene biscyclohexanamine (PACM) as amine hardeners then they are converted into thermoset polymers. When BOF reacts with PACM, the epoxy group starts reacting with the amine group and the polymerization with an epoxy ring opening reaction (Hu et al., 2014). If epoxy derivatives of vanillin monomer are melted and mixed with amine hardener isophorone diamine (IPDA) at 100 C
125 C then the monomer converts into a highly crosslinked structure through chemical reaction, as shown in Scheme 4.1 (Fache et al., 2015).

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SCHEME 4.1 Synthesis of crosslinked materials by epoxy/amine reaction (Fache et al., 2015).

Another synthesis of biobased thermoset polymers was observed when biobased monomers were cured with amine hardeners. The first biobased monomers, that is, isosorbidediferulate (IDF), butanediddiferulate (BDF), glyceroltriferulate (GTF), were synthesized with ferulic acid and various biobased diols, such as ethanol, and then functionalized with epoxy groups. When epoxy containing IDF, BDF, and GTF reacted with a diamine curing agent they were converted into having crosslinked structures (Me´nard et al., 2017). The schematic synthetic route for this synthesis is shown in Scheme 4.2. Curing of Bis-N-phenylbenzoxazine derivatives and N-propyl benzoxazine derivatives at 120 C
180 C and 140 C
200 C respectively, produces highly crosslinked thermosets. This synthesis reaction involves oxazine ring opening and gives some phenolic structures with Mannich bridges, which is represented in Scheme 4.3 (Tu¨zu¨n et al., 2016).

4.1.3 PROPERTIES OF THERMOSETTING POLYMERS Thermosetting polymers exhibit a highly crosslinked, fishnet-like chemical structure and due to this characteristic property, they show many other specific properties in comparison to other materials that make them different from others. Some of the properties of thermoset polymers are discussed here:

4.1.3.1 Formulations The formulation of thermoset polymers is highly complex. Their formulation involves curing, crosslinking, and many reactions between monomers or oligomers with other monomers, crosslinking agents, or hardeners, such as ring opening reactions, free radical reactions, and catalytic reactions. Thermosetting polymers show multistage processing, including the preparation of thermoset monomers with different functionalities, the preparation of thermoset resins with

4.1 Thermosetting Polymers

SCHEME 4.2 Syntheses of the different biobased epoxy precursors GTF3EP, BDF2EP, BDF(amide)2EP, and IDF2EP from ferulic acid and various biobased diols (Me´nard et al., 2017).

SCHEME 4.3 Thermal curing of bis-benzoxazine monomers and representative structure of the resulting networks (Tu¨zu¨n et al., 2016).

mixtures of different necessary monomers, oligomers, hardeners, and fillers, and then the curing process for long periods at different temperatures in appropriate environments.

4.1.3.2 Solvent resistant Thermosetting polymers show excellent resistance against solvents due to their chemical structures and high degree of crosslinking. For example, uncrosslinked

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polyacrylamide can be dissolved in water, but crosslinked polyacrylamide only adsorbs water.

4.1.3.3 Melt viscosity When thermoset monomers, prepolymers, or resins are heated they show high viscosity, but during the curing process after the gel point, the viscosity of resin rapidly increases to a high level or infinity and the resin enters a solid state. It has been seen from various research that when thermoset resins are heated, they do not melt and at higher temperatures they degrade through random scission.

4.1.3.4 Mechanical properties Thermoset polymers show fair to good mechanical properties at room temperature. They are much more stable then thermoplastic even at high temperatures. Most thermosets are brittle in nature while some break under external force (Li and Dingemans, 2017). The mechanical strength of thermosets can be improved using nanofillers, such as CNT, nanofibers, clay-like materials, and certain thermoplastics.

4.1.3.5 Fiber impregnation Fiber impregnation in thermosetting is easy compared to with thermoplastic (melt viscosity 100
1000 Pa) because uncured thermoset resins show low viscosity (1
10 Pa) (Karger-Kocsis, 1999).

4.1.3.6 Processing cycle The processing cycle for the synthesis of thermoset polymers is long. Mostly they take long a time for curing and hardening. After the curing process, most thermosets require a postcuring period of 2
5 hours to ensure complete crosslinking. For example, Tan et al. synthesized various phosphorus-containing polybenzoxazine derivatives as thermoset prepolymers by melting them at 160 C and 170 C to remove the air followed by a curing process at 180 C for 4 hours, then 190 C and 200 C for 2 hours each, finally postcuring them at 215 C for 1 hour to prepare a crosslinked thermoset polymer (Tan et al., 2017).

4.1.4 CHARACTERIZATION OF THERMOSET POLYMERS Thermosetting polymers are characterized using various sophisticated analytical instrumental techniques. The physical, chemical, mechanical, and thermal properties of these components, which have a broad effect on the ultimate properties of polymers, like dimensional stability, moisture and solvent resistance, and mechanical strength, were characterized using techniques such as FTIR, ultraviolet light (UV) spectroscopy, nuclear magnetic resonance spectroscopy (NMR), X-ray fluorescence spectroscopy (XRF), thermogravimetric analysis (TGA), dynamic mechanical thermal analysis (DMTA), and differential scanning calorimetry (DSC). (Forrest, 2003).

4.1 Thermosetting Polymers

4.1.4.1 Fourier transform infrared spectroscopy Infrared (IR) is an extensive identification technique, which is used for the identification of functional groups and the nature of chemical bonds, etc., in chemical compounds. Infrared spectra able to identify the presence of functional groups and chemical bonds present in cured or uncured samples. With the help of FTIR, the nature of reactions during curing and of evolved gases during degradation at high temperatures can be identified. For example, when allyl alcohol lactic acid (ALA) oligomer resin is cured in order to synthesize its thermoset, the changes in chemical structure are identified using FTIR spectra. FTIR spectra of ALA resin showed adsorption band at 1640 cm21 due to double bond of allyl alcohol, band at 3428 cm21 due to
OH stretching and 1759 cm21 due to CQO (Bakare et al., 2015). For DGEBA and epoxidized hemp oil (EHO) cured with citric acid (CA) and tartaric acid (TA) as the hardeners and triethylbenzylammonium (TEBAC) as a catalyst, the involved chemical process was studied with FTIR and is shown in Fig. 4.8. The authors explained that the presence of bands at 3409 cm21 of
OH occurs due to the opening of epoxy rings under the action of carboxyl groups. The band at 915 cm21 of the epoxy group completely disappears from DGEBA due to the same reaction. The presence of bands in cured EHO at 2800 and 3000 cm21 confirms the asymmetric and symmetric stretching of CH, CH2, and CH3 of the EHO structure. The presence of a CQC stretching band at 1609 cm21 shows the presence of an aromatic ring in cured epoxy resin of DGEBA. An absorption band at 1738 cm21 specifies the ester group of EHO in the cured

FIGURE 4.8 IR spectra of: (A) DGEBA; (B) crosslinked DGEBA/EHO/TA; (C) EHO (Mustata et al., 2016).

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thermoset (Mustata et al., 2016). So, based on the spectra of these cured and uncured resins, FTIR spectroscopy is confirmed as having an important role in the study of chemical reactions and structure of cured and uncured thermosetting resins.

4.1.4.2 Nuclear magnetic resonance spectroscopy NMR is the most sensitive and powerful tool for the determination of polymer structures and the presence of functional groups in polymer chains. This technique indicates the types of protons in cured and uncured thermoset resins. It is also used to determine the functionalization and position of functional groups, and it is possible to use solid state NMR for thermoset products. For example, when hemp oil is epoxidized with peroxyacetic acid, NMR has been proven as the best technique for the confirmation of the epoxidation of hemp oil. Proton NMR (1H NMR) spectrum of epoxidized hemp oil shows the appearance of epoxy protons, located peaks at 2.860 and 3.031 ppm, which was not observed in unepoxidized hemp oil (Mustata et al., 2016). The 1H NMR spectrum of epoxidized hemp oil is shown in Fig. 4.9. Another interesting example is the preparation of a renewable isosorbidebased monomer for the preparation of the corresponding thermoset. When the isosorbide is modified using 5-norbornene-2-yl(ethyl) chlorodimethylsilane and 5-norbornene-2,3-dicarboxylic anhydride then norbornyl functionalized isosorbide

FIGURE 4.9 1

H NMR spectrum of EHO (Mustata et al., 2016).

4.1 Thermosetting Polymers

SCHEME 4.4 Synthetic route of norbornenyl-functionalized isosorbide ISN and IN (Wang et al., 2016a,b).

is produced, known as ISN and IN respectively. The structure of functionalized ISN and IN can be determined by 1H NMR. The synthetic route for the synthesis of ISN and IN is shown in Scheme 4.4, while the 1H NMR spectra for ISN is represented in Fig. 4.10. NMR pattern of ISN shown two resonances at δ value of 6.11
6.09 ppm and 5.90
5.87 ppm of two different protons aCHQCHa of Norborne moiety, while peaks located at δ value of 4.45
3.44 ppm was correspond to eight protons of two fused rings of isosorbide segment. The NMR peaks at δ value of 2.78
2.73 and 1.92
1.78 ppm were caused by end protons (CH) and middle protons (CH2) of
CH
CH2
CH
of Norborne moiety. In addition, observed multiples at around 1.11
1.01, 0.67
0.50 and 0.15
0.05 ppm were observed for
CH2
CH2
Si(CH3)2
protons. These NMR signals confirm the proposed structure of ISN (Wang et al., 2016a,b).

4.1.4.3 Differential scanning colorimetry DSC is a useful tool for thermal analyses of thermosetting plastics using changes in heat capacity results due to exothermic or endothermic reactions. DSC can be used to identify the glass transition of thermosets. It can also be used in thermal stability studies of thermoset products and those that investigate the effectiveness of antidegradants and fire retardants. Specifically, in thermosets, DSC is used to determine the cure behavior of thermoset monomers or oligomers. The use of DSC for determining the curing behavior of maleimidobenzoxazine monomer 1-(3-phenyl-3,4-dihydro-2H-benzo[e][1,3]oxazin-6-yl)maleimide (Mal-Bz) showed that exothermic transition begins from 167 C with a maximum 214 C. The total amount of exotherm was 75 cal g21. This exothermic transition corresponds to the polymerization of benzoxazine via the ring opening of oxazine rings and the addition homopolymerization of maleimide (Agag and Takeichi, 2006). DSC is

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FIGURE 4.10 1

H NMR spectrum of biobased monomer ISN (Wang et al., 2016a,b).

also used to measure the Tg of thermoset polymers. The effect of hardeners on the thermal properties of cured materials can also be examined through DSC. The effect of the epoxy/amine ratio on the Tg of cured epoxy monomer diglycidyl ether of methoxyhydroquinone is represented in Fig. 4.11, (Fache et al., 2015). The authors reported that the maximum Tg was observed for a 2:1 epoxy/ amine ratio. This can be helpful in the synthesis of appropriate flame-retardant polymers with appropriate properties. The nature of the reactions in the curing process can also be determine with the help of DSC. The DSC of DGEBA/EHO at ratio 70/30 (w/w) in presence of TEBAC as a catalyst, using CA and TA as hardeners shows single peak. This shows that the reactions during curing, that is, opening of epoxy ring, formation of ester bonds, and reaction between hydroxyl groups, epoxy groups, and carboxy groups, are completed simultaneously (Mustata et al., 2016).

4.1.4.4 Thermogravimetric analysis This technique is used for measuring mass changes as a function of temperature or time. This analysis technique is used in research and development toward determining the thermal stability of various substances and engineering materials that are both solid and liquid. This technique has been used in the quality control and assurance of raw materials and incoming goods, as well as in the failure analysis of finished parts, especially in the polymer processing industry (Forrest, 2003). The thermal stability and thermal degradation of thermoset polymers is determined using this technique. TGA cured neat DGEBA with isophorone

4.1 Thermosetting Polymers

160

140

Tg = 132ºC Tg = 125ºC

120

Tg = 113ºC

Temperature (ºC)

Tg = 102ºC Tg = 97ºC

100

80

Tg = 88ºC Tg = 78ºC

Tg = 71ºC

60

40

20

2.0/0.6

2.0/0.8

2.0/1.0

2.0/1.2 2.0/1.4 Epoxy/amine ratio

2.0/1.6

2.0/1.8

2.0/2.0

FIGURE 4.11 Tg of diglycidyl ether of methoxyhydroquinone-based materials as a function of the epoxy/ amine ratio used (Fache et al., 2015).

Table 4.2 TGA Data for DGEBA/IPDA and P3SP (Me´nard et al., 2015) Thermosets

10% Weight Loss at Temperature ( C)

% of Char Residue Remaining at 700 C

DGEBA/IPDA 1% P(P3SP) 2% P(P3SP) 3% P(P3SP)

364 352 337 325

9 13 16.1 19.3

diamine (IPDA) as a hardener showed a 10% weight loss from its initial mass at 364 C and the char residue was 9% of its initial mass at 700 C. When trithiophosphonate phloroglucinol (P3SP) was added in ratio of 1%, 2%, and 3% (w/w) to DGEBA/IPDA, the temperature for 10% weight loss of these combinations shifted to lower temperature which is represented in Table 4.2 and TGA curves are shown in Fig. 4.12 (Me´nard et al., 2015). TGA of cured ALA oligomer functionalized with methacrylic anhydride (MLA) showed that it loses 10% of its initial mass at about 290 C and 50% of its weight at about 400 C, while cured ALA oligomer pentaerythritol functionalized with methacrylic anhydride (PMLA) looses 10% of its weight at about 302 C and 50% of its weight at about 444 C. TGA curves for MLA and PMLA are represented in Fig. 4.13 (Bakare et al., 2015).

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polymer composites

FIGURE 4.12 Thermograms of P3SP-containing thermosets (Me´nard et al., 2015).

FIGURE 4.13 TGA curve of MLA, PMLA, GLA, and UPE resin cured (Bakare et al., 2015).

4.1.4.5 Dynamic mechanical thermal analysis DMTA or DMA is a sophisticated analytical instrumental technique which is used to record the viscoelastic properties of polymers, that is, it records the storage modulus and tan δ versus temperature for polymeric materials. The effects on the

4.1 Thermosetting Polymers

viscoelastic properties of thermoset materials over a wide temperature range 2150 C to 1200 C can be observed through DMTA. This analysis technique investigates the effect of temperature and loading on the mechanical properties of thermoset materials. For example, DMTA analysis of neat DGEBA and DGEBA/ PEI-PCLX-B formulation shows that the storage modulus and tan δ vs temperature of neat DGEBA was higher and for PEI-PCL10-B formulation both storage modulus and tan δ vs temperature were shifted to lower temperature which indicated that glass transition (Tg) was slightly decreased (Acebo et al., 2014). This shows that the tan δ curves become broader due to the heterogeneity of the structure at the molecular level. The DMTA curves for neat DGEBA and DGEBA/ PEI-PCLX-B formulations are shown in Fig. 4.14. Another example for DMTA in the literature is observed for carbazole containing a DGEBA/TTMP mixture (Korychenska et al., 2016). DMTA curves show a unimodal relaxation, typical of homogeneous materials, but as the concentration of carbazole moieties increases, the tan δ curves shift to higher temperatures and become broad, which indicates lower homogeneity. Tan δ curves of DMTA are shown in Fig. 4.15 (Korychenska et al., 2016). The loading of other polymers is affected by the viscoelastic properties of thermoset polymers, for instance, when triblock copolymer PVPy-b-PVK-b-PVPy was loaded into epoxy polymer DGEBA, it affects the storage modulus and tan δ value (Xiang et al., 2016). DMTA of neat DGEBA shows α-transition at 159 C, which was attributed to the glass
rubber transition of the thermoset. Upon loading of PVPy-b-PVK-b-PVPy at different weight percentages to DGEBA, the α-transitions shifted to a lower temperature as loadings of triblocked copolymer

FIGURE 4.14 Storage modulus and tan δ curves against temperature for neat DGEBA.

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FIGURE 4.15 Tan δ versus temperature at 1 Hz for obtained thermosets with different contents of carbazole moieties (5
30 wt.%) (Korychenska et al., 2016).

increased, but the tan δ is increased for 20 wt.% loading of triblock copolymer. This transition was responsible for the PVK nanophases (Xiang et al., 2016). DMTA curves for the neat epoxy thermoset and loaded triblock copolymers are shown in Fig. 4.16.

4.1.4.6 X-ray fluorescence spectroscopy XRF is a useful technique, which is used for obtaining semiquantitative elemental data from thermoset products and thin ashes. This analysis technique helps to identify inorganic fillers and pigments in samples. Usually, this technique is used in conjugation with IR.

4.1.5 APPLICATIONS OF THERMOSET POLYMERS Thermoset materials are those materials that are made up of polymers jointed together by chemical bonds, thereby acquiring a highly crosslinked polymer structure. These polymers cannot be melted or dissolved and possess excellent thermal stability and rigidity. Thermoset materials are directly responsible for high mechanical and physical strength as compared to thermoplastic or elastomer materials. Some thermosets have good mechanical properties and show high ionic conductivity, which can be synthesized from ionic liquid epoxy monomers. Some examples of thermoset polymers include: urea
formaldehyde resins, melamine
formaldehyde resins, polyurethanes, epoxy resins, phenol
formaldehyde resins, UPEs resins, polyelectrolytes, etc. The uses of thermosetting polymers are broad and are explained here:

4.1 Thermosetting Polymers

FIGURE 4.16 DMA curves of the control epoxy and nanostructured epoxy thermosets containing PVPy-b-PVK-b-PVPy triblock copolymer (Xiang et al., 2016).

4.1.5.1 Urea
formaldehyde resin Several examples of urea
formaldehyde uses include in textiles, paper, foundry, sand molds, wrinkle resistant fabrics, cotton blends, rayon, corduroy, etc. It is also widely used as an adhesive to glue wood together. Urea
formaldehyde is mostly used in electrical appliances, casings, and desk lamps, etc. Urea
formaldehyde is also used in the agricultural field as a source of nitrogen fertilizer, which decomposes into CO2 and NH3. This is performed by the action of microbes, which are found naturally in soils (Nuryawan et al., 2017).

4.1.5.2 Melamine
formaldehyde resins This resin is similar to urea
formaldehyde resin, but melamine resins are more moisture-resistant, harder, and stronger. Melamine
formaldehyde (MF) resins are widely used in laminate flooring, countertops, cabinetry, surface coatings, textile

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finishes, paper processing, impact resistant crockery (e.g., for hospitals and picnics), toilet seats, pan handles and knobs, stain and cut resistant decorative laminates, etc. (Doudiuk and Goodman, 2013).

4.1.5.3 Phenol
formaldehyde resin Phenol
formaldehyde thermosets are used in different industries. It is mainly used in circuit board production for making circuit boards, like PCB, etc. Phenolic resins are used in electrical equipment, for example, caps, handles, buttons, radio cabinets, furniture, knobs, vacuum cleaners, cameras, ashtrays, and engine ignition equipment. It is also used in laminated material, like laminated sheets, rods, and tubes which are made in a great variety, that is, from fabric, paper, wood veneers, etc. (Doudiuk and Goodman, 2013).

4.1.5.4 Polyelectrolytes Polyelectrolytes show many applications in fields, such as in water treatment as flocculation agents, in ceramic slurries as dispersant agents, and in concrete mixtures as super-plasticizers. Furthermore, many shampoos, soaps, and cosmetics contain polyelectrolytes. Certain polyelectrolytes are also added to food products, for example, as food coatings and release agents. Some examples of polyelectrolytes are pectin (polygalacturonic acid), alginates (alginic acid), and carboxymethyl cellulose, of which the last one is of natural origin. Polyelectrolytes are water soluble, but when crosslinking is created in polyelectrolytes they are not dissolved in water. Crosslinked polyelectrolytes swell in water and work as water absorbers and are known as hydrogels or superabsorbent polymers when slightly crosslinked. Superabsorbers can absorb water up to 500 times their weight and 30
60 times their own volume (Bolto and Gregory, 2007; Dobrynin and Rubinstein, 2005).

4.1.5.5 Polyurethane Polyurethanes have many different applications. In modern times, with advances in the different techniques for producing this polymer, manufactures are able to make a wide range of polyurethane apparel, including manmade skin and leathers, which are used for garments, sports clothes, and a variety of accessories. Polyurethanes are most commonly used in major appliances, such as in rigid foams for refrigerator and freezer thermal insulation systems. In addition to the foam that makes car seats comfortable, bumpers, interior “headline” ceiling sections, car bodies, spoilers, doors, and windows, all use polyurethanes. Polyurethane is also used as a household material which includes floors, flexible foam padding cushions. Polyurethanes play a major role in modern material science, such as in composite woods. Polyurethane-based binders are used in composite wood products to permanently glue organic materials to oriented strand board, medium-density fiberboard, long-strand lumber, laminated-veneer lumber, and even strawboard and particleboard, etc. (Guo, 2012).

4.2 Thermoset Metal
Polymer Composites

4.1.5.6 Epoxy resins Epoxy is a wonderful chemical having a wide variety of applications due to its unique chemical and physical properties. It is broadly used in industrial fields, for instance, epoxy resins are used in the paint industry, as structural or engineering adhesive, in the construction of aircrafts, automobiles, and boats, as a coating, encapsulates, casting materials, potting compounds, and binders. Some of their most interesting applications are found in the aerospace and recreational industries (Lagunas et al., 2014; Acebo et al., 2014).

4.1.5.7 Unsaturated polyester resin UPEs, which are usually strengthened by fiberglass or ground minerals, are used in the manufacturing of structural components, such as boat hulls, pipes, and countertops. The principal products include boat hulls, appliances, business machines, automobile parts, automobile-body patching compounds, bathtubs and shower stalls, flooring, translucent paneling, storage tanks, corrosion-resistant ducting, and building components. These UPEs are also used as a filler with ground limestone or other minerals and cast into kitchen countertops and bathroom vanities. Bowling balls are made from UPEs by casting into molds with no reinforcement (Mark, 1999).

4.2 THERMOSET METAL
POLYMER COMPOSITES 4.2.1 INTRODUCTION Currently, the development of polymer nanocomposites with adequate properties is one of the most active research areas in polymer science. Nanoparticles show various properties focused particularly on strengthening electrical conduction and barrier properties to temperature, mechanical properties, and the possible improvement in fire behavior on gases and liquids (Marquis et al., 2011). The polymer composites are combinations or compositions which comprise of two or more materials as separate phases at least one of which is a polymer. Combining polymers to other materials, like as glass, carbon, metal nanoparticles, ceramics, or another polymer, is generally to obtain unique combinations. Glass, carbon, or polymer fiber-reinforced thermoplastic or thermosetting resins, polymer blends, silica or mica reinforced resins, and polymer-bonded or impregnated concrete or wood, etc., are typical examples of synthetic polymeric composites. It is also often useful to consider materials such as coatings (pigment-binder combinations) and crystalline polymers (crystallites in a polymer matrix) as composites. Typical naturally occurring composites include wood (cellulosic fibers bonded with lignin) and bone (minerals bonded with collagen) (Schwartz, 2002). The addition of fillers and reinforcements has played an important role in the polymer industry. Many different types of fillers have been introduced into

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polymers, which provide a synergistic improvement to their processability and the properties of the final products, like tensile strength, heat distortion temperatures, thermal and electrical conductivity, and enhanced gas barrier properties. It has also been confirmed that the addition of high fractions of micron sized fillers results in considerable changes in rheological properties. Examples of these fillers are small solid particles of black carbon, calcium carbonate, glass-fibers, and talc, and their particle sizes usually range within the micron-level (Gupta and Bhattacharya, 2008). Nanoplate fillers may be natural or synthetic clays, as well as phosphates of transition metals. The most broadly used reinforcement is clay due to its natural abundance and its high form factor. Clay added nanocomposites exhibit an overall improvement in physical performance (Marquis et al., 2011). Thermoset polymers are mostly used in polymer nanocomposites, including phenolic resin, epoxy resins, and UPE resins, etc. UPE resins show relatively poor mechanical and thermal properties, which moderates their use in advanced composite systems (Mark, 1999). Thermoset composite materials show many important properties, such as high strength, UV resistance, light weight, nonconductance, corrosion resistance, electrical and exceptional thermal properties. There is increased interest in polymer/metal composites due to their properties, such as multifunctionality, ease in processing, potential in large-scale fabrication, and being lighter as compared to metals. In these composites, when the metal nanoparticles are embedded in the polymer matrix, they show the mechanical, electrical, and chemical properties of metal and polymers. The combination of metal or metalloid atoms and polymers, which can also exhibit newly synthesized products, properties such as mechanical, physical, thermal, electrical, and aesthetic., come under the category of polymer/metal composites (Bhattacharya, 1986). Polymer/metal nanocomposites can be used in several areas, such as catalysis, sensors, electronics, optics, medicine, and biotechnology. Many researches are on-going on novel metals, such as gold, silver, copper, and platinum. These noble metals play an important role in stable dispersion and applicability. There are two different type of metal nanoparticles used for the synthesis of polymer nanocomposites, that is, monometallic, which are formed from single metal elements and bimetallic, which are composed of an allowed or core
shell structure that differs from the two metals (Yadav and Gautam, 2017; Singh et al., 2016). Polymer composites with fillers are of great interest for many fields of engineering. As already mentioned, polymer metals show intrinsic advantages, including being low cost and easily processed among others. On the other hand, metallic materials also show useful properties and characteristics, such as high conductance, mechanical strength, and thermal conductivity. (Delmonte, 1990).

4.2.2 SYNTHESIS OF THERMOSET COMPOSITES Thermoplastic polymer nanocomposites are synthesized using various synthesis methods, but thermoset polymers have certain limitations, that is, thermosetting

4.2 Thermoset Metal
Polymer Composites

polymers cannot be melted, remolded, or reshaped after being subjected to a curing process so there are limitations associated with thermosetting plastic and the synthesis of their composites. Thermoset polymer/metal nanocomposites can be prepared using two different techniques, that is, ex situ and in situ techniques. In the ex situ method, the polymerization of monomers and the formation of metal nanoparticles occur separately and then they are mechanically mixed to form nanocomposites. In this method metal nanoparticles show wide distribution and exhibit poor dispersion in the polymer matrix. This technique cannot be applied to all types of thermoset polymers. In the in situ method, metal particles are generated inside a polymer matrix by decomposition (e.g., thermolysis, photolysis, radiolysis, etc.) or a metallic precursor is dissolved into a polymer by chemical reduction. A commonly applied in situ method is a dispersion process in which the solutions of the metal precursor and the protective polymer are combined, and the reduction is subsequently performed in solution. This method is more effective with a lower cost for the performance improvements of polymers as compared to the ex situ method. These approaches are used in many types thermoset polymers (Mittal, 2013). It has been seen from various researches and literature that thermoset polymer composites are generally prepared by mixing fillers into thermoset resins followed by a curing with hardeners or crosslinking agents and catalysis. In this method, fillers such as metal nanoparticles, glass fibers, silica particles, nanoclay, etc., are mixed through mechanical stirring in uncured resin vigorously to obtained a homogeneous mixture of fillers with resins. This mixture is cured and molded in various molds to obtain the desire shape and size, etc. An epoxy nanocomposite with Al2O3 nanoparticles and CaSiO3 microparticles was prepared through the mixing of these nanoparticles and microparticles in an uncured epoxy resin (Wetzel et al., 2003). After the incorporation of these nanofillers and microparticles into the resin, this mixture was mechanically dispersed to distribute the components homogeneously within the matrix. This mixture was cured with cycloaliphatic polyamine as a curing agent at 70 C
120 C for several hours. Organic
inorganic hybrid nanocomposites of epoxy were synthesized by blending DGEBA with nanostructured polyhedral octa aminopropyl silsesquioxane (POSS-NH2) and cured with diamino dimethyl sulfone (DSS) as a curing agent at 100 C for 25 minutes and poured into a mold (Zhang et al., 2007). Another interesting example of the synthesis of a thermoset composite is when Gd2O3 nanoparticles were mixed with DGEBA epoxy resin via mechanical mixing at 100 C for 1.5 h to disperse these nanoparticles homogeneously. This solution was cured with polyoxyalkylene amine as a curing agent at 120 C (Ma et al., 2011). It has been seen from the literature that nanoparticles also play the important role of a curing agent or hardener, for example, when DGEBA was mixed with methyl isobutyl ketone containing 30 wt.% silica nanoparticles (MIBK-ST), then MIBK was removed under reduced pressure and a viscous solution was obtained. This viscous solution was heated at 170 C to produce a thermal curing reaction. In this curing process, silica nanoparticles played the role of a curing

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agent (Liu and Li, 2005). A composite thermoset polymer was also prepared with SiC nanoparticles with EPOBOND epoxy resin through mechanical stirring and curing with polyamino as an amine hardener in a ratio of 50:25 (Nassar and Nassar, 2013). Rubber microparticles and silica nanoparticles were also used in the synthesis of thermoset polymers and fiber composites. In this method, the epoxy resin DGEBA was blended with silica nanoparticles containing 40 wt.% and carboxyl-terminated butadiene acrylonitrile (CTBN) rubber. This mixture was cured with the methylhexahydro phthalic acid anhydride as a curing agent to obtained the thermoset composite (Hsieh et al., 2010). The synthesis of a threephase polymer
ceramic
metal composite was also marked in polymer science, in which the polymer used was epoxy resin, the metal was silver, and the ceramic was Ca[(Li1/3Nb2/3)0.8Ti0.2]O32δ (CLNT) (George and Sebastian, 2009). For the synthesis of the epoxy-CLNT-Ag composite, the fine powder of sintered CLNT and Ag were mixed with uncured epoxy resin and hardener. This mixture was mechanically mixed for 30 minutes to uniformly disperse the ceramic powder and Ag in this matrix. This solution was poured in to a cylindrical mold and cured at 70 C for 2 hours. A thermoset polymer embedded fumed silica was prepared by adding fumed silica into a monomer of bisphenol E cyanate ester (BECy) through mechanical stirring to obtain the suspension solution. This solution was mixed with a catalyst at 2000 rpm for 2 minutes and poured into a silicon rubber mold followed by curing at 60 C for 1 hour (Goertzen and Kessler, 2008). Another thermoset nanocomposite of silica nanoparticles and epoxy thermoset polymer has been seen in the literature, which was prepared from the mixing of silica particles with DGEBA through mechanical stirring at 26,000 rpm. After some time, polyamine was added as a hardener and this solution was cured at 70 C for 2 hours then postcured at 120 C for 2 hours (Bondioli et al., 2005). A novolac-type phenolic/SiO2 hybrid organic
inorganic nanocomposite was also synthesized. For this, two types of materials were used and these materials were prepared by mixing two solutions. These solutions consisted of phenolic resin/THF and tetraethoxysilane/H2O/THF/HCL. Both these solutions were mixed and cured with hexamethylene tetraamine as the curing agent. The final solution was poured into aluminum dishes and aged at room temperature for 2
3 days (Chiang et al., 2003).

4.2.3 PROPERTIES OF THERMOSET POLYMER COMPOSITES When nanoscale sized nanofillers were mixed into thermosetting resins, it affected many properties of those polymers. For instance, nanoclay improved the thermal and mechanical properties, CNTs and graphene oxide (GO) improved the mechanical properties, while silver nanoparticles improved the antibacterial and conductivity properties of thermosetting polymers. Some of the properties of thermoset polymer composites are discussed here. After the loading of nanofillers into a polymer matrix, they affect the mechanical and thermal properties of those polymers. Some nanofillers decrease the

4.2 Thermoset Metal
Polymer Composites

mechanical property, while some increase the mechanical properties. For application in biomedical engineering, in this section, the effect of fillers on the mechanical properties of composites is discussed here:

4.2.3.1 Tensile strength Researchers have reported that when fillers, such as Cu, Al, and Zn were added to polyester resin, a sharp decrease in tensile strength was observed after the incorporation of 10% filler into the polyester composite. In addition, in a Cu containing composite the tensile strength was higher as compared to Al and Zn containing composites, which shows that every filler has different effects. It was reported that, the hardness of a composite increased with increasing filler content up until 20% filler loading and then it decreased with further loading. This may contribute to poor surface contact between the filler and polymer matrix (Mansour et al., 2007). Considering SiC nanoparticles, when the nanoparticles were mixed with epoxy resin at 10
20 wt.% variation, the tensile strength was decreased and further increased when the weight percentage reinforced (Nassar and Nassar, 2013). It has been seen that, tensile strength was enhanced after the incorporation of CuO nanoparticles in a vinyl
ester nanocomposite (Guo et al., 2007). Research shows that the incorporation of TiO2 nanoparticles improves the tensile strength of polymers, for example, SiC and TiO2 were used to prepare a thermoset nanocomposite and when the SiC was blended with the epoxy resin, the prepared nanocomposite showed a significantly lower impact strength as compared to the neat matrix, but after the incorporation of 7.5 vol.% TiO2 nanoparticles into the epoxy resin, the impact strength was remarkably improved, while when a 17.5 vol.% of TiO2 was mixed with epoxy resin, the material showed increases in impact strength (Wetzel et al., 2001).

4.2.3.2 Fracture surface Fillers show great impact on the fracture properties of thermosets, for example, when nanoparticles of SiC and TiO2 were added into an epoxy resin, the SiC containing epoxy composite showed the fracture surface to be much rougher than that of the epoxy/TiO2 nanocomposite. After some time, the particle agglomerates can clearly be obtained within the fracture surface (Wetzel et al., 2001). Another study was done with TiO2/polyester resin. With the incorporation of TiO2 nanoparticles into the polyester resin with a loading of 1, 2, and 3 vol.%, increases in toughness of 57%, 42%, and 41% respectively were observed as compared to virgin polyester. The result obtained for 1 vol.% of TiO2 nanoparticles was the highest value compared to that of neat polyester (Evora and Shukla, 2003).

4.2.3.3 Stress
strain behavior The influence of particulate fillers on the stress
strain behavior of polymers is well known, at least for fillers in the micrometer size range and larger. On the one hand, rigid microfillers commonly increase stiffness, but on the other hand, they may have a detrimental effect on the stress
strain behavior leading to

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breakages. The flexural strength of microparticle-filled composites may also be reduced with rising filler content, especially in cases where the load transfer between the matrix and the particles is insufficient and the interface is weak (Fridrich et al., 2005). For example, the loading of flax fibers in thermoset polyester resin modified the stress
strain properties of that polymer composite. When these fibers were loaded into the polymer matrix, it was reported that the tensile stress at break and the strain at maximum stress were increased from 31 to 304 MPa and from 0.68% to 1.73% respectively compared to neat polymer, but when the filler was replaced with E-glass fibers then the tensile stress at break increased to 695 MPa and the strain at maximum stress to 2.37 % (Hughes et al., 2007).

4.2.3.4 Dynamic mechanical properties Dynamic mechanical tests, over a wide temperature range are highly sensitive to the physical and chemical structure of polymers and composites. They allow for the study of the viscoelastic properties against the temperature of thermosetting polymers. It was observed from various researches that the interaction of some nanoparticles can increase these properties, while the interaction of other nanoparticles can decrease these properties; for example, ZnO nanoparticles loaded with epoxy resin showed a slightly improved storage modulus, while nanoparticles of barium titanate (BT) loaded with epoxy exhibited decreases in storage modulus (Medina et al., 2016; Li and Zheng, 2016).

4.2.3.5 Wear performance In general, the friction and wear properties depend on the whole tribiological system rather than a single material property. After the incorporation of fillers, composites exhibit improved wear rates. For example, TiO2 and SiC were both added into epoxy resin in a uniform distribution of particles, and it was observed that they help to enhance the wear rate (Wetzel et al., 2001). TiO2 possesses good filler characteristics, that is, when TiO2 microparticles are mixed with polyester resin, it improves the sliding wear resistance of the composite (Satapathy et al., 2010).

4.2.4 CHARACTERIZATION OF THERMOSET POLYMER COMPOSITE Many properties of thermoset composites are characterized through different sophisticated analytical techniques, for instance, morphology can be determined through scanning electron microscopy and transmission electron microscopy, while thermal properties or behavior against temperature can be characterized through TGA, different thermal analysis, DSC. Mechanical and viscoelastic properties can be determined through DMTA. The thermoset nanocomposites can be characterized against structural properties by FTIR and NMR, etc. For example, high-resolution transmission electron microscopy analysis confirmed that ZnO nanoparticles synthesized via an arc-discharge method were prismatic and rod

4.2 Thermoset Metal
Polymer Composites

shaped with average particle sizes of 56 nm (Medina et al., 2016). When these nanoparticles were loaded in epoxy resin, it was observed that the Tg of the ZnO/ epoxy composite was slightly increased, which was determined by DSC. The storage modulus and rubber zone of these composites were determined by DMTA, which showed that the ZnO nanoparticles slightly increased these parameters for the epoxy polymer. The interaction between a GO and hyper branched epoxy (HBE) system was determined through IR spectroscopy. FTIR of GO showed that the bands at 1733 cm21 for CQO, 1396 cm21 for O
H bending, 1222 cm21 for C
O
C str., 855 cm21 for C
O
C of the epoxy ring, etc., confirm the presence of functional groups, like epoxy, carbonyl, hydroxyl, etc., in GO. FTIR of pristine HBE showed bands at 3433 cm21 for O
H str., 915 cm21 for C-O str. of the epoxy ring, and 842 cm21 for C
O
C str. of the epoxy ring, but in HBE/GO composites, the shifting of the O-H band to a lower wave number 3419
3418 cm21 confirmed that GO had interacted with HBE via H-bonding (Baruah and Karak, 2016). The thermal stability of GO/HBE composites was determine through TGA, which showed that the incorporation of GO nanoparticles in the HBE matrix improved the thermal stability, while the Tg of the nanocomposites was lower than that of pristine HBE, which was determine through DSC. GO nanoparticles also improved the tensile strength, elongation at break percentage, toughness, and adhesive strength of the GO/HBE composite compared to that of pristine HBE. The incorporation of barium titanate (BT) nanoparticles improved the dielectric constant of epoxy thermosets (Li and Zheng, 2016). BT nanoparticles decreased the Tg and thermal degradation of an epoxy polymer, which was determine by DSC and TGA. DMTA of BT/epoxy nanocomposites showed increased BT nanoparticle loadings in the polymer matrix decreased the storage modulus of the composite compared to that of the neat epoxy polymer. Broadband dielectric spectroscopy showed that the dielectric properties of neat epoxy was 5.24 and 4.09 at frequencies of 103 and 106 Hz, but it increased with the loading of BT nanoparticles into the epoxy. For a 14.1 wt.% loading of nanoparticles, the dielectric constant was increased by up to 14.6 at a frequency of 103 Hz.

4.2.5 APPLICATIONS OF THERMOSET POLYMER COMPOSITES Imbedding fillers into thermoset polymers typically affect some changes to composites, such as low electrical conductivity and low fracture toughness. These composites also exhibit typically poor resistance to lightning strikes and crack growth (Ladani et al., 2015). There are many advantages of thermoset polymer matrix composites. Thermoset polymer composites show better economics properties than thermoplastic composites and these composites have high temperature properties, good wetting, and adhesion to reinforcement. They cannot be reshaped and melted as thermosets are stable against temperature and exhibit their properties even at high temperatures, that is, strength, mechanical properties, wear resistance, etc. Due to these characteristic properties, these polymers and polymer

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composites are used in various applications, such as in biomedical engineering, the electronic industry, automobile engineering, aircraft engineering, scaffold production, and in a variety of electronic devices, etc. Thermoset cynate ester (CE) composites typically exhibit high temperature resistance and these composites are found in the aerospace sector. The main applications of CE composites include redoes for fighter aircrafts, missile nose cones, skin covering phase array radar, as well as in missile fins, nozzle flaps, fairings, cowls, and inlet guide vanes in jet engines, gear cases for helicopters, etc., (Mangalgiri, 2005). Polyester composites containing glass reinforced fillers are used in several areas, including automobile-body panels, seats and panels for transit cars, boat hulls, bathroom shower and bathtub structures, chairs, architectural panels, agricultural seed and fertilizer hoppers, tanks, and housings for a variety of consumer and industrial products. Glass fiber
reinforced epoxy composites are used in filament-wound pipes and tanks, and circuit boards, while carbon fiber (CF) based epoxy composites show light weight, high strength and modulus, and also exhibit excellent fatigue properties. Due to these characteristic properties CF/epoxy composites are used in military aircraft aerospace components, fuselage panels for military aircrafts, cargo doors for space shuttles, and high-priced sports equipment, such as tennis-racquet frames, golf-club shafts, skis, and archery bows.

4.3 APPLICATIONS IN BIOMEDICAL ENGINEERING Materials used in biomedical fields should have certain special properties as well as being tunable to meet the need for selected applications, for example, they should be biocompatible, noncarcinogenic, corrosion-resistant, and should have low toxicity (Teoh et al., 2004). The selection of an appropriate material is dependent on the application of that material, for example, polymers used in scaffolds should be biodegradable so that as cells generate their own extracellular matrices and a patient’s own tissue will completely replace the polymeric material with time, but in the case of polymeric heart valves, the polymeric material should be wear-resistant and nonbiodegradable so it may remain stable and not degrade with time. These materials should not disturb or induce the opposite response from the host. These materials should be sterilizable and should not decompose or emit toxic gases during the sterilization process. These materials should meet with all the functional and mechanical requirements in order to be applied in the biomedical field, some such applications are discussed here:

4.3.1 IN DENTISTRY The development of restorative materials for dental problems for use in dentistry is challenging due to the environment of the mouth. The temperature of the

4.3 Applications in Biomedical Engineering

human mouth varies from 32 C to 37 C, moreover the intake of hot or cold food changes the temperature of mouth from anywhere between 0 C and 70 C. Another important factor that affects the environment of mouth is pH, which ranges from 4 to 8.5 and varies from 2 to 11 with the intake of various foods. The success and failure of dental material is dependent on the “selection of appropriate material for a given application” and the “ability to carry out manipulative procedures to arrive at the optimum properties” of those materials. Dental materials can be broadly categorized into ceramics, polymer composites, and metals, etc. Due to this, many researchers have tried to prepare dental material from thermoset polymers, such as Bis-GMA. A dental material was prepared from Bis-GMA and triethylene glycol dimethacrylate (TEGDMA) with a silanized glass filler (Atai and Watts, 2006). For this, 65 wt.% Bis-GMA and 35 wt.% TEGDMA was used in a matrix with 0.5 wt.% camphorquinone and 0.5 wt.% N,N0 -dimethyl aminoethyl methacryate as a light-curing initiator. Silanized glass in 3.4 μm average particle size was also mixed into this matrix to decrease the shrinkage-strain and maximum shrinkagestrain rate. Polyurethane resins are also used in dental applications, but they show high shrinkage. To reduce the shrinkage of these resins, Atai et al. modified urethane di(meth)acrylate (UDMA) with isophorone diisocynate (IPDI) to prepare a dental material, and it was observed that the shrinkage of the synthesized material was low, roughly near to that of Bis-GMA/TEGDMA resin (Atai et al., 2007). Another researcher modified Bis-GMA/TEGDMA resin with liquid rubber as a filler for dental applications to improve fracture toughness (Mante et al., 2010). Marsich et al. modified Bis-GMA/TEGDMA with silver coatings to improve its antibacterial properties for orthopedic and dental applications (Marsich et al., 2013). The coatings of Ag nanoparticles were done with polysachride-1-deoxylactil-1-yl chitosan. Another thermoset resin of Bis-GMA, methyl methacrylate (MMA), N,N-cynomethyl methylaniline (CEMA), and camphorquinone (CQ) with different weight percentages of silanized zirconia fibers was used to prepare dental material (Wang et al., 2016a,b). Polyhedral oligomeric silsesquioxane (POSS) was also used to reduce the shrinkage of Bis-GMA/TEGDMA resin for dental material. POSS decreased the shrinkage from 3.53% to 2.18% and it also increased the mechanical properties of the Bis-GMA/TEGDMA thermoset resin (Wu et al., 2010).

4.3.2 IN PROSTHETIC HEART VALVES Thermoset polymers and polymer nanocomposites have been used in the preparation of prosthetic heart valves and artificial hearts by various researchers and scientists. Thermoset polyolefins, that is, crosslinked poly(styrene-block-isobutylene-block-styrene), was used for the preparation of a polymeric heart valve (PHV) (Claiborne et al., 2013). The use of crosslinked polyolefins were also reported by Zhou et al. for making a PHV (Claiborne et al., 2012). Another

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polymer composites

thermoset, that is, poly(styrene-cobleck-4-vinylbenzocyclobutene)-polyisobutylenepoly(styrene-coblock-4-vinylbenzocyclobutene) was synthesized in order to make an efficient PHV (Sheriff et al., 2015), and it has been seen that various thermoset polyurethanes are also used for making heart valves (Rogers, 2005; Lambert et al., 2001).

4.3.3 IN BONES Bone fractures and the development of materials that meet the required properties of bone are burning issues in bone science. Researchers and scientists are continuously attempting to develop the ideal prosthetic material for bone repair. They use metals, alloys, ceramics, polymers, and polymer composites for repairing bone. Thermoset polymers and their composites along with various fillers play an important role in bone repair, for example, bone morphogenetic proteins transduced in injectable thermoset hydrogels were used to repair segmental defects in rat femurs (Rutherford et al., 2002). Epoxy nanocomposites were also used in bone applications, that is, multiwalled carbon nanotubes (MWCNTs) were loaded into epoxy-based resin to enhance the strength and modulus for making an efficient material for application in the sockets in transfemoral amputees (Arun and Kanagaraj, 2016). The use of CF/flax/epoxy hybrid thermoset composites was marked for orthopedic long bone fracture plates as an alternate to metal plates (Bagheri et al., 2013). These CF/flax/epoxy plates were closer to human cortical bone compared to clinically used metal plates. A thermoset material, CORTOSS, was used as a cortical bone void filler (Pomrink et al., 2003). It is a glass
ceramic reinforced composite that consists of Bis-GMA, 2,2-bis[4-(2-methacryloxyethoxy)]phenylpropane (Bis-EMA), TEGDMA, and 2,6-di-tert-butyl-p-cresoln (BHT). This material was primarily developed as a cortical bone void filler. E-glass fibers reinforced with a Bis-GMA/TEGDMA thermoset composite was used in the formation of intramedullary rods or intramedullary nails, which are used to treat fractures in the long bones of the body (Moritz et al., 2014).

4.3.4 IN BONE GRAFTING Many thermoplastic and thermoset polymers and their composites have been used in bone regeneration by many researchers. A biodegradable polyurethane acrylate/2-hydroxyethyl methacrylate (HEMA) grafted nanodiamond (ND) composite was synthesized for bone regeneration (Alishiri et al., 2016). It was reported that this composite showed high modulus and strength and also did not cause any negative effect on proliferation. Hybridized carbon nanofibers (CNFs) containing calcium phosphate (CaP) nanoparticles performed an important role in improving the interfacial adhesion of epoxy resin for bone repair (Gao et al., 2016). It was reported that these CNF/CaP nanoparticles also enhanced the flexural properties of epoxy composites and that these materials play an important role in bone repair. Another researcher prepared a bioactive thermoset composite from Bis-GMA reinforced with E-glass fibers to replace metallic implants for bone

4.3 Applications in Biomedical Engineering

(Kulkova et al., 2016). The fatigue resistance and mechanical properties of this composite match the properties of bone. Researchers have tried to synthesis a promising material for bone grafting from natural sources. For example, Natarajan et al. synthesized a biodegradable poly(ester amide) from soybean oil for modulated release and bone regeneration (Natarajan et al., 2016). Biodegradable star-shaped polylactide scaffolds were also synthesized for bone tissue regeneration (Timashev et al., 2016). These scaffolds provide a beneficial microenvironment for osteogenic mesenchymal stem cell differentiation in vitro and support de novo bone formation in vivo.

4.3.5 IN PROSTHETIC SOCKETS Polymeric materials were also used to develop prosthetic sockets. An MWCNT reinforced epoxy composite was developed for application in prosthetic sockets in transfemoral amputees (Arun and Kanagaraj, 2016). This synthesized thermoset material increases the comfort level by decreasing the metabolic cost of the socket in transfemoral amputees. Thermosetting polyester resin was used for lamination in prosthetics to match the skin tone of patients (Saikia et al., 2015). Natural fibers are also used to synthesize thermoset polymer composites for prosthetic sockets. A bioactive banana pseudo stem fiber reinforced epoxy composite was synthesized and used as a material to replace transtibial prosthetic sockets (Odusote et al., 2016). Another biodegradable thermoset composite was marked as a material for prosthetic sockets. Pineapple leaf fibers were loaded in different variations into thermoset polyester and epoxy resin (Odusote and Oyewo, 2016). These composites showed some efficient mechanical properties for the development of prosthetic sockets.

4.3.6 IN MEDICAL DEVICES The use of thermoset polymers and their composites have also been reported in the manufacturing of medical devices. Magnetic nickel zinc ferrite particles were loaded with ester-based thermoset polyurethane (Buckley et al., 2016). CNT fabricated polydimethylsiloxane composites were used in the manufacturing of electrodes for electrocardiography (Liu et al., 2015). Thermoset polyurethanes were also studied for application in medical devices. A polyurethane thermoset, MP5510, was investigated and reported as being particularly well suited for medical applications, especially deployment devices, that is, as stents or clot extractors (Baer et al., 2007). There are many electrical devices used in medical field. These devices are prepared using various metals, polymers, ceramics, etc. Many researches have been performed for making suitable polymer compositions for these devices, for instance, graphene/epoxy composites were synthesized for electrical applications due to their adequate mechanical and electrical properties (Wajid et al., 2013).

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Further Reading

Wu, X., Sun, Y., Xie, W., Liu, Y., Song, X., 2010. Development of novel dental nanocomposites reinforced with polyhedral oligomeric silsesquioxane (POSS). Dent. Mater. 26, 456
462. Xiang, Y., Li, L., Zheng, S., 2016. Photophysical and dielectric properties of nanostructured epoxy thermosets containing poly(N-vinylcarbazole) nanophases. Polymer (Guildf). 98, 344
352. Yadav, S., Gautam, J., 2017. Review on Undoped/doped TiO2 nanomaterials: synthesis, Photocatalytic and antimicrobial activity. J. Chin. Chem. Soc. 64, 103
116. Zhang, K., Ishida, H., 2015. Smart synthesis of high-performance thermosets based on ortho-amide
imide functional benzoxazines. Front. Mater. Available from: https:// doi.org/10.3389/fmats.2015.00005. Zhang, Z., Gu, A., Liang, G., Ren, P., Xie, J., Wang, X., 2007. Thermo-oxygen degradation mechanisms of POSS/epoxy nanocomposites. Polym. Degrad. Stab. 92, 1986
1993.

FURTHER READING Kotsilkova, R., 2007. Thermoset Nanocomposites For Engineering Applications. Smithers Rapra Technology Limited, UK.

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Epoxy composites in biomedical engineering

5

Satheesan Bobby and Mohammed Abdul Samad Mechanical Engineering Department, King Fahd University of Petroleum and Minerals, Dhahran, Saudi Arabia

5.1 INTRODUCTION Since their discovery in the late 1930s, epoxy systems in bulk and coating forms have been widely employed in a range of applications spanning several industries. They are typically produced by condensation or a step-growth polymerization reaction between compounds containing epoxide groups and those possessing amino groups. This crosslinking reaction between a base and solidifier, as they are generally known, is exothermic to such an extent that efforts are often directed toward reducing the peak exotherm temperatures that result from the curing process. High-cure temperatures prove to be quite detrimental to the mechanical properties of the epoxy polymer, thereby, reducing its usable life. For almost all practical purposes, the cured epoxy system is identified by the nature of the reactants involved. Although several types of epoxy systems exist, three popular ones are those based on bisphenol-A, bisphenol-F, and novolac and are distinguished primarily by the type of compounds undergoing the reaction. Additionally, the properties of the product are also largely dependent on the type of amine used to complete the curing process. The choice of the type of amine to be used depends on the intended area of application and the operating conditions such as the temperature and process media involved, loads acting on the system under consideration, and the general tribo-environment, among others. Pristine epoxy systems devoid of additives or fillers are known to possess excellent temperature-resistance characteristics combined with exceptional chemical resistance. However, their low load-bearing capacity and inability to retain thermal stability under high levels of sustained loading render them unsuitable for several practical applications involving sliding or impact loads. Over the years, researchers have tried to overcome these issues by reinforcing the epoxy matrix with nano- and micron-sized fillers (Bobby and Samad, 2016). In this chapter, a focused review of the works in the field of composites for biomedical applications has been carried out. An effort has been made to highlight the recent developments in the field of biomedical research involving epoxy bulk and composite coatings. A detailed description of the techniques used by various researchers for Materials for Biomedical Engineering: Thermoset and Thermoplastic Polymers. DOI: https://doi.org/10.1016/B978-0-12-816874-5.00005-0 © 2019 Elsevier Inc. All rights reserved.

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formulating such advanced epoxy composites has also been provided. A point to be noted is that irrespective of the properties obtained by using fillers to reinforce the epoxy matrix, limited works have been carried out by researchers over the past few years owing to the toxicity of the monomers in partially cured epoxy resin. Recent studies have, therefore, focused efforts on developing epoxy systems using biomaterials (Huo et al., 2016). Though biobased feedstock can be used to substitute the bisphenol-A monomer, the thermal properties of the cured system is affected, triggering a need for the incorporation of fillers or additives to improve the desired properties. The flow chart presented in Fig. 5.1 provides the reader with an outline of this chapter.

FIGURE 5.1 Flowchart representing the outline of this chapter.

5.2 Artificial Implants and Bone Fixation Plates

5.2 ARTIFICIAL IMPLANTS AND BONE FIXATION PLATES 5.2.1 ARTIFICIAL IMPLANTS In the world of polymers, modern thermoplastics such as ultra-high molecular weight polyethylene (UHMWPE) and polyetheretherketone (PEEK) have been widely used in both pristine or composite forms to produce artificial implants. One of the major challenges faced during the development of implants, apart from the requirement of biocompatibility, is the need to ensure that wear rates are kept to a minimum. This is due to the fact that the production of wear debris under various contact conditions leads to osteolysis (bone loss) with third body particles increasing the rate of wear, thereby, causing aseptic loosening of the implant and eventual failure of the joint (Dattani, 2007). The use of epoxy-based composites modified with multiwalled carbon nanotubes (MWCNT) for developing sockets in transfemoral amputees has been proposed in a recent study (Arun and Kanagaraj, 2016). A transfemoral prosthetic device serves as an artificial limb replacement where the knee joint has been removed—the individual still has a portion of the femur or thigh bone intact— and consists of a socket, rotator, knee joint, pylon, and foot (John, 1999). Some discomforting statistics were presented by Aruna and Kanagaraj on the percentage of the population in various countries undergoing amputations thereby, further highlighting the need for carrying out research in this field, with a special emphasis on ergonomic factors like socket design and overall comfort. In their work (Arun and Kanagaraj, 2016), the authors used MWCNT of 97 wt.% purity with an outer diameter in the range of 10 20 nm and a length of 5 15 μm. The epoxy composite with a sandwich structure was prepared using a combination of techniques as outlined in Table 5.2. An important finding from the study was the huge increase in mechanical properties, particularly compressive strength and modulus at a filler loading of 0.3 wt.%. The authors cited four reasons for the improvement and a summary of these is: (1) the presence of MWCNT in the composite which possesses excellent inherent properties such as a high modulus of 950 GPa and strength of up to 1 TPa; (2) the interfacial bonding between the reinforcement and the matrix; (3) the increased crystallinity of the composite; and (4) uniform dispersion of the fillers in the matrix. Points (2) and (3) were also cited as the reasons for improved thermal properties. However, it was found that when the filler concentration was increased further, micro voids began to develop in the epoxy matrix due to improper wetting of the reinforcement fillers owing to an increase in specific surface area of the filler particles. Once the optimum epoxy-MWCNT was formulated through iterative trials, characterization of epoxy/MWCNT/Eglass woven fabric or stockinette was carried out. A sandwich composite using pristine epoxy was also fabricated to act as a control specimen. The thermal conductivity and diffusivity of the composites containing 0.3 wt.% MWCNT was found to be greater by B30% and 52%, respectively, compared to composites prepared using pristine epoxy. A practical aspect of this result was defined by the

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authors that such high heat transfer characteristics would improve the comfort level of sockets by decreasing metabolic cost. Furthermore, the improvement in flexural properties by about 11% in composites reinforced with MWCNT and possessing 4 10 layers of stockinette was also pointed out. Osseointegration, in simple terms, can be defined as a process which leads to the formation of a connection, both structural and functional, between a living bone and an implant. Furthermore, an implant is said to be fully osseointegrated when no relative movement occurs between the implant and the bone to which it is in direct contact (Mavrogenis et al., 2009). In a detailed review on problems associated with osseointegration using titanium implants, it was pointed out that two grades of titanium—commercially pure titanium indicated as CPTi (with varying grades of oxygen content from 0.18% to 0.40%) and Ti6Al4V alloy (consisting of 6% aluminum and 4% vanadium)—were workhorses of the implant industry with CPTi being used for dental applications owing to its capability to form a stable titania oxide layer and Ti6Al4V being popular in total hip replacement (THR) and total knee replacement (TKR) surgeries (Petersen, 2014). Despite its many advantages, the review laid out various reasons for the need to consider replacements to titanium, one being the availability of a new and improved alternative—an epoxy/carbon fiber-reinforced composite for the manufacture of artificial implants. The author touched upon several aspects in this regard and are summarized here for informative purposes: (1) corrosion: even though the titanium dioxide layer formed ensures passivity of the metal, the authors explained that this passive layer could be easily broken down owing to harsh acidic environments generated by the organelle mitochondria during periods of lower concentrations. Furthermore, upset conditions such as an inflammatory response triggered by the body could result in less oxygen availability at the surface of the implant resulting in an even lower pH at the implant site leading to the rapid breakdown of the passive layer; (2) infection: another issue stated was the aseptic loosening of the implant resulting from bacterial infection; and (3) coating: the presence of an oxide layer on the surface, which can attain a thickness of approximately 200 nm, is considered to be the primary reason assisting bone growth on the surface of the implant, a point also highlighted in a later work (Raghavendra and Dhinakarsamy, 2015). In the referenced review paper (Petersen, 2014), the author also mentions that a hydroxyapatite (HA) coating is usually sprayed onto the metallic implant to improve the surface roughness (and, thereby, the surface index) of the implant to provide a better substrate for bone growth; however, a combination of an unstable coating film of HA coupled with an adequate surface profile which can safely harbor bacteria leads to tissue damage around the implant—a condition commonly referred to as peri-implantitis in dentistry. A point cited in favor of polymer matrix composites (such as the epoxy/carbon fiber composite) was that these fibers were found to promote bone formation over the implant with a much higher degree of osseointegration when compared to Ti6Al4V. The adapted graphs shown in Fig. 5.2 give a comparison

5.2 Artificial Implants and Bone Fixation Plates

FIGURE 5.2 Percentage bone growth over various substrates.

of the percentage bone area observed at a distance of 0.1 and 0.8 mm from the implant (Petersen, 2014): In addition, the bone structure growth over the epoxy/carbon fiber composite was found to be pore bearing, implying a better degree of oxygen and nutrient accessibility compared to the titanium alloy which lacked such features. Another interesting point stated in the article was that the polarity of cell membrane— composed of lipids, proteins, and carbohydrates—which would come into contact with an implant was more closely matched to that of the epoxy composite as opposed to the metal alloy. To conclude, the authors stated that although there were quite a few advantages when compared to the metallic alloy, much more research was required to be undertaken prior to full-fledged clinical trials. In another study, Hou et al. proposed a novel surface-coating technique to enhance the surface properties of titanium implants (Hou et al., 2016). In the referenced work, epoxy and epoxy/polyester hybrids were prepared by the addition of nano-titania (n-TiO2), micro-titania (μ-TiO2), and calcium oxide (CaO) fillers in various quantities as detailed in Table 5.1. It was expected that n-TiO2 particles would play the role of flow modifier (by acting as miniature spacers between particles) while CaO would act as a functional additive. The functional additive loading was limited to B5 wt.% of the formulations to maintain the integrity and homogeneity of the coating surfaces. μ-TiO2 particles were expected to improve the overall biocompatibility of the system. The coatings were tested for enhancement in biocompatibility and osteoinductivity (the property by which a graft material or implant induces de novo bone growth through biomimetic substances) and comparisons were drawn to specimens of untreated CPTi. Another

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Table 5.1 Formulation of Reinforced Composites Formulation

#1

#2

#3

#4

Epoxy Polyester n-TiO2 CaO μ-TiO2

ü

ü ü ü ü

ü

ü ü ü ü ü

ü ü

ü ü ü

CaO was added to the blend at B5 wt.%; n-TiO2 at B0.5 wt.% and µ-TiO2 at B25 wt.%.

important point stated by the authors was the choice of not using 1,3,5-triglycidyl isocyanurate (TGIC), a low molecular weight multifunctional crosslinker that acts as the curing agent in polyester-based coatings and known to be quite hazardous. Hou et al. resorted to using minute quantities of another curing agent, dicyandiamide, which has been safely used in food over a number of years. However, the researchers found that a great degree of crosslinking was expected to occur between the epoxy and polyester systems eliminating the need for a curing agent in epoxy or polyester hybrids. As seen from Table 5.1, all powdered formulations (1), (2), (3), and (4) were enriched with n-TiO2 and CaO and mixed thoroughly using a high-shear mixer followed by passing through a 35 μm sized mesh. The dimensions of the final particles obtained were verified by using a laser particle analyzer. Finally, the ultrafine-powdered particles were electrostatically sprayed (20 kV) onto circular shaped CPTi disks (grade 2, 0.5 mm thick, 24 mm diameter) using a Corona Gun. The dry powder-coated disks were then cured at 200 C for 10 minutes in a highperformance air flow oven. Testing involved: (1) comparing the capacity of both coated disks and CPTi control surfaces to support cell attachment and spreading in short-term cell in vitro cultures; and (2) comparing the capacity of both coated disks and CPTi control surfaces to support the initiation and progression of biomineralization in long-term cell cultures. All of the ultrafine dry powder coatings supported cell attachment and spreading as confirmed through optical microscopic evaluation. The human mesenchymal cells grown on the surface of coated specimens revealed varying degrees of biomineralization with alizarin red staining tests confirming moderate amounts of mineral deposits on the epoxy/polyester pure hybrid, epoxy/polyester hybrid-TiO2 enriched, and also on the epoxy-TiO2 surfaces, within 2 weeks of growth and differentiation. It was concluded that both epoxy and epoxy/polyester hybrids enriched with TiO2 and CaO would serve as an ideal coating material for titanium implants by enhancing their biocompatibility as well as osteoinductivity.

5.2.2 FIXATION PLATES, SCREWS, AND INTRAMEDULLARY NAILS Bone fixation plates, screws, and intramedullary nails have been widely used in the medical realm to stabilize and support fractured bones. Such internal splints

5.2 Artificial Implants and Bone Fixation Plates

(as they are known) support and stabilize the broken bone until they are capable of supporting the body’s weight and movement. Titanium, stainless steel, and cobalt-chrome alloys have been a preferred choice in manufacturing such supports owing to their resistance to corrosion from body fluids and biocompatibility. However, a common problem associated with these types of in vivo supports is osteoporosis or bone atrophy caused by the “stress-shielding” effect—a phenomenon which results in a decrease in bone density in response to the reduced load placed on it owing to a support carrying this load, resulting in recurrence of fracture over a period of time—caused mainly due to the huge difference in elastic modulus of the internal splint compared to the human cortical bone. Recent research programs in this field have been aimed at developing composites using carbon fiber/flax on an epoxy backbone (Bagheri et al., 2013, 2014a, 2015). In the referenced pioneering work, a composite possessing a “sandwich structure” was developed. In the first of the series of a large number of tests carried out across several years, the authors evaluated the mechanical properties under uniaxial tension and three-point bending, to simulate the loads experienced by orthopedic femur fracture plates. Rockwell hardness tests were also carried out and scanning electron microscopy (SEM) scans were used to characterize the fracture mechanisms. It was postulated that the particular composite in question would be a perfect alternative to long-bone, metallic, fixation plates. The composite was designed such that the epoxy/flax composite would constitute the core with thin sheets of the epoxy/carbon fiber system attached to the outer surfaces of the core. The detailed preparation technique is provided in Table 5.2. It was observed that the flexural properties (flexural strength of 510.6 MPa) of the developed composite were superior to tensile properties (ultimate tensile strength of 399.8 MPa); however, both these values were found to justify the recommendation of this composite as femoral-bone fixation plates due to the fact that clinicaltype loads on the human femur rarely exceed 2.5 3-times the body weight. Both tension and bending tests revealed brittle fracture for all specimens with linear stress strain behavior under tensile loading culminating in an abrupt decrease in stress commonly associated with a catastrophic failure after reaching peak value. This was attributed to the brittle nature of the epoxy resin system combined with the unidirectional fiber orientation which were also confirmed by SEM scans. An important dimension to this research was the biocompatibility tests including cytotoxicity and its effect on osteogenesis that were carried out a year later. Cytotoxicity evaluation was carried out using an MTT (3-(4,5-dimethylthiazol-2yl)-2,5-diphenyltetrazolium bromide) assay, direct and indirect tests, details of which are provided in the cited paper. To summarize, the MTT assay is a colorimetric analysis that relies on the ability of NAD(P)H-dependent oxidoreductace enzyme (secreted by mitochondria of living cells) to reduce the tetrazolium dye (MTT) to its insoluble purple-colored form, formazan. The general finding from this research was that the proposed composite would not negatively affect osteogenesis (the process of bone formation) when compared with conventional metallic fracture plates.

151

Table 5.2 Composite Preparation Techniques for Applications in Implants and Fixation Plates Reference

Application

Composite Preparation Technique

Bagheri et al. (2013)

Long-bone fixation plate and intramedullary nails

Cochran et al. (1994)

Long-bone fixation plate

Arun and Kanagaraj (2016)

Transfemoral prosthetic socket

This novel composite was prepared by adding 16 layers of prepeg epoxy/flax composite sheets (58% 60% vol. fraction) in between two outer layers of prepeg epoxy/carbon fiber sheets (57% vol. fraction), inserting all into a stainless-steel mold followed by hot pressing at 500 kPa pressure and 150 C for 60 min resulting in a composite 260 3 250 3 3.6 mm in dimension. Two types of prepeg epoxy/aramid plates, labeled as type I and type II plates depending on the plies and stacking sequences were prepared. For both types, a prepeg fabric ply with an orthogonal weave was rotated to ensure that its fibers were 6 45 degrees with respect to the long axis of the plate. Certain predetermined plies were aligned with fibers along the axial direction of the plate representing a 0-degree orientation with respect to the long axis of the plate. The fabric was so prepared that the orthogonal and axial fibers would be placed over each other resulting in different stacking sequences for each plate. The layered structure was then compressionmolded under heat to obtain the final composite. MWCNT chemically treated via heating with a mixture of nitric acid and sulfuric acid (1:3 volume ratio) at 140 C in an oil bath with continuous stirring for 30 min followed by deionized water wash to achieve a pH of 7. Fine powder of MWCNT was then recovered by heating the mixture in an oven at 100 C to remove the moisture content. FTIR tests reveal the presence of various functional groups on the surface of the MWCNT. Meanwhile, the epoxy resin is dissolved in acetone using a bath sonicator for 30 min and MWCNT dispersed in acetone using a tip sonicator. They were mixed together and sonicated for another 45 min. Acetone removed using a vacuum-drying process followed by the epoxy resin being hand-mixed with the hardener, poured into molds, and allowed to cure. The composite epoxy sandwich is then prepared using the hand lay-up technique using stockinette layers, E-glass woven fabric, and epoxy resin with MWCNT reinforcement.

5.3 Tribological Characterization of Green Composites

A notable development in the field of internal fixation plates has been the development of an epoxy/aramid system which was been found to possess acceptable short-term biocompatible properties (Cochran et al., 1994). The procedure that was utilized for fabricating the composite plates is outlined in Table 5.2. Furthermore, the catastrophic failure mode observed in epoxy/flax/carbon fiber composites was also not present in the epoxy/aramid composite when the system was subjected to mechanical tests. However, a major disadvantage was the lowyield strength in bending which, according to the authors, could be possibly improved by changing the fiber orientation. Due to the lower flexural strength that was observed, the authors found that the composite would be more suitable to support stable fractures as opposed to normal skeletal loads without undergoing permanent deformation.

5.3 TRIBOLOGICAL CHARACTERIZATION OF GREEN COMPOSITES FOR BIOMEDICAL APPLICATIONS When it comes to selecting a polymer and counterpart, friction and wear shouldn’t be the only factors influencing suitability for using a tribo-pair in biomedical applications like implants. The importance of biocompatibility and corrosion resistance are of prime importance. It is worth noting that several tribological characterization studies involve titanium alloys (Ti6Al4V) or stainless steel due to their extensive use in THR or TKR surgeries (Ruggiero et al., 2016). These alloys feature either as the head component in THR or as the femoral condylar components in TKR and any tribological testing on polymeric composites need to involve one of these materials as the counterface (Ruggiero et al., 2015; Guezmil et al., 2016). It was observed during our review that intricate details, such as the selection of counterface materials, have been paid attention to only by a very limited number of researchers who carried out tribological testing on epoxy systems.

5.3.1 BULK COMPOSITES As stated earlier, in the development of composite systems for biomedical applications, an evaluation of the tribological properties—the wear and frictional performance of a system under a range of varying environments—is essential. However, minimal studies from a tribological point of view on epoxy-based composites for biomedical applications have been documented in the literature compared to those available for thermoplastic systems. In one study, epoxy composites were prepared using cellulose nanofibers (CNF) followed by mechanical, thermal, and tribological characterization (Barari et al., 2016). The authors pointed out that cellulose as a reinforcement entity has gained much importance only in recent years owing to earlier processes requiring a huge amount of energy

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to be spent to separate cellulose from pulp. However, the use of modern technologies has rendered the production of CNF-based aerogels quite cost-effective. Aerogels are materials that are highly porous and lightweight. They are prepared by removing the liquid solvent from their matrix leaving behind a solid skeleton with a large number of pores. The commonly used production processes for aerogels are freeze drying or vacuum drying. In this study, silylated and nonsilylated CNF aerogels were used to reinforce a biobased epoxy matrix (named so because of the presence of biocontent exceeding 50% due to the epoxy used in this experiment being derived from waste products). The epoxy/CNF composites were prepared using a process known as liquid composite molding. The two main techniques involved in liquid composite molding are: (1) resin transfer molding process which utilizes positive pressures to wet out the reinforcements using the liquid (epoxy) matrix; and (2) vacuum-assisted resin transfer molding process— commonly known by its abbreviated form VARTM, and uses positive injection pressures while the mold is placed under vacuum—of which the latter was employed in this study. A detailed description of the process that was used is provided in Table 5.3. The effect of silane treatment on the mechanical properties of CNF composites were also investigated by the authors. Tests carried out using a tensile testing machine revealed the silylated CNF composites possessed a higher ultimate tensile strength and elastic modulus compared to untreated CNF composites. It was also noted that 1.4 vol.% of CNF filler loading produced consistently higher values of tensile strength as compared to 0.9 vol.% reinforcement concentration. As evident from the SEM micrographs shown in Fig. 5.3 (Barari et al., 2016), silane treatment improved the wettability of the fibers by the resin. Furthermore, the presence of silane ( SiO2) functional groups on the surface of the CNF also helped improve the mechanical entanglement between the fibers and the matrix during the curing process. Tribological tests were carried out using a pin-on-disk tribometer with a rotating steel disk as the counterface and under dry contact conditions. Loads of 4, 7, and 10 N were applied at varying sliding speeds of 0.15, 0.25, and 0.35 m s21. Since it is widely known that tribological properties are greatly influenced by mechanical properties, the authors limited their work to the tribological performance evaluation of the silylated CNF nanocomposite. Tests were repeated for varying filler concentrations of 0.9 and 1.4 vol.%. At low sliding speeds of 0.15 m s21, it was observed that under all loading conditions, the nanocomposites showed a reduction in the coefficient of friction values with a significant percentage reduction in the frictional coefficient occurring at high loads for 1.4 vol.% filler loaded epoxy nanocomposite. However, when the sliding speed was increased to 0.25 m s21 it was observed that although the frictional coefficient values decreased for 0.9 vol.% filler loading, it shot up again with increasing filler concentration at normal loads of 4 and 7 N. Under sliding speeds of 0.35 m s21, a trend similar to the one obtained at 0.25 m s21 with minimal exceptions were noted. In general, it could be concluded that for varying sliding speeds, the pristine epoxy displayed the best coefficient of friction

Table 5.3 Bulk Composite and Composite-Coatings Preparation Techniques for Tribological Characterization Bulk Composites Reference

Filler

Composite Preparation Technique

Barari et al. (2016)

Cellulose nanofibers

Omrani et al. (2015)

Carbon fabric

Dinesh et al. (2014)

Sisal fiber

The biobased epoxy used in the study was designed to be cured at room temperatures; a point highlighted by the authors as a necessity for VARTM processability. In this process, the free-dried cellulose aerogel preform was placed in the mold cavity with a polycarbonate cap used for fastening at the top. The inlet vent was connected to a beaker containing the epoxy resin while the outlet vent connected to a resin trap ensured that the resin would not leak into the vacuum pump. Next, the resin was forced to flow through the preform which, being an aerogel, has a porous structure. Initial curing for 24 h was carried out at room temperature followed by postcuring at 130 C for 20 min. The resulting composite was found to have volume fractions varying in discrete steps of 0.9% 1.4%. The composite was then machined to the required dimensions to obtain the final specimen suitable for testing purposes. In this method, the RTM technique was employed. The carbon fabric layers were stacked in the mold cavity made of aluminum and covered with plexi-glass (which also played the role of an air-seal) which was then fastened using bolts and clamps. An important point stated by the authors was the need to ensure proper fit of the carbon fabric within the mold cavity which if not done could seriously affect the resin flow and fiber/fabric wetting. The resin which was pushed using positive pressure into the mold cavity through an inlet located at the top of the center of the plexiglass, flowed horizontally and came out through vent holes provided at the ends. The specimen was left to cure for 24 h at ambient temperatures followed by machining to the required dimensions. Sisal fiber strands were cleaned using distilled water followed by drying under the sun. The dried natural fibers were then subjected to chemical cleaning by being dipped in a mixture of 80 vol.% NaOH and 20 vol.% distilled water followed again by drying under the sun. The dried fibers were then cut to 500 mm lengths and used for preparing the fabric for wetting out with resin by employing the hand lay-up technique followed by curing of the composite system at elevated temperatures and slight pressure. The specimens were then cut to dimensions stipulated by the ASTM G-99 standard. Composite Coatings

Reference

Filler

Yılmaz Atay et al. (2013)

Chitosan polymer

Substrate Glass

Composite Preparation Technique Varying amounts of chitosan with 84.5% deacetylation were dissolved in acetic acid solutions of two different concentrations (1 and 3% v/v) to produce colloids by stirring using an ultrasonic cleaner. The solutions were then stirred again in air using magnetic balls for 20 min. The resulting solutions were loaded into the epoxy and left to cure for 24 h at ambient temperatures followed by coating the same onto a glass substrate.

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FIGURE 5.3 (A) non-silylated CNF composite. (B) silylated CNF composite.

values at low, normal loads of 4 N while the nanocomposite reinforced with 1.4 vol.% CNF yielded the best frictional performance at high loading levels of 10 N. In other words, the pristine epoxy system would not be suitable for demanding tribological situations involving high loads and sliding speeds owing to their poor frictional performance. An increase in wear resistance was noted with increasing filler content. As expected, increasing normal loads produced a higher degree of wear in all tested specimens irrespective of filler-loading concentration. The reason for improvement in wear resistance with filler concentration was attributed to the improved load-bearing ability occurring in the reinforced matrix and the increased elastic modulus of the reinforced epoxy system. The wear mechanism was found to be primarily of the abrasive mode in the case of the neat pristine. A point highlighted by the authors was that the neat epoxy system under varying sliding velocities produced a significant amount of sharp wear debris which contributed to microcutting and removal of material from the epoxy matrix through abrasion. Furthermore, deep grooves were also observed on the surface of the pristine epoxy which was believed to be caused by the plowing action of the wear debris resulting in a huge amount of plastic deformation with the formation of ridges. On the other hand, only a minimal amount of wear debris was found to be produced for the reinforced epoxy composites, with the predominant wear mode having an adhesive nature wherein a transfer film was formed on the steel counterface within a few minutes of running-in. At higher normal loads, however, the detachment of CNF fibers from the matrix resulted in a reduction of the load-carrying capacity of the matrix resulting in higher wear rate. It must, however, be noted that the wear volume-loss was still significantly smaller than that of the pristine epoxy. Tribological evaluation of biobased epoxy composite systems were also considered by Omrani et al. (2015). In this study, carbon fabric (CF) was used as the reinforcement and the epoxy/CF composite was prepared using the resin transfer molding process, details of which are provided in Table 5.3. A pin-on-disk test

5.3 Tribological Characterization of Green Composites

rig was used to characterize the frictional and wear properties of the system under unlubricated contact conditions at ambient temperatures. Cylindrical samples, 3 mm in diameter, were prepared by machining and tests were carried out against a hardened 440C stainless steel disk as the counterpart with an initial surface roughness (Ra) of 0.3 6 0.05 μm. Applied normal loads were in the range of 10 20 N and sliding velocities were varied between 0.15 and 0.35 m s21 over a sliding distance of 1000 m. The experiments were designed using the Taguchi method. The studies also evaluated the effect of increasing CF reinforcement volume (10, 20, and 30 vol.%) on the mechanical and tribological properties. One of the points highlighted by the authors was the increase in Tg (glass transition temperature) of the system as confirmed by differential scanning calorimetry (DSC) tests (from 45 C corresponding to the pristine epoxy to about 86 C in the 30 vol. % CF composites) and was mainly attributed to the matrix-stiffening effect caused by the reduction in mobility of polymer chains due to the presence of CF reinforcements. Furthermore, three-point bending tests in suitably prepared composite specimens revealed a significant improvement in bending strength from 15 to 95 MPa with increasing CF content from 0 to 30 vol.% respectively. Tribological tests presented a decrease in frictional coefficient values with increasing CF content. At 30 vol.% the coefficient of friction was found to be approximately 0.15. This was primarily attributed to two reasons: (1) the enhanced load-bearing ability of the matrix owing to the presence of CF; and (2) due to the ability to reduce the direct contact area between the pin and disk by acting as a solid lubricant. It has been concluded from earlier studies using short carbon fiber reinforcements that delamination of carbon fibers from the matrix occurs due to a combination of multistep sequential events beginning with matrix wear and followed by fiber thinning, breakage, and delamination from the matrix (Chang and Zhang, 2006). This phenomenon can easily be extrapolated to CF which basically constitutes carbon fibers stacked into layers. An interesting finding was from the research was that (as opposed to the pristine epoxy matrix) the coefficient of friction values decreased with increasing normal loads at a certain sliding velocity with increasing reinforcement concentration. This was attributed to greater delamination of carbon fibers from the reinforced matrix (with increasing concentration of CF and normal loads) resulting in a higher amount of lubricating transfer film being generated. On the other hand, the wear volume was found to increase with increasing normal loads. This was attributed to an increase in adhesive forces at higher filler-loading resulting in a higher degree of fiber pull-out and accompanied by a greater degree of plowing action by the detached fibers, which is demonstrated by an increase in the depth and width of the grooves. Sisal (botanical name Agave sisalana) is a species of Agave. Sisal fibers have been selected by certain researchers to reinforce epoxy matrices to produce composites with potential use in orthopedic implants (Dinesh et al., 2014). Various reasons including biocompatibility, nontoxicity, processing ease, and recyclability were cited by the researchers as the reason for why sisal was chosen as the

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reinforcement. The authors also listed a set of desirable characteristics that must be possessed by any orthopedic implant; some of the key characteristics in addition those stated are: (1) noncarcinogenic; (2) possessing exceptional wear resistance; (3) appropriately designed in form; and (4) resistance to corrosion. In this study, the composites were prepared using the technique presented in Table 5.3. Three different weight fractions of sisal fibers namely, 10, 20, and 30 wt.% were used to prepare the composites. Tribological characterization was carried out using a pin-on-disk apparatus with a rotating EN31 high-carbon, alloy, steel disk as the counterface. Normal loads ranging from 20 to 80 N in four steps were applied and the wear volume and frictional coefficient values were recorded. To summarize, the increase in filler concentration from 10 to 30 wt.% was accompanied by a reduction in wear resistance and an increase in frictional coefficient. While noting that the best results were produced by the composite reinforced with 10 wt.% sisal fibers, further insight such as type of wear debris produced or wear mechanisms weren’t provided by the authors.

5.3.2 COMPOSITE COATINGS In our previous review work, we observed that the tribological characterization of composite epoxy coating systems constituted only a minor fraction of the total number of works carried out on epoxy systems from a tribological point of view (Bobby and Samad, 2016). It follows, therefore, that the number of works involving a combination of coatings (epoxy based), tribological characterization, and biomedical applications should be even more scarce. An intensive review proved that while both UHMWPE (Samad and Sinha, 2010, 2011; Ravi et al., 2015; Minn and Sinha, 2008) and PEEK (Nunez et al., 2011; Song et al., 2016) were definitely preferred materials in the composite coatings domain, even bioceramic coatings based on materials such as hydroxyapatite and titanium (Choudhuri et al., 2009) aren’t far behind when we consider research works focused on dental and orthopedic implants. This definitely points to a major gap in academia when it comes to thermosetting coatings based on the epoxy generic for biomedical applications. The significance of the investigations carried out by the few authors into the self-healing properties of chitosan-reinforced epoxy composite coatings is worth mentioning (Yılmaz Atay et al., 2013). Chitosan, as described by the authors, is a linear polysaccharide obtained by the deacetylation of chitin—a material that occurs naturally in the exoskeletons of crustaceans and in several other life forms including insects, internal shells of cephalopods, and soft tissues of fish. The reaction for the same is presented in Fig. 5.4 adapted from another work (Usman et al., 2016). The authors also state that it is often not possible to completely deacetylate chitosan and, therefore, the only real feature that distinguishes chitosan from chitin is the former’s ability to dissolve in diluted aqueous solutions. Chitosan was selected to be the reinforcement agent for various reasons, namely: (1) biocompatibility; (2) low toxicity; and (3) biodegradability. These

5.3 Tribological Characterization of Green Composites

FIGURE 5.4 Production of Chitosan from Chitin.

characteristics render it as potentially useful for a wide range of biomedical applications. According to a technical article on the Sigma-Aldrich website, the biocompatible properties of a polymer are often determined by its degradation entities. A second reason cited by Yılmaz Atay et al. for the selection of chitosan was due to its self-healing abilities. It was stated that such materials when embedded in the polymer matrix as an “active” state would respond to the formation of any microcracks or voids arising from damage or external stimulus by initiating a repair mechanism involving a mobile phase which would then move to the damaged site and close up the cracks by recreating the broken chemical bonds. Thirdly, both chitosan and chitin are known to possess excellent antimicrobial properties, though applications using this feature left much to be researched. The composite was prepared using the methodology detailed in Table 5.3. Six different specimens (with varying amounts of chitosan or acetic acid concentrations) were prepared. For two of the six specimens, the chitosan content was kept constant at 1.25 vol.%; however, the colloid was prepared for both using different amounts of acetic acid to determine whether there would be an effect on the selfhealing abilities due to the acid used for dissolving the chitosan. Since the authors’ primary aim was to investigate the self-healing properties, the coated specimens were subjected to a scratch test only (achieved using a thin pin). The characterization of coatings was carried out using the Fourier Transform Infrared Spectroscopy technique. The degree of the induced damage and subsequent healing were analyzed using SEM (field emission type). One of the first observations made by the authors was that different acid concentrations were found to not influence the coating morphology (scanning carried out using secondary electron imaging feature). SEM scans were taken at 30 minutes, 7, 20, and 35 days after the initial scratch was made to study the dynamic crack-healing behavior. In all samples except for the pristine epoxy, healing was found to have been initiated almost immediately with no change occurring after 30 minutes which led the authors to conclude that chitosan should have been the agent responsible for initiating the healing mechanism. As shown in Fig. 5.5 adapted from the referenced work (Yılmaz Atay et al., 2013), the healed crack was found to possess a morphology similar to the branches of a tree. The healing mechanism was elucidated by drawing references to earlier works where such branches spanning the crack were found to assist molecular diffusion of numerous chain segments across the polymer polymer interface. A final point

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FIGURE 5.5 (A) Epoxy composite (2.5 vol. % Chitosan by 3% v/v acetic acid) scratched using pin. (B) healed via intermolecular diffusion.

observed was that thin cracks healed better than wide ones. The conclusion drawn by the authors was that chitosan, while improving the scratch-resistance properties of the epoxy matrix by virtue of its self-healing abilities could be used to bridge thin cracks; however, it would not be possible to satisfactorily compare the degree of self-healing by simply measuring the distances between the splits due to the dependence on a wide range of factors namely, branch numbers, shapes, and thickness. No friction or wear tests were carried out.

5.4 DENTAL APPLICATIONS Epoxy composite systems have also been proposed by several authors for a range of applications in the dental field (Malquarti et al., 1990; Viguie et al., 1994; Tey and Lui, 2014). In one such study, the authors evaluated the effect of reinforcing pristine epoxy matrices with carbon fiber reinforcements as potential materials for fixed partial dentures (dental prosthesis used to replace missing teeth and attached to adjacent teeth using fixed bridges that may be screwed in or manually attached) (Viguie et al., 1994). In this work, carbon fiber fillers in three forms—(1) short (,5 mm) randomly distributed; (2) long unidirectional; and (3) woven type— were used to formulate the epoxy composites. Flexion and elastic modulus tests carried out on the specimens revealed that the composite reinforced with long unidirectional fibers presented the highest values with an increase in: (1) elastic modulus values by about 180% and 330%; and (2) flexural strength by about 129% and 230% when compared to woven and short fibers, respectively. Furthermore, it was noted that the composite containing long unidirectional fibers resisted

5.5 Research Works Based on Bioepoxy Resins

forces applied perpendicular to it in the three-point bending test as opposed to the composite reinforced with woven fibers, implying that the connectors of a fixed partial denture or crown manufactured from such composites should ideally contain a greater number of long fibers. However, since it is known that in addition to forces acting perpendicular to the occlusal plane, there may be also multiaxial forces coming into play in a real-world situation. The ideal fiber reinforcement proposed by the authors was a combination of fiber alignments including the unidirectional and woven type, with the former required to be in larger proportions. A point stressed by the authors was the need to carefully choose the reinforcement type based on the clinical situation at hand. It was pointed out that for fabricating posts and cores, a material with an elastic modulus closely matched to dentin (B18 GPa) would need to be considered. This point was also highlighted by Tey and Lui who prepared dowel posts from glass fiber-reinforced epoxy composites and found that a similar moduli (comparisons drawn between the composite and dentin) would mean better distribution of functional stresses across the bonding interfaces of the dentin/dowel/core system to the tooth with a potential to reinforce the weakened tooth in addition to reducing the probability of root fracture (Tey and Lui, 2014). From the tests, it was quite evident that the specimens prepared using short randomly distributed fibers, albeit presenting the lowest value of the parameter from the series of tests, met this requirement (B15.7 GPa). In a related study, three sets of prosthesis were prepared, namely: (1) individual crowns to determine the esthetic qualities, compressive strength, and wear of the material; (2) fixed partial dentures to verify the flexural strength of the material; and (3) splint bars to ensure that teeth are held in position before surgery and during the healing time (Malquarti et al., 1990). For preparing the composites, PAN-based fibers were used and subjected to superficial oxidation treatment to improve the bonding with the epoxy matrix. The major findings from the work on which monitoring was carried out across a time span of 3 years were: (1) the biocompatibility of the epoxy-carbon composite was found to be excellent; (2) medical examination revealed that magnetic resonance imaging could be performed without image distortion; (3) the esthetic properties were retained even after 3 years with only slight discoloration of the crown; (4) low density; and (5) excellent resistance to corrosion. Of the few limitations that were pointed out, one was that the abrasion resistance of this composite was less than that of ceramics or metal alloys.

5.5 RESEARCH WORKS BASED ON BIOEPOXY RESINS Any review on epoxy composites for biomedical applications is virtually incomplete without a brief mention of the works carried out using bioepoxy resins. The negative effects of bisphenol-A (BPA), its role in particular as an endocrine

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disruptor and as a toxic agent adversely affecting the cardiovascular and reproductive systems, has been extensively studied and documented by several authors as well as research organizations including the US National Institute of Environmental Health Sciences. The detailed study by Bacle et al. found that the risk of exposure to BPA via the intravenous route for a large number of patients suffering from end-stage renal disease and undergoing hemodialysis is extremely high due to the presence of this compound in water and medical devices used for the production of dialysate fluid involved in the treatment (Bacle et al., 2016). A point of concern raised by many researchers carrying out works on epoxy composites has been the overall toxicity of the system arising from the monomers in the partially cured resin (Scholz et al., 2011). In this regard, epoxy resins based on rosin—obtained from exudation of pine trees and containing B90% abietic acid or its monomers (Atta et al., 2004)—have been used by a set of researchers for preparing non-BPA-based epoxy composites for potential applications in the medical field (Huo et al., 2016). Epoxies generally cure via a polymerization reaction between a base containing bisphenol A or F and amines (straight chain, branched, or cyclic). In the referenced work, however, since the base component was rosinbased, the authors resorted to using a polyurethane blend—obtained by the reaction between castor oil (a triglyceride with hydroxyl groups) and toluene-2,4-diisocyanate—for crosslinking the epoxy resin. It was postulated that the presence of polyurethane groups would reduce the brittleness of the resulting cured system while improving the impact strength several fold. Carbon nanotubes are known to improve the mechanical and tribological properties of pristine epoxy matrices (Khun et al., 2013; Garg et al., 2015). To enhance the properties of the pristine matrix, the authors also considered the addition of multiwalled, carbon nanotubes to the curing resin to improve its properties. Cytotoxicity and thermal characterization of the composites were carried out using various techniques. The detailed preparation techniques pertaining to the base and solidifier components are provided in their paper; however, since this is not within the scope of the current review, only the dispersion technique used for the incorporation of CNT and the overall curing mechanism of the epoxy resin will be discussed. Incorporation of CNT within the matrix was achieved by dispersing 0.01 g of CNT (equivalent to 0.2 wt.% of the total resin weight) in 25 mL of dichloromethane followed by ultrasonication for 1 hour. This was followed by the addition of 3 g of base (maleopimaric acid epoxy resin; MPAER) and 2 g of solidifier (castor oil-based polyurethane; COPU) to the CNT/dichloromethane solution while stirring constantly, followed by another 1 hour cycle of ultrasonication. The solvent was removed from the mixture using a vacuum and stored in a refrigerated environment. Prior to testing, all specimens were subjected to a cure temperature of 160 C for 3 hours. Samples with different weight proportions of CNT were also produced using the same technique for comparison purposes. Ultrathin films of the composites were prepared and moved to copper grids followed by transmission electron microscopic (TEM) analyses to study the morphology and dispersion of CNT particles within the matrix. Meanwhile, for dynamic mechanical analyses, glass fiber

5.6 Composite Shape Memory Polymers for Biomedical Applications

laminates impregnated with MPAER, COPU, or CNT resin were prepared by the hot press method. Any solvent traces were removed by vacuum. For cytotoxicity studies, HeLa cells were exposed to different concentrations of the composites for 48 and 72 hours and cell-viability was determined using the MTT assay. To begin with, the curing reaction, was found to take a complex path with the first reaction occurring between the epoxy group of the epoxy resin and the primary amine structure on the polyurethane. This reaction led to the generation of a secondary amine which underwent another reaction with the epoxy leading to the generation of a tertiary amine to produce the final crosslinked structure. An interesting point noted by the authors was that the Tg (the temperature at which a polymer transforms from a hard “glassy” state to a soft “rubbery” state) of the pristine epoxy was 97.6 C whereas this value went up to around 150 C for the composite possessing a filler loading of 0.4 wt.%. This was attributed to two factors. Firstly, to the decreased mobility of the polymer chains arising from: (1) the π-π interactions between the resin molecules and CNTs; and (2) the reaction between hydroxyl groups on the surface of the CNT with the epoxy groups of the MPAER. Secondly to the improvement in thermal degradation properties due to the absorption of free radicals generated by the π-bond in CNT leading to the termination of the free-radical reaction. A practical significance of this value is that the composite could be recommended for applications involving higher temperatures without the risk of degradation. However, a further increase in CNT content was found to bring this value down to B140 C owing to the agglomeration of nanoparticles as seen from the TEM scans. Cytotoxicity test results were also encouraging with a cell-viability value .90% after 48 and 72-hour incubation.

5.6 COMPOSITE SHAPE MEMORY POLYMERS FOR BIOMEDICAL APPLICATIONS Shape memory polymers (SMPs) belong to a class of materials called smart materials that exhibit stimuli-dependent behavior, and extensive reviews are available on this topic (Xie et al., 2016; Pretsch, 2010). Such materials assume a temporary shape when external stimuli, such as a mechanical load, electric or magnetic field, is imposed upon them. Once the applied stimulus is removed, the material returns to its original form. Innovative concepts such as temporary implants and drug release control systems using SMP are currently being researched. In one study on composite an epoxy-based shape-memory system targeting potential applications in the biomedical field, Zhang et al. utilized a novel technique to design an epoxy-polycaprolactone (PCL) (a type of polyester) system (Zhang et al., 2015). PCL was selected as the reinforcement due to its biodegradable properties (Yeganeh et al., 2005) combined with low cost, easy availability, and due to possessing the ability to impart hydrophobic properties to the polymer matrix to which it has been coupled. Besides these, the shape fixity and recovery

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FIGURE 5.6 Set-up of Epoxy/PCL electrospinning process.

ratios—the respective ability of an SMP to fix and recover shapes measured in terms of different strain ratios—of PCL were also cited to be excellent, thereby making it an ideal candidate for applications requiring shape-memory properties. In this study, the epoxy/PCL composite was prepared using the electrospinning technique because of its known efficiency in the production of fibers with micro/ nano level diameters. The formulation process involved mixing two compounds, namely dichloromethane (CH2Cl2) and dimethylformamide (DMF), in a volumetric ratio of 4:1, followed by the dissolution of PCL in the mixed solvent at a concentration of 15 wt.%. Afterwards, an UV initiator, benzophenone, was added to the mixture at 10 wt.% concentration. The prepared solution was then stirred using a magnetic stirrer at room temperature. Next, to obtain a core/shell structure, a set-up as shown in Fig. 5.6 was used (Zhang et al., 2015). The feed-rate of the outer shell forming PCL was set at 0.002 mm s21. Specimens having four different core diameters were prepared by varying the feed-rate of the epoxy. Ratios of 1:1, 2:1, 2.5:1, and 4:1 of feed-rates of the epoxy to PCL were used to prepare the composites. Finally, the obtained composites were cured at 70 C in an oven for 75 hours. Characterization was carried out using a combination of techniques which will not be elaborated further here. However, an important point stated by the authors is that they were able to successfully prepare epoxy fibers (not conducive to electrospinning) through the coaxial electrospinning process. To clarify, the shell liquid (PCL) was drafted by the electrostatic field (through the formation of the Taylor cone) and the core liquid (epoxy) was drawn into fibers by frictional force. This was further confirmed by the fact that when the feed-rate of the epoxy system was increased to match with that of PCL, there was virtually no core of epoxy to the “composite” and only the PCL outer shell was visible in the TEM scans. Furthermore, Zhang et al. also mentioned that the immiscibility of the two phases combined with the lower feed-rates of the epoxy core could aid the formation of continuous fibers. DSC tests revealed a higher melting point of the epoxy/ PCL composite compared to the pristine PCL system. Improvements were also

5.6 Composite Shape Memory Polymers for Biomedical Applications

seen in the values of mechanical strength and elongation at break, as confirmed by nanotension tests. Shape-memory behavior of the produced composite was verified using SEM imaging. The original shape was stretched along the horizontal direction leading to the attainment of a temporary shape. Heating of this polymer composite to above the Tg value was found to result in shape-memory recovery. Another interesting point revealed by the SEM scans was the presence of pores throughout the matrix of the polymer composite leading the authors to propose this composite system for tissue-engineering applications. In addition to this, the cytotoxicity of both the pristine PCL and epoxy/PCL composite was evaluated using the CCK-8 assay. The cell viability of both samples was found to be greater than 100% after 24 hours and greater than 80% after 96 hour cultivation. In a recent work, an SMP was synthesized from biobased epoxy resins (Li et al., 2016). As explained by the authors, the main reason for selection of rosin to produce the epoxy was due to its nontoxic properties compared to the diglycidylether of bisphenol-A (DGEBA) precursor commonly used to produce epoxies. Another reason cited by the authors was due to its ability to impart shapememory abilities to the polymer matrix as confirmed in previous works by the same authors. In this study, the authors first produced acrylic pimaric acid (APA) through a series of steps which are well-detailed in the research. The APA so produced was, in turn, used to produce diglycidylether of APA (DGAPA). For comparison purposes, another monomer namely, diglycidylether of terephthalic acid (DGT) was also prepared. This was followed by crosslinking of the monomers through the addition of a curing agent, D230, poly(propyleneglycol)-bis (2-aminopropyl ether) in a volumetric ratio of 1:1. This was followed by the addition of CH2Cl2 to the base-solidifier mixture to dissolve all components before stirring for about 20 minutes at room temperature, followed by heating under vacuum in an oven at 60 C for about 2 hours to remove CH2Cl2, and then post-curing at elevated temperatures to obtain the final specimens. Shape-memory tests revealed both high shape-fixity and shape-recovery ratios. Though not a composite, this recent work was included in the review to highlight the importance of biobased epoxy systems in the production of SMP. Epoxy composites possessing shape-memory properties have also been prepared using CF reinforcements (Fejos and Karger-Kocsis, 2013). The motivation for this line of research was due to the lack of works in temperature-sensitive, shape-memory composites prepared from asymmetric layering of fabrics. Epoxy/ CF composites, two layered and four layered, were prepared by placing the fabrics on one side (top or bottom) of the epoxy matrix resulting in a reinforcement asymmetry. The thickness of all specimens was set at B2 mm. Furthermore, all specimens were subjected to both three-point bending as well as shape-recovery tests. It was noticed that the recovery stress (maximum stress measured during reheating of the composite) was enhanced by the presence of CF reinforcement with a higher value observed when the CF was located on the tension side of the bent specimens compared to when the reinforcements were on the compressive side. The improvement in recovery stress was found to compromise the shape-

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recovery effect, albeit marginally, as a function of the fabric reinforcement content. This effect was also confirmed by earlier studies on other fiber-reinforced thermosetting resins (Ivens et al., 2011).

5.7 GENERAL BIOMEDICAL APPLICATIONS Several other biomedical applications are also worth mentioning; however, we will only provide a brief mention of a few of these. One area where epoxy systems have found much use is in the development of piezo-composites for ultrasonic systems in medical imaging (Cannata et al., 2006; Ritter et al., 2002; Gururaja et al., 1985). Epoxy systems coated with 60 nm thick nanolayers of titania have been used to produce positive replicas with extremely minimal variation in surface topographies to improve the reproducibility of cell cultural assays (Schuler et al., 2009). Certain older works have also mentioned the use of epoxy resins based on a range of precursor chemistries for preparing bioadhesives for tissue bonding and hemostasis (Sung et al., 1999; Meyer et al., 1979). Another area where epoxy composites have found use is in the manufacture of electrodes and platforms for biosensing applications (Castan˜eda et al., 2007; Xiaomi and Stephen, 2009; Ninoska and Eliana, 2011; Smith and Lamprou, 2014). A detailed review paper on composite electrodes is available in the literature (Ninoska and Eliana, 2011). In this comprehensive work, the authors begin by explaining the basic principle of a biosensor followed by the role of the three phases, namely: (1) receptor (incorporating the biological or biomimetic recognition element); (2) transducer (also known as the conductive phase which works by converting the biochemical signal owing to the interaction between an analyte of interest with a biological entity into a reading/measurement); and (3) amplifier for boosting the signal. This study also provides a detailed explanation of the properties of various forms of carbon utilized in the synthesis of composites for such applications along with an emphasis on the preparation techniques. The significant gap in hybrid bactericidal polymers was addressed by Santhosh and Natarajan who prepared an epoxy composite using silver ion dopedTiO2 (Santhosh and Natarajan, 2015). The reasons stated by the authors for the selection of titania as one of the fillers were its proven ability to photocatalytically inactivate bacteria such as Staphylococcus aureus and Escherichia coli (which are known to form biofilms over a variety of substrates). However, the authors reasoned from their detailed literature review that such bare TiO2 films (in doped or undoped form) when applied as a coating wouldn’t be able to retain their photocatalytic efficiency for long due to the mass transfer effect. It was postulated that immobilization of titania by dispersing it into a polymeric phase would hold the key to retaining the said properties for longer. In addition to this, and in order to achieve a synergistic effect, doping of titania using silver ions was

5.7 General Biomedical Applications

considered. This was intended to ensure a combination of pthe hotocatalytic effect of the former coupled with the improved biocidal abilities imparted by the latter. Furthermore, it was also mentioned that the addition of titania would serve to improve the mechanical properties of the pristine epoxy matrix along with an enhancement in crack resistance and photostability of the epoxy polymer under UV exposure. In this study, six different specimens—pristine epoxy, epoxytitania, and epoxies with four different filler loading concentrations of silver ions (epoxy-silver ion doped titania)—were compared. One of the first observations made from SEM characterization studies was that the undoped titania particles retained their nanosizes whereas silver ion-modified titania particles were found to possess dimensions in the order of micrometers. However, while both the unmodified as well as modified titania particles were found to be dispersed homogenously through the polymer matrix, the degree of agglomeration was found to be less for the silver ion-modified titania fillers. A reason apart from sonication, as indicated by the authors, was the improvement in dispersion of the fillers provided by lowering the viscosity of the base resin that was achieved by the addition of a diluent. The Tg of the epoxy nanocomposite prepared by blending in doped titania was found to be B107 C and higher than that observed for the pristine epoxy by about 15 C. Antibiofilm formation properties of the epoxy polymeric samples were tested under bright (UV-A irradiation) and dark conditions and the following was observed: (1) under exposure to UV light, the highest antibiofilm activity was exhibited by the silver-titania epoxy nanocomposite whereas the pristine epoxy matrix showed the highest biofilm formation; (2) under dark conditions, the reduction in the photokilling effect resulted in lowered antibiofilm activity for both the modified and unmodified titania nanocomposite coatings; however, these values were still better than the pristine epoxy system. In short, the doping of titania using silver ion particles and dispersing them within the epoxy matrix were cited to open new doors of research in the biomedical engineering arena. Tissue engineering is another discipline within the biomedical realm where epoxy composites have found much use. Tissues in the body that may be damaged by disease, injury, or trauma may be treated by various methods namely, autograft (in which tissue is transplanted from one location to another within the same patient), allograft (in which tissue is transplanted from one individual to another), and, more recently, tissue engineering which seeks to regenerate the damaged tissues by relying on porous 3D scaffolds that provide the right environment for the growth of tissues and organs (O’Brien, 2011). In the referenced work, O’Brien explains that the role of the scaffold is essential in providing a template for the growth of tissues over it. For a tissue to grow, it must be seeded with cells and subjected to biophysical stimuli using a bioreactor. The cell-seeded scaffolds may then be cultured in vitro after which they can be implanted in the injury site or in vivo where the body’s systems will ensure regeneration. To summarize, the combination of cells, signals, and scaffolds constitute the tissue engineering triad. Some of the essential features, according to the author, that a

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scaffold must possess are: (1) biocompatibility; (2) biodegradability; (3) mechanical properties; (4) scaffold architecture with optimized critical pore sizes; and (5) manufacturing technology. In a study by Barua et al., hyperbranched epoxy-nano clay composites were developed by embedding the nanocomposite with silver nanoparticles and was found to improve their antimicrobial efficiency and use as scaffolds in vivo (Barua et al., 2015). The cytocompatibility of the nanocomposite with primary cardiac and liver cell lines was verified using MTT assay tests which revealed an average cell survival rate of over 90% for every sample tested. This was attributed to the presence of glycerol and nanoclay in the matrix structure, which is known to possess biocompatible properties. Covalent functionalization of single walled carbon nanotubes (SWCNT) and surface modification of nanodiamonds have been carried out by Khabashesku et al. (2005). Various techniques were presented in the referenced paper for sidewall functionalization of SWCNT and nanodiamonds. The functionalization of both the SWCNT and nanodiamonds were achieved in two steps: (1) fluorination of the reinforcement; and (2) addition of dangling bonds by chemical treatment. The authors reasoned that such functionalization would improve the bonding between the reinforcement and the epoxy matrix and, in turn, reduce the probability of fiber pull-out and ensuring an effective transfer of the load from the matrix to the reinforcement, thereby enhancing the mechanical properties.

5.8 SUMMARY OF RESEARCH WORKS IN EPOXY COMPOSITES FOR BIOMEDICAL APPLICATIONS Table 5.4 provide the reader with a summary of 15 important works reviewed in this chapter in a concise form for quick reference:

5.9 CONCLUSIONS Epoxy composite systems have found their use within a wide range of disciplines, both directly and indirectly, in various biomedical applications. An important point observed during this focused review was the sheer number of works carried out recently which further underlines the gaining popularity of composites and, in particular, thermosetting resins such as epoxies as a material for in vivo use. It has been pointed out that in developing nations, for example, affording a highly expensive carbon fiber/epoxy composite prosthetic or implant may not be always feasible. However, with advanced technologies becoming available for production and dispersion of nanoparticles within the epoxy matrix at a reasonable cost, it is expected that some of the reviewed works can definitely be scaled-up to volumes justifying commercial ventures.

5.9 Conclusions

Table 5.4 Summary of Research Works Reference

Composite Type

Arun and Kanagaraj (2016)

Sandwich

Petersen (2014)

Filler dispersed Filler dispersed

Hou et al. (2016)

Reinforcement

Application (Section)

Epoxy/MWCNT/Eglass woven fabric/ stockinette Epoxy/carbon fiber

Transfemoral prosthetic socket (5.2.1) Review paper (5.2.1)

Epoxy-n-TiO2CaO; Epoxy-nTiO2-μ-TiO2-CaO; Epoxy/Polyester-nTiO2-CaO; Epoxy/Polyester-nTiO2-μ-TiO2-CaO Epoxy/carbon fiber/flax

Coatings on titanium implants (5.2.1)

Bagheri et al. (2013), Bagheri et al. (2015), Bagheri et al. (2014b)

Sandwich

Cochran et al. (1994)

Sandwich

Epoxy/aramid

Barari et al. (2016)

Aerogel

Omrani et al. (2015)

Sandwich

Dinesh et al. (2014)

Sandwich

Epoxy/cellulose nanofibers Epoxy/carbon fabric Epoxy/sisal fabric

Yılmaz Atay et al. (2013)

Filler dispersed Filler dispersed

Epoxy/chitosan

Filler dispersed Filler dispersed

Epoxy/glass fiber

Filler dispersed Sandwich type

MPAER/COPU/ CNT; MPAER/COPU/ CNT/glass laminate

Viguie et al. (1994)

Tey and Lui (2014) Malquarti et al. (1990)

Huo et al. (2016)

Epoxy/carbon fiber

Epoxy/PAN-carbon fiber

Long-bone fixation plates and intramedullary nails (5.2.2) Long-bone fixation plate (5.2.2) Tribological studies (5.3) Tribological studies (5.3) Tribological studies (5.3) Tribological studies (5.3) Fixed partial-denture, crown, post, and core (5.4) Post and core (5.4) Fixed partial-denture, crown, splint bars (5.4) Biobased epoxy composites (5.5)

Various biocompatibility and mechanical tests carried out on these advanced epoxy composites combined with the encouraging results observed could also help assuage any concerns potential users could have while opting to go with implants, tissue scaffolds, and prosthesis manufactured from such materials.

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It may be observed that only a limited number of works involve such tests and this definitely presents a significant gap in the literature when it comes to presenting an epoxy-composite solution to problems posed by the biomedical field. In comparison, thermoplastics such as UHMWPE and PEEK, which are also quite popular biomaterials, have a large number of works focused only on the tribological aspect—an important dimension that needs to be added to any research work involving composites for implants or prosthesis due to the potential of the generated wear debris producing an inflammatory response. Such an extensive study was found to be carried out by a very limited number of researchers during the course of this review, probably because tribological testing on epoxy composites have been mainly limited to enhancing the frictional and wear characteristics for diverse industrial applications. It is hoped that future work along these lines could address the biocompatibility and cytotoxicity aspects of such composites and vice versa.

REFERENCES Arun, S., Kanagaraj, S., 2016. Performance enhancement of epoxy based sandwich composites using multiwalled carbon nanotubes for the application of sockets in trans-femoral amputees. J. Mech. Behav. Biomed. Mater. 59, 1 10. Available from: https://doi.org/ 10.1016/j.jmbbm.2015.12.013. Atta, A.M., Mansour, R., Abdou, M.I., Sayed, A.M., 2004. Epoxy resins from rosin acids: synthesis and characterization. Polym. Adv. Technol. 15, 514 522. Available from: https://doi.org/10.1002/pat.507. Bacle, A., Thevenot, S., Grignon, C., Belmouaz, M., Bauwens, M., Teychene, B., et al., 2016. Determination of bisphenol A in water and the medical devices used in hemodialysis treatment. Int. J. Pharm. 505, 115 121. Available from: https://doi.org/10.1016/ j.ijpharm.2016.03.003. Bagheri, Z.S., El Sawi, I., Schemitsch, E.H., Zdero, R., Bougherara, H., 2013. Biomechanical properties of an advanced new carbon/flax/epoxy composite material for bone plate applications. J. Mech. Behav. Biomed. Mater. 20, 398 406. Available from: https://doi.org/10.1016/j.jmbbm.2012.12.013. Bagheri, Z.S., Tavakkoli Avval, P., Bougherara, H., Aziz, M.S.R., Schemitsch, E.H., Zdero, R., 2014a. Biomechanical analysis of a new carbon fiber/flax/epoxy bone fracture plate shows less stress shielding compared to a standard clinical metal plate. J. Biomech. Eng. 136, 091002. Available from: https://doi.org/10.1115/1.4027669. Bagheri, Z.S., El Sawi, I., Bougherara, H., Zdero, R., 2014b. Biomechanical fatigue analysis of an advanced new carbon fiber/flax/epoxy plate for bone fracture repair using conventional fatigue tests and thermography. J. Mech. Behav. Biomed. Mater. 35, 27 38. Available from: https://doi.org/10.1016/j.jmbbm.2014.03.008. Bagheri, Z.S., Giles, E., El Sawi, I., Amleh, A., Schemitsch, E.H., Zdero, R., et al., 2015. Osteogenesis and cytotoxicity of a new Carbon Fiber/Flax/Epoxy composite material for bone fracture plate applications. Mater. Sci. Eng. C 46, 435 442. Available from: https://doi.org/10.1016/j.msec.2014.10.042.

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Huo, L., Wang, D., Liu, H., Jia, P., Gao, J., 2016. Cytoxicity, dynamic and thermal properties of bio-based rosin-epoxy resin/castor oil polyurethane/carbon nanotubes bionanocomposites. J. Biomater. Sci. Polym. Ed. 5063, 1 15. Available from: https://doi. org/10.1080/09205063.2016.1183332. Ivens, J., Urbanus, M., De Smet, C., 2011. Shape recovery in a thermoset shape memory polymer and its fabric-reinforced composites. Express Polym. Lett. 5, 254 261. Available from: https://doi.org/10.3144/expresspolymlett.2011.25. John, A., 1999. Prosthetic primer: transfemoral prosthetics, above knee. InMotion 9. Khabashesku, V.N., Margrave, J.L., Barrera, E.V., 2005. Functionalized carbon nanotubes and nanodiamonds for engineering and biomedical applications. Diam. Relat. Mater. 14, 859 866. Available from: https://doi.org/10.1016/j.diamond.2004.11.006. Khun, N.W., Zhang, H., Lim, L.H., Yang, J., Liu, E., 2013. Mechanical and tribological properties of epoxy matrix composites modified with microencapsulated mixture of wax lubricant and multi-walled carbon nanotubes. Friction 1, 341 349. Li, T., Liu, X., Jiang, Y., Ma, S., Zhu, J., 2016. Bio-based shape memory epoxy resin synthesized from rosin acid. Iran. Polym. J. 25, 957 965. Available from: https://doi.org/ 10.1007/s13726-016-0482-0. Malquarti, G., Berruet, R.G., Bois, D., 1990. Prosthetic use of carbon fiber-reinforced epoxy resin for esthetic crowns and fixed partial dentures. J. Prosthet. Dent. 63, 251 257. Mavrogenis, A.F., Dimitriou, R., Parvizi, J., Babis, G.C., 2009. Biology of implant osseointegration. J. Musculoskelet. Neuronal Interact. 9, 61 71. Available from: https://doi. org/10.1016/j.joms.2007.05.013. Meyer, G., Muster, D., Schmitt, D., Jung, P., Jaeger, J.H., 1979. Bone bonding through bioadhesives: present status. Biomater. Med. Dev. Artif. Organs 7, 55 71. Minn, M., Sinha, S.K., 2008. DLC and UHMWPE as hard/soft composite film on Si for improved tribological performance. Surf. Coatings Technol. 202, 3698 3708. Available from: https://doi.org/10.1016/j.surfcoat.2008.01.012. Ninoska, B., Eliana, A., 2011. Graphite-composites alternatives for electrochemical biosensor. In: John, C. (Ed.), Met. Ceram. Polym. Compos. Var. Uses. Nunez, E.E., Yeo, S.M., Polychronopoulou, K., Polycarpou, A.A., 2011. Tribological study of high bearing blended polymer-based coatings for air-conditioning and refrigeration compressors. Surf. Coatings Technol. 205, 2994 3005. Available from: https://doi.org/ 10.1016/j.surfcoat.2010.11.008. O’Brien, F.J., 2011. Biomaterials & scaffolds for tissue engineering. Mater. Today 14, 88 95. Available from: https://doi.org/10.1016/S1369-7021(11)70058-X. Omrani, E., Barari, B., Dorri Moghadam, A., Rohatgi, P.K., Pillai, K.M., 2015. Mechanical and tribological properties of self-lubricating bio-based carbon-fabric epoxy composites made using liquid composite molding. Tribol. Int. 92, 222 232. Available from: https://doi.org/10.1016/j.triboint.2015.06.007. Petersen, R.C., 2014. Titanium implant osseointegration problems with alternate solutions using epoxy/carbon-fiber-reinforced composite. Metals (Basel) 4, 549 569. Available from: https://doi.org/10.3390/met4040549. Pretsch, T., 2010. Review on the functional determinants and durability of shape memory polymers. Polymers (Basel) 2, 120 158. Available from: https://doi.org/10.3390/ polym2030120.

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Usman, A., Zia, K.M., Zuber, M., Tabasum, S., Rehman, S., Zia, F., 2016. Chitin and chitosan based polyurethanes: a review of recent advances and prospective biomedical applications. Int. J. Biol. Macromol. 86, 630 645. Available from: https://doi.org/ 10.1016/j.ijbiomac.2016.02.004. Viguie, G., Malquarti, G., Vincent, B., Bourgeois, D., 1994. Epoxy/carbon composite resins in dentistry: mechanical properties related to fiber reinforcements. J. Prosthet. Dent. 72, 245 249. Xiaomi, X., Stephen, G.W., 2009. Carbon fiber/epoxy composite ring-disk electrode: fabrication, characterization and application to electrochemical detection in capillary high performance liquid chromatography. J. Electroanal. Chem. 630, 75 80. Available from: https://doi.org/10.1016/j.pestbp.2011.02.012.Investigations. Xie, F., Huang, L., Leng, J., Liu, Y., 2016. Thermoset shape memory polymers and their composites. J. Intell. Mater. Syst. Struct. 27. Available from: https://doi.org/10.1177/ 1045389X166342111045389X16634211-. Yeganeh, H., Lakouraj, M.M., Jamshidi, S., 2005. Synthesis and properties of biodegradable elastomeric epoxy modified polyurethanes based on poly(??-caprolactone) and poly(ethylene glycol). Eur. Polym. J. 41, 2370 2379. Available from: https://doi.org/ 10.1016/j.eurpolymj.2005.05.004. Yılmaz Atay, H., Eral Do˘gan, L., C ¸ elik, E., 2013. Investigations of self-healing property of chitosan-reinforced epoxy dye composite coatings. J. Mater. 2013, 1 7. Available from: https://doi.org/10.1155/2013/613717. Zhang, F., Zhang, Z., Liu, Y., Cheng, W., Huang, Y., Leng, J., 2015. Thermosetting epoxy reinforced shape memory composite microfiber membranes: fabrication, structure and properties. Compos. Part A Appl. Sci. Manuf. 76, 54 61. Available from: https://doi. org/10.1016/j.compositesa.2015.05.004.

CHAPTER

Polyethylene and polypropylene matrix composites for biomedical applications

6

Aravinthan Gopanna1,2, Krishna Prasad Rajan2,3, Selvin P. Thomas1,3 and Murthy Chavali4,5 1

Advanced Materials Laboratory, Yanbu Research Center, Royal Commission for YanbuColleges and Institutes, Yanbu Industrial City, Kingdom of Saudi Arabia 2School of Chemical Engineering, Vignan’s Foundation for Science, Technology and Research University (VFSTRU; Vignan’s University), Guntur, India 3Department of Chemical Engineering Technology, Yanbu Industrial College, Royal Commission Colleges & Institutes, Yanbu Industrial City, Kingdom of Saudi Arabia 4Shree Velagapudi Ramakrishna Memorial College, Acharya Nagarjuna University, Guntur, Andhra Pradesh, India 5MCETRC, Tenali, Guntur, Andhra Pradesh, India

6.1 INTRODUCTION Polyolefins are a class of commodity thermoplastics that are the most widely produced and consumed all over the world. Their low cost, easy availability, ease of processing, light weight, and easy recyclability are some of the reasons behind their wide popularity, mainly as packaging materials. Polyolefins can be categorized, on the basis of their monomeric repeating units and polymer chain structures, into ethylene-based polyolefins, propylene-based polyolefins, higher polyolefins, and polyolefin elastomers (Gahleitner, 2001). Ethylene-based polyolefins are predominantly of two types: mostly linear highdensity polyethylene (HDPE) usually manufactured under conditions of low pressure utilizing transition metal catalysts, and mostly branched low-density polyethylene (LDPE) prepared under conditions of high pressure using oxygen or peroxide initiators (Nwabunma and Kyu, 2008). Transition metal-based catalysts are used for the production of propylene-based polyolefins, which results in linear chain structures having a stereospecific arrangement of propylene units. Transition metal-based catalysts used for the preparation of higher polyolefins also result in linear and stereospecific chain structures. Metallic or single-site catalysts are also used for the production of polyolefin elastomers containing a mixture of ethylene and propylene. Dienes are normally included in these types of elastomers for the purpose of crosslinking. These types of polymers are predominantly amorphous and possess a heterogeneous type of phase structure and high molecular weights Materials for Biomedical Engineering: Thermoset and Thermoplastic Polymers. DOI: https://doi.org/10.1016/B978-0-12-816874-5.00006-2 © 2019 Elsevier Inc. All rights reserved.

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(Vasile, 2000). It can be summarized that a polyolefin can be considered as a homopolymer, copolymer, or terpolymer based on the type of monomers used in the polymerization process. The crystallinity or amorphous region in the polyolefin depends on the arrangement of the polymer chains, chain configuration, and also on the conditions applied during processing (Brydson, 1999). Research advancements in the area of single-site transition metal-based catalysts have helped in the development of novel polyolefin homopolymers, copolymers, and terpolymers with controlled molecular architecture, microstructure with a wide range of molecular weights, and molecular weight distributions. These advancements have helped expand the range of applications of polyolefins. At present, polyolefins and polyolefin-based polymeric materials are finding uses in various applications that are spread across many fields. These applications include packaging, consumer products (toys, household items, appliance body parts, etc.), transportation (automotive and aerospace components), biomedical, communication and electronics, cable and wire coatings, thermal, electrical, and acoustic insulation, building and construction products, etc., to name a few (AlMa’adeed and Krupa, 2015). Polyolefins can be easily processed into many shapes. They can be extruded into fibers or filaments, blown films and cast films, and pipes and profiles. They can easily be compression molded or injection molded into various shapes. They can also be foamed into cellular plastics with the help of various foaming additives and physical or chemical blowing agents. They can also be coated onto other materials. Polyolefin homopolymers, copolymers, and terpolymers can be produced through free radical or ionic polymerization of corresponding alkenes with the help of conventional free radical initiators, such as peroxides and organometallic-type catalysts (Ziegler Natta or metallocenes). Developments in the field of polyolefin polymerization technology and unique catalyst systems have helped produce polyolefins with a wide range of structures, configurations, morphologies, molecular weights, and molecular weight distributions. Subsets of polyolefin homopolymers are polyethylene (PE), polypropylene (PP), polybutylene (PB), poly-1 methylpentene (PMP), and higher polyolefins. Out of these polyolefin materials, PE and PP are the most produced polymers by all worldwide polyolefin manufacturers (White and Choi, 2005a,b). PE can again be subdivided based on its chain structures, crystallinity, and density, into HDPE, LDPE, linear low-density polyethylene (LLDPE), ultralow density polyethylene (ULDPE), high molecular weight polyethylene, and ultrahigh molecular weight polyethylene (UHMWPE). PP and the other higher polyolefins are commercially manufactured with three main stereospecific arrangements: isotactic, syndiotactic, and atactic (Nwabunma and Kyu, 2008).

6.2 POLYOLEFIN COMPOSITES Polyolefin composites can be considered as a subclass of polymer composites (Nwabunma and Kyu, 2008). Polyolefin-based composites are developed to

6.4 Biocompatibility Evaluation of Polyolefin-Based Biocomposites

address the demand for higher load-bearing materials that are required for engineering applications which cannot be satisfied by polyolefins alone. In addition to the base polyolefin matrix, polyolefin composites contain at least one nonpolymeric additive which acts as a reinforcement. The incorporation of this additive can lead to the development of polyolefin composites that show enhanced overall properties of the resulting composite structure. Nonpolymeric additives that are used for the fabrication of polyolefin composite materials can be of different sizes and shapes, natural or synthetic origins, particulate, fibrous (long, short, oriented), flakes, etc. Other than fillers, various other additives are also incorporated into polyolefin composites, such as stabilizers, plasticizers and processing aids, antioxidants, antiozonants, UV stabilizers, coupling agents, blowing agents, flame retardants, pigments, and fungicides, etc. (Tolinski, 2015). The various types of fillers used for the preparation of polyolefin composites include various natural fibers, glass fibers, minerals of clay types, carbon black, carbon fibers, single or multiwalled carbon nanotubes, graphite, titanium dioxide, magnesium hydroxide, calcium phosphate (hydroxyapatite, HA), aluminum trihydroxide, calcium carbonate, silica, etc. (Rothon and DeArmitt, 2017). The properties of the resulting polyolefin composites mainly depend on the properties of the matrix polymers, weight or volume fractions of the reinforcements, and also on the interactions between the matrix and the reinforcements.

6.3 BIOMEDICAL ENGINEERING The branch of engineering that deals with the application of engineering concepts and design methodologies in the field of medicine and health care is known as biomedical engineering. The areas under this field of engineering deal with various clinical sectors, such as diagnosis, monitoring, and treatment of diseases. The most popular applications in the field of biomedical engineering include scaffolds and tissue engineering, design and development of biocompatible implants, different types of therapeutic, diagnostic, or monitoring devices, medical imaging devices, drugs and pharmaceutical engineering, and drug administration techniques, etc., to name a few. Polyolefins as engineering materials find applications in most of the mentioned areas either as standalone materials or as composites in combination with fibrous or particulate fillers (Ramakrishna et al., 2001).

6.4 BIOCOMPATIBILITY EVALUATION OF POLYOLEFIN-BASED BIOCOMPOSITES Biocompatibility, or the evaluation of the biological responses of the prospective biomaterial, is an important step in the development of materials and devices for biomedical engineering related applications. The main challenge in the evaluation of biocompatibility of composites is due to the fact that composites contain more

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than one distinct phase. In composites, the matrix is the continuous phase and the reinforcement is the discontinuous phase. The matrix can be in macro- or microdimensions and the reinforcement usually exist on a micro- or nanolevel and the contributions of each of these phases toward tissue/material interactions should be carefully monitored. The biocompatibility of a material or device is defined as its ability to exhibit a favorable host response during its intended application. The surface of the device or material plays a vital role in controlling the host response. For conventional types of composite materials with one matrix at the tissue/device interface, the evaluation of biocompatibility is much easier than for composites with more than one matrix and many other additives as previously mentioned. There are two major challenges involved in the evaluation of biocompatibility of polyolefin composite materials. These are related to the evaluation of biocompatibility of the composite as a whole (including the individual contributions provided by the matrix and the reinforcement phase) and the surface characterization of the composite structure (Anderson and Voskerician, 2009). Developments in the field of nanotechnology, tissue engineering, and regenerative medicine pose more challenges to the evaluation of biocompatibility of polyolefin composite structures.

6.4.1 TESTS FOR BIOCOMPATIBILITY The tests for evaluation of biocompatibility are broadly divided into three types. These are primary (level I), secondary (level II), and preclinical (level III) tests (de Moraes Porto, 2012). Level I tests are generally in vitro cytotoxicity studies. These cell culture tests are considered as primary screening tests for the material and involve the interaction of the biomaterial with a selected cell line outside of the biological environment. There are three primary cell culture assays for the evaluation of biocompatibility. These are direct contact, agar diffusion, and elution (extract dilution) (Ratner et al., 2004). The cell lines used for cytotoxicity studies include mouse fibroblasts, lymphocytes, human lymphocytes, polymorphonuclear leukocytes and mixed leukocytes, mouse macrophages, mouse embryo cells, etc. (de Moraes Porto, 2012). The selection of a cell line for testing a material intended for a specific application should be based on the type of assay, measurement endpoints, such as viability, enzymatic activity, species receptors, etc., and also on the practical experience of the investigator (Morais et al., 2010). Level II or in vivo tests include tests for irritation, intracutaneous reactivity, systemic toxicity, subchronic toxicity, genotoxicity, implantation, hemocompatibility, chronic toxicity, carcinogenicity, reproductive and developmental toxicity, biodegradation, and immune responses (Anderson, 2004). The results from in vivo tests are based on the assessment of tissues from test animals that have received the implants. The selection of these test methods depends on the type of application of the material/device and the nature or duration of body contact as described by ISO 10993. A generalized description of various tests as per ISO and FDA is given in Table 6.1.

Table 6.1 Generalized Description of Various Tests as per ISO and FDA

External communicating devices

Indirect blood path Tissue/bone/ dentin Circulating blood

Implant devices

Bone/tissue

Blood



O Evaluation required by ISO and FDA. Additional evaluation required by FDA.

O O O O O O O O O O O O O O O O O O O O O O O

O

O











O



O O O



O

O



O O

O O

O O



O O O

O

O



O O O O O



O

O O

O O

O O

O O O

O O O O O

O O

O O O O O

O O O

O

O

O

O

O

O



O O O O O

O O

Biodegradation

O



Reproductive/ Developmental Toxicity

 

Carcinogenicity





Chronic Toxicity



Hemocompatibility

Implantation

Irritation

O O O O O O O O O O O O O O O O O O O O O O O O

Genotoxicity

Breached surfaces

O O O O O O O O O O O O O O O O O O O O O O O O

Subchronic Toxicity

Mucosal membrane

A B C A B C A B C A B C A B C A B C A B C A B C

Systemic Toxicity

Skin

Contact Duration A: ,24 h; B: 24 h for 30 days; C: .30 days.

Sensitization

Body Contact Surface contacting devices

Supplementary Evaluation

Initial Evaluation

Cytotoxicity

Device Categories

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6.5 FABRICATION TECHNIQUES FOR POLYOLEFIN BIOMEDICAL COMPOSITES Polyolefin matrix composites have achieved commercial accomplishment in biomedical applications. A vital step in the preparation of biomedical polyolefin matrix composites is fabrication. Fabrication is defined as the technology of transferring raw polymer into substances in a preferred shape and size. An important step in the thermoplastic processing method is to produce flowability. In the heating stage of the processing method, the polymer molecules slip past each other and produce flow. In the cooling stage of the processing method, the polymer molecules solidify and are shaped. The phases of flowing, shaping, and hardening of thermoplastic materials can be carried out in the fabrication machine. Composites consist of reinforcement materials held in place by a matrix system. Fabricating a composite entails the process of incorporating reinforcement material in the polymer matrix with a definite orientation to provide specified characteristics for the finished product and, thereby, execute its design function. The growth of polyolefin-based biomedical composites both in a number of applications and in volume is related to the ease of their processability. While high processing temperatures, high melting viscosity, lack of drape, etc., cause complications during the polyolefin composite fabrication process. Polyolefin composites are stronger than thermosets composites with the contribution of their amorphous and semicrystalline structures. The amorphous portion has a number of entangled chains and provides enhanced toughness to the composites. The semicrystalline portion contributes to the plasticizing consequence of the materials. Polyolefin composites offer a longer shelf life and a shorter processing time. Polyolefin high molecular weight characteristics result in high processing temperatures and high melting viscosity during processing. The shear thinning flow behavior of polyolefin materials prevents easy fiber wet out and causes complexity during fabrication. Conventional processing techniques are modified to the fabrication of thermoplastic biomedical composites. Polyolefinbased composites are shaped into various biomedical devices through processing methods such as injection molding, compression molding, extrusion, electrospinning, etc. Parameters like softening temperature, size, and shape are considered when choosing a processing method. In general, biomedical composites are fabricated in a clean room to diminish the incorporation of inappropriate substances into the products (Olabisi and Adewale, 2016; White and Choi, 2005a,b; Biron, 2012; Nwabunma and Kyu, 2008; Ambrosio, 2009). A schematic representation of processing techniques for polyolefin-based composites is shown in Fig. 6.1. This section will explore modified techniques used in the manufacturing of polyolefin composites. The basic processing steps as well as the advantages and disadvantages of each processing method are discussed.

6.5 Fabrication Techniques for Polyolefin Biomedical Composites

FIGURE 6.1 Flowchart of processing techniques for polyolefin-based composites.

6.5.1 MOLDING A mold is a hollow structure that imparts its shape to finished products. The term “molding” includes injection, compression, blow, and rotational molding processes. Injection molding is the most commonly used processing method used for the manufacturing of polyolefin biomedical composites. Injection molding is a rapid technique to produce products by injecting polymer materials into a mold. In the injection molding process, thermoplastic filled materials are melted in a chamber to a temperature that leads the materials to flow and they are then pumped into a closed mold through a rotating screw. The polymer melt solidifies on cooling and the mold is opened for the removal of the finished product. The injection molding technique provides excellent product consistency, little postproduction scrap, good dimensional control, high production rate, etc. The disadvantages of the injection molding process include the requirement of expensive equipment investment and product design restrictions, etc. (Bryce, 1996; Olmsted and Davis, 2001; Rosato and Rosato, 2012). Short-fiber polyolefin composites are injection molded, however, fiber orientation and distribution are difficult to control. Haneef Mohammed et al. prepared HDPE composite materials for use in orthopedic applications with hybrid reinforcement materials, such as titanium dioxide and alumina particles using an injection molding machine (Haneef et al., 2013). In the compression molding process, compounded pellets are placed between stationary and movable molds and then pressure and heat are applied to achieve a consistently shaped composite. The compression molding technique offers the least expensive tooling, large part production, little material wastes, etc. The disadvantages of the compression molding process include its slower process

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FIGURE 6.2 Flowchart of molding process.

time, its inappropriateness for complex product design, high labor cost, etc. Juhasz et al. developed biomedical composites for use as implant materials that consist of glass ceramic apatite wollastonite particulate matter reinforced with HDPE using a compression molding technique (Juhasz et al., 2004). All ceramic particle-reinforced polyethylene composites without chemical coupling are manufactured through compression molding (Wang et al., 1998a). Fig. 6.2 shows a flowchart of the molding process.

6.5.2 EXTRUSION Extrusion is a versatile process used for manufacturing products with a uniform cross-section, like hoses, films, sheets, etc., by forcing materials through a die under controlled conditions. The five main components of the extrusion process are a drive system, a feed system, a barrel system, die assembly, and a control system. The polymer extrusion process includes single-screw and twin-screw extruders. Melting, compression, and metering are the main steps of the extrusion technique. In the melting section, pellets are transferred from a hopper and converted into molten polymer. In the compression section, the molten polymer is compressed and mixed with the different additives. The metering section is to produce the desired product through a shape-forming die. The final products are shaped and cooled (Carley, 1989; Nakajima, 1997; Rauwendaal, 2014). Films are formed through film blowing thin-walled tubes or drawing cast films. The extrusion technique is also used for the purpose of compounding polyolefin-based biomedical composites, for example, hydroxyapatite particulate reinforced UHMWPE nanocomposite with superior toughness for orthopedic applications is compounded using twin-screw extrusion (Fang et al., 2006). Ram extrusion consists of a hopper that permits material to go into a heated cavity, a reciprocating ram, and a die unit. UHMWPE-based biomedical composite products can be manufacture using a ram extrusion process (Kurtz, 2004). In hydrostatic extrusion, the workpiece is fully enclosed by pressurized liquid. When the ram moves forward, the delivered force pressurizes the liquid and applies pressure to all surfaces of

6.5 Fabrication Techniques for Polyolefin Biomedical Composites

FIGURE 6.3 Schematic representation of hydrostatic extrusion process.

the workpiece that pushes the work through the die. Fig. 6.3 shows a schematic representation of the hydrostatic extrusion process. Hydrostatic extrusion is useful for enhancing the mechanical properties of hydroxyapatite/HDPE composites for load bearing implant applications. Polyethylene chains can align in the extrusion direction to produce high strength and high modulus parts (Wang et al., 1998b; Ladizesky et al., 1997).

6.5.3 MELT ELECTROSPINNING Electrospinning is a technique used to produce fibers with a diameter less than 100 nm. Electrospinning includes a high voltage power supply, grounded collector, and positively charged capillary packed with polymer fluid. In this process, a high voltage electric field is applied to form fibers from polymer fluid that is conveyed through a capillary. A liquid polymer jet is formed and elongated under the action of electrostatic repulsion and finally deposited on a grounded collector which serves as an electrode. Electrospun fibers have great prospective in biomedical applications, such as in scaffolds, drug delivery, wound dressing, implants, etc. Melt electrospinning is an interesting alternative technique to conventional solvent electrospinning (Ramakrishna, 2005). In the melt electrospinning process, the polymer melts, which is normally more viscous when polymer solution is used. The molten electrified jet requires cooling for solidifying and results in micron diameter fibers. Melt electrospinning is suitable for nonsoluble polymers, like polyolefin, and thereby ensures that no solvent is present in the final fiber. Organic solvents that lead to cellular cytotoxicity are not used in this process and benefit from the

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FIGURE 6.4 Schematic representation of melt electrospinning setup.

applications in biomedical sectors (Ferrari et al., 2007; Agarwal et al., 2008). Melt electrospun fibers show excellent properties, like large surface area, outstanding mechanical performance, high length/diameter ratio, etc. Fig. 6.4 shows a schematic representation of a melt electrospinning setup.

6.5.4 FILAMENT WINDING Filament winding is a process that involves winding strands under tension over a male mandrel, which is suitable for producing open or closed-end structures. The filament winding technique is used to produce parts with excellent repeatability, high performance, quality internal surface, high fiber-to-resin ratios, minimum labor involvement, etc. The disadvantages of this process include the requirement for a high initial investment and the inability to generate reverse curvature, etc. In the thermoplastic filament winding process, fiber strands unwind and continuously pass a thermoplastic resin tank. In the resin tank, the fiber strands are completely impregnated with thermoplastic resin. These resin impregnated strands are wound around the mandrel in a controlled manner and in a definite fiber orientation (Hoa, 2009; Strong, 2008). Kazanci et al. developed a butane ethylene copolymer reinforced with UHMWPE fibers using a filament winding technique, and these composites were potentially intended for ligament or tendon prostheses (Kazanci et al., 2002a,b). Fig. 6.5 shows a representation of fiber wound on a mandrel in a filament winding method.

6.5 Fabrication Techniques for Polyolefin Biomedical Composites

FIGURE 6.5 Representation of fiber wound on a mandrel in filament winding method.

FIGURE 6.6 Flow diagram represents the steps in the thermoplastic pultrusion process.

6.5.5 THERMOPLASTIC PULTRUSION Pultrusion is a continuous and cost-effective processing technique used for the production of composites with close-dimensional cross-sections. Pultrusion is an ideal process for the manufacturing of either solid or hollow profile-like flat bars, channels, pipes, tubing, rods, etc. (Hoa, 2009; Strong, 2008). In the thermoplastic pultrusion method, preheated continuous fiber strands are pulled into the impregnated apparatus in order to wet out the fibers. Then the melt impregnated reinforcements are passed through a cooling die, which controls the shape, size, and finish of the finished products. A puller is used to control the speed of the process and generate a dragging force on the products. Finally, a pelletizing system is used to cut the final products (Tao et al., 2015). The pultrusion technique is wellsuited for parts that need good dimensional tolerance, high fiber volume fractions, excellent reinforcement alignment, precise control of resin and fiber, low scrap rates, etc. Some of the limitations of this process are the requirement for a high initial investment and for skilled labor, etc. Fig. 6.6 shows the steps involved in the thermoplastic pultrusion process.

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6.6 POLYETHYLENE MATRIX Polyethylene is produced from ethylene monomers having a molecular weight of 28. The chemical formula of polyethylene is (C2H4)n , where n is the number of repeat units. A representation of the chemical structures of ethylene and polyethylene is shown in Fig. 6.7. Polyethylene is available in several forms, such as LDPE, LLDPE, HDPE, and UHMWPE, which are produced with different molecular weights and chain structural designs (Peacock, 2000). Schematics of the molecule alignment in different forms of polyethylene are shown in Fig. 6.8.

FIGURE 6.7 Representation of the chemical structures of ethylene and polyethylene.

FIGURE 6.8 Schematics of molecule alignment in different forms of polyethylene.

6.6 Polyethylene Matrix

LDPE has a branched chain structural design and LLDPE has a linear chain structural design with a molecular weight of less than 50,000 g mol21. HDPE has linear chain architecture with a molecular weight of 200,000 g mol21 and a crystallinity of 60% 80%. HDPE resin is naturally translucent and offers good low-temperature toughness, creep resistance, moisture barrier, impact resistance, chemicals resistance, etc. The molecular weight of UHMWPE cannot be calculated by conventional techniques, however, it is inferred by its intrinsic viscosity. UHMWPE generally has a viscosity average molecular weight of 6 million g mol21 with a crystallinity of 50% 60%. UHMWPE is an odorless, tasteless, and nontoxic material made up of extremely long chains of polyethylene that all align in the same direction. UHMWPE resin provides high ultimate strength, good impact strength, excellent wear resistance, resistance to corrosive chemicals, extremely low moisture absorption, low coefficient of friction, etc. (Kurtz, 2004; Brydson, 1999). The significant mechanical properties of HDPE and UHMWPE compared to other polyethylenes contribute to the development of their composites as matrices or reinforcements for various biomedical applications, such as knee/hip/shoulder joints. Typical average properties of different forms of polyethylenes are given in Table 6.2.

6.6.1 HDPE-BASED BIOMEDICAL COMPOSITES HDPE material is chosen for various biomedical and technical applications due to its excellent chemical and creep resistance characteristics. Bones and joints made up of natural composite materials are often fractured due to diseases, impact stress, traumatic situations, etc., and thereby need to be temporarily or permanently restored. The search for a bone replacement material that combines the mechanical and biological necessities has been extensively carried out in ceramics, metals, and polymers (Ambrosio, 2009). Composites of HDPE reinforced with HA have been considered as materials for bone replacement, middle ear prostheses, and orbital floor implants without any inflammatory response (Hule and Pochan, 2007; Zhang et al., 2007). The concept of fabricating bioactive composites for bone replacements was developed in the early 1980s by William Bonfield and coworkers with bioactive HA-reinforced HDPE composites (Bonfield et al., 1981; Wang and Porter, 1994). HDPE reinforced with 40 vol.% of hydroxyapatite particles was commercialized with the trade name HAPEX. Hydroxyapatite particles within composite materials were observed to perform as microanchors, providing encouraging cell attachment sites (Huang et al., 1997). An HA content of 40 vol.% concentration in HDPE with controlled surface topology provides greater bioactivity with enhanced osteoblast proliferation, skeletal tissue restoration, and durability of implants (Di Silvio et al., 1998, 2002; Dalby et al., 2000, 2002). Zhang et al. carried out in vitro biocompatibility of 30% of HA/HDPE composites with human osteoblast cells from femoral heads and skull bones and observed normal growth cycles on the composites. Rough surface HDPE/HA composite implants offer superior cellular response compared to

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Table 6.2 Typical Average Properties of Different Forms of Polyethylene Property

LDPE

LLDPE

HDPE

UHMWPE

0.910 0.925 ,0.01

0.92 0.94 ,0.01

0.941 0.965 ,0.01

0.928 0.94 ,0.01

16 90 800 0.20 0.40 0.08 0.60 41

30 500 0.30 0.70 0.60 1 45

38 20 1000 0.60 1.40 1 2 31

41 200 500 0.7 0.8 1 1.7 No break

81 97 Shore A

44 70 Shore D

33 66 Rockwell R

66 Shore D

105 118 -

122 124 -110

126 135 -110

130 135 -110

32

35

54

48

5.6 12.2

8

6.1 7.2

7.8

460 700

600

450 500

900

2.25 2.35 0.0002 10 15

2.36 0.0002 10 15

2.30 2.35 0.0003 10 15

2.30 2.35 0.0002 10 18

135 160

200 230

200 250

250 350

1.51 4 50

68 92

1.54 10 50

-

Physical Density (g cm23) Water absorption, 24 h, 1/8 in. thick (%) Mechanical Tensile strength (MPa) Elongation at break (%) Tensile modulus (GPa) Flexural modulus (GPa) Izod impact, notched (kJ m22) Hardness Thermal Melting temperature (Tm) ( C) Glass transition temperature (Tg) ( C) Deflection temperature at 1.8 MPa ( C) Coefficient of linear thermal expansion ( 3 1025 in. (in.  F)21 Electrical Dielectric strength (V mil21) short time, 1/8 in. thick Dielectric constant at 1 kHz Dissipation factor at 1 kHz Volume resistivity (ohm-cm) at 73 F, 50% RH Arc resistance (sec.) Optical Refractive index Transmission, visible (%)

smooth-surfaced implants, and their adequate impact characteristics make them a likely contender for skull implants. (Zhang et al., 2007). Wang et al. investigated the effects of HA particle size on composite properties and found that HA particle-reinforced composites with smaller sized HA particles had higher tensile strength and tensile modulus and lower strain to failure (Wang et al., 1998b). The

6.6 Polyethylene Matrix

modulus of HDPE/HA composites closely matches to bone, thus, providing a practical solution for the bone resorption problem (Wang and Bonfield, 2001). Zhang and Tanner studied the impact properties of HDPE/HA composites and found that the impact strength of HDPE/HA composites decreased with increasing HA content due to weak filler-matrix interfaces resulting in creation of voids and the propagation of cracks (Zhang and Tanner, 2003). The uniaxial and biaxial fatigue behaviors of 40 vol.% HA-reinforced HDPE composites were investigated by That et al. (That et al., 2000a,b). The uniaxial fatigue test results found that the ultimate strengths were uppermost in compression and torsion. The composites were ductile in torsion, but more brittle in tension. The weaker filler/matrix resin interface was liable for the failure in torsional fatigue. The biaxial fatigue was dominated by a shear mechanism and the test results revealed that out-of-phase loading was less destructive than in-phase loading, while failure was not observed at 25% of ultimate tensile strength and ultimate shear stress. The incorporation of a compatibilizing agent and a surface treatment of HA favor the dispersion, distribution, and interfacial interaction of the filler particles in a polymer, which enhances the mechanical properties of HDPE/HA composites. Wang et al. reported the enhancement of HDPE/HA composite properties through mechanical interlocking at the matrix-reinforcement interface by the incorporation of acrylic acid grafted HDPE as a compatibilizer and through surface treatment of HA with silane coupling agents (Wang et al., 2000; Wang and Bonfield, 2001). Balakrishnan et al. explored the use of maleated high-density polyethylene (mHDPE) as a compatiblizing agent for HDPE/HA composites for bone replacement applications. The addition of mHDPE as a compatibilizer enhanced the interfacial adhesion between HDPE and HA particles through the formation of HDPE fibrils and this fibril network was interconnected with the HA particles. Albano et al. investigated composites based on HDPE reinforced with surface treated HA (Albano et al., 2009). The surface treatment was carried out with an ethylene acrylic acid copolymer and acrylic acid. An in vitro study of the composites with the ethylene acrylic acid copolymer treated HA showed improved cell adhesion. The results of the surface treated HA composites were credited to the interactions between the ethylene acrylic acid copolymer and the HA. Joseph et al. studied the effect of an HDPE matrix on the rheological behavior of HA filled composites (Joseph et al., 2001a,b, 2002). The incorporation of HA powder into the HDPE matrix increased the plateau value and reduced the melt compressibility. Carmen et al. found that a zirconate coupling agent was better executed than a titanate coupling agent in HDPE/HA composites based on its mechanical properties. Zirconate as a coupling agent has a richer electronic density than a titanate coupling agent, which means the elevated electronic richness of the metal-oxygen bonds contributed to the enhanced interactions between the polymer and the reinforcement (Carmen et al., 2007). The formation of an apatite layer, a “seaweed-like” structure, was observed with increasing stimulated body fluid (SBF) immersion time on the outside of HDPE/HA composites with mHDPE as a compatibilizer. The enhanced growth of

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apatite layers indicates the excellent biocompatibility properties of the composites and highlight its potential for use in bone implantation (Balakrishnan et al., 2013). Apatite layer formation on the outside of HDPE/HA composites in SBF exhibits in vitro bioactivity behavior and are capable of tying to living bone face in the human body (Kokubo and Takadama, 2006; Rea et al., 2004). The development of an apatite layer is an effect from the ion exchange process between ions (Ca21 and HPO422) from SBF and HA during immersion (Cao and Hench, 1996). The growth of apatite layers is induced by HA and is associated with the size of the HA particles. The smaller particle size of HA offers greater surface area to the SBF solution and induces the increased growth of apatite layers (LeGeros, 2008; Fang et al., 2006; Espigares et al., 2002). HA/HDPE composite material is used in orbital floor replacement devices and has accomplished the necessary requirements, like the ability to attach well to the orbital floor and exhibiting an unchanged volume of implant (Downes et al., 1991; Tanner et al., 1994). In the first design, the HA/HDPE composite material was compression molded as a disc and used to seal the bottom of the eye socket following rupture of the orbital floor, thus, avoiding the extrusion of the soft tissues into the sinus spaces. In the second design, for patients who had lost an eye, HA/HDPE composite material can be used as a space-filling implant (Ambrosio, 2009). An HA/HDPE composite was used to replace UHMWPE in middle ear implants as shafts which were cut into the necessary lengths to fit on staples that convey and strengthen sound vibrations from the outer ear to the inner ear. HAPEX shafts increase the long-term stability of implants, make it easier to trim intraoperatively due to the presence of the HA particles, and increase the sound transfer through the increased density of the composites (Dornhoffer, 1998; Goldenberg, 1994; Goldenberg and Driver, 2000). HA powder reinforced HDPE composites were compression molded into preferred plates and irradiated with different doses by Smolko and Romero (Smolko and Romero, 2007). The incorporation of HA powder increased the young’s modulus and tensile strength at break, whereas it decreased the yield strength and elongation. However, the increase in dose of radiation enhanced the tensile and yield strength and reduced the elongation from 800% to 5%. HDPE/nanoparticle hydroxyapatite composites with 10 30 wt.% of fillers were prepared by Fouad et al. (2013). It was observed that increasing the content of reinforcement decreases the melting temperature and the crystallinity of composites due to the restriction of the mobility of molecules. However, with ageing, HDPE-based nanohydroxyapatite composites showed a decrease in melting temperature and crystallinity increase due to chain scission and oxidation. The addition of HA nanoparticles improved the hardness and wear resistance of the composites. Li and Tjong prepared HDPE/HA nanorod composites, fabricated by melt compounding with 20 wt.% of reinforcement (Li and Tjong, 2011). The effective incorporation of hydroxyapatite nanorods (HANRs) enhanced hardness, young’s modulus, yield strength, thermal stability, and wear resistance compared to pure HDPE significantly. HA and alumina restricted to a total of 40 vol.%

6.6 Polyethylene Matrix

were added into an HDPE matrix by Nath et al. (2009). It was found that an elevated elastic modulus, superior hardness, and higher wear resistance were obtained in a HDPE/20 vol.% HA/20 vol.% alumina composite. In vitro cell culture studies confirmed the constructive cell adhesion properties in the prepared hybrid composites. The wear resistance of HDPE/GNP composites was examined using pin-on-disc wear testing equipment under various sliding velocities by Liu et al. (2014). A wear resistance enhancement of about 97% under 1.3 m s21 sliding velocity was realized in silanized GNP-reinforced HDPE composites. Haneef et al. developed hybrid polymer matrix composites using HDPE as a matrix material with titanium oxide and alumina particles as fillers. It was found that the hybrid polymer matrix composite offered superior mechanical and tribological characteristics, which are required for bone substitution materials. HDPE with 10 wt.% titanium oxide and with various proportions of alumina hybrid composites (5%, 10%, 15%, and 20%) were prepared and subjected to mechanical and tribological characterization The study reported that overall better mechanical and tribological properties were attained with a 10% titanium oxide and 20% alumina reinforced HDPE hybrid composite (Haneef et al., 2013). Spinal diseases and loads experienced by the spine through daily activities are common problems that affect the intervertebral discs (IVDs) and other spinal parts (Martz et al., 1997). Many manmade devices are developed to restore spinal stability and function. IVDs offer spine flexibility and provide an extensive variety of postures to the body. Each IVD has soft “nucleus pulposus” surrounded by “annulus fibrosus.” IVD degeneration causes back pain due to dehydration of the nucleus associated with tiny tears in the annulus (Hukins, 2005; Tsantrizos et al., 2005). Various IVD prostheses has been developed with materials like metals, polymers, and ceramics considering their properties, such as biocompatibility, resistance to compressive creep, and endurance (Bao and Yuan, 2000; De Santis et al., 2000; Ramakrishna et al., 2001; Traynelis, 2002; Gloria et al., 2008; Shikinami et al., 2004). Poly(2-hydroxyethylmethacrylate) (PHEMA) hydrogels were used in IVD prosthesis, however, the mechanical properties of PHEMA in hydrated conditions are not appropriate for biomedical applications with a necessity for high mechanical strength (Netti et al., 1993; Ambrosio et al., 1998; Peppas et al., 2000; Hoffman, 2012). The incorporation of hydrophobic components, like poly(caprolactone) (PCL) and polymeric fibers improved the mechanical properties of polymer hydrogels (Davis et al., 1992; Ambrosio et al., 1998; De Santis et al., 2004). Ambrosio et al. found that PHEMA/poly(methylmethacrylate) (PMMA) semi-interpenetrating polymer networks with poly(ethylene terephthalate) (PET) fibers show considerable potential for use in IVD replacements with fabrication by filament winding and molding processes (Ambrosio et al., 1998; Ambrosio, 2009). The stiff and hard metal endplates that cause bone resorption are replaced with HA reinforced polyethylene composites, which help to fasten the device to the vertebral bodies (De Santis et al., 2004; Ambrosio et al., 2007; Gloria et al., 2007). The addition of bioactive materials, like HA and/or calcium phosphate, is beneficial and provides the stiffness required for the endplates (Giordano et al., 2006).

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FIGURE 6.9 IVD prosthesis consisting of PHEMA/PMMA semi-interpenetrating polymer networks hydrogel reinforced with PET fibers and two hydroxyapatite reinforced polyethylene composite endplates (Ambrosio, 2009).

Fig. 6.9 shows composite biomimetic total IVD prosthesis made up of PHEMA/ PMMA semi-interpenetrating polymer network hydrogels reinforced with PET fibers and two HA reinforced polyethylene composite endplates. The mechanical and tribological properties of HDPE/MWCNT (multiwall carbon nanotube) composites were investigated by Kanagaraj et al. (2007). The composites showed improvements in mechanical properties with an increase of carbon nanotube (CNT) content as well as good load transfer effect and interface link between reinforcement and polymer. Theoretical calculation of volumetric wear rate was estimated by Ratner’s correlation, Wang’s model, and reciprocal toughness. It was observed that the volumetric wear rate of the composites decreased with the incorporation of CNTs. The reciprocal toughness and volumetric wear rate showed a linear relationship that maintains the theoretical calculation of the microscopic wear model. Composites with bioactive reinforcement and ductile polymer matrices were popular for implant applications due to their superior mechanical and biological properties. Glass-ceramic apatite wollastonite (A W) is a bioactive ceramic material that helps bone regeneration and offers strong interfacial linking between the implant and host tissue in biomedical applications (Yamamuro, 1993, 1995). Juhasz et al. studied the effect of filler content and particle size on the mechanical properties of glass-ceramic A W particulate reinforced HDPE composites. The manufacturing process of these composites involved blending and compounding through a twin-screw extruder followed by centrifugal milling and compression molding. HDPE-based composites consisting of glass-ceramic A W with volume fractions varying from 10% to 50% with average particle sizes of 4.4 and 6.7 mm were prepared and compared with HAPEX, a commercially available composite of HA and HDPE, with 40 vol.% filler content. It was observed that raising the glass-ceramic A W content

6.6 Polyethylene Matrix

increases the young’s modulus, yield strength, bending strength, and decreases the strain to failure behavior, but the young’s modulus, yield strength, and bending strengths were found to be slightly reduced with increasing filler particle size. In general, implant materials should preferably exhibit the characteristic of ductile behavior with a high strain to failure in order to evade any disastrous failure in the body. The results showed that glass-ceramic A W particulate reinforced HDPE composites with 50 vol.% have the potential for implant applications (Juhasz et al., 2004).

6.6.2 UHMWPE-BASED BIOMEDICAL COMPOSITES UHMWPE shows a low friction coefficient against steel, superior wear resistance, and good impact strength, which are essential for biomedical applications. It’s excellent load bearing characteristics make it useful in joint endoprostheses in combination with metal or ceramic fillers. A UHMWPE composite reinforced with carbon fibers was developed and used in orthopedic implants. UHMWPEbased biomedical composites reinforced with chopped carbon fibers with random orientation were manufactured through a compression molding technique (Kurtz, 2004). Ainsworth et al. found that carbon fiber-reinforced biomedical composites have lower wear rates, greater stiffness, flexural yield strength, and elastic modulus with the capability to endure higher compressive loads (Ainsworth et al., 1977). However, the poor fiber-matrix adhesions in carbon fiber-reinforced UHMWPE composites contributed to short-term clinical failures and they were ultimately neglected for use in joint replacement. Interestingly, studies at Drexel University revealed that long-term implanted carbon fiber-reinforced UHMWPE composites showed well-consolidated fibers in the matrix (Kurtz, 2004). The observed results from the long-term clinical test create an impulse to reexamine carbon fiber-reinforced UHMWPE biomedical composites for hip and knee joint replacement. UHMWPE homocomposites used in joint replacement applications were manufactured by sintering oriented UHMWPE fibers together or through the reinforcement of polymer matrices with UHMWPE fibers. Self-reinforced UHMWPE materials offer elevated tensile strength, tensile modulus, and abrasion resistance with respect to UHMWPE bulk materials (Price et al., 1997). The arrival of crosslinked UHMWPE and challenges encountered during the processing of the UHMWPE homocomposites pose difficulties for the commercialization of self-reinforced UHMWPE for orthopedic applications (Chang et al., 2000). UHMWPE matrix composites reinforced with nanoparticles, nanotubes, and nanofibers are used in orthopedic bearing applications. UHMWPE material filled with Al Cu Fe powders and chemically crosslinked UHMWPE reinforced with quartz particles of micron size were developed and are at present in the experimental analysis phase for orthopedic applications (Anderson et al., 2002; Xie et al., 2003; Liu et al., 1999). The fibrous reinforcement in composite materials improved the yield strength and elastic modulus of the surface treatment of the filler materials by plasma or chemical etching, usually, prevent filler

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aggregation issue and thereby improve reinforcement-matrix adhesion (Hofste et al., 1998). Anderson et al. observed that Al Cu Fe/UHMWPE composites have better wear resistance to volume loss as compared to pure UHMWPE and alumina/ UHMWPE composites. The volume loss of pure UHMWPE was due to the removal of the polymer materials during wear. Al Cu Fe/UHMWPE composites showed a 35% decrease in volume loss as compared to pure UHMWPE. The enhanced wear resistance of Al Cu Fe/UHMWPE composites has been recognized in the high young’s modulus, high hardness, and low coefficient of friction of the composites (Anderson et al., 2002). UHMWPE-based hybrid composites reinforced with bioactive HA, bioinert aluminum oxide, and CNTs were fabricated using compression molding by Gupta et al. The hybrid composites exhibited higher elastic modulus and toughness compared to that of pure UHMWPE (Gupta et al., 2013). The abrasive wear performance of UHMWPE reinforced with quartz powder was studied (Liu et al., 1999). The quartz powder offered increased surface hardness and enhanced ploughing and cutting resistance to the composite. The incorporation of quartz filler in UHMWPE reduced deep furrow formation and improved wear resistance by about a maximum of four times compared to unfilled UHMWPE. Larger filler particles are superior to smaller particles for the improvement of wear resistance. UHMWPE/quartz composites with vinyl triethyloxyl silane were prepared (Xie et al., 2003). Vinyl tri-ethyloxyl silane acts as a crosslinking agent for the UHMWPE matrix. The crosslinking of UHMWPE leads to an enhancement of wear resistance in the composites. The mechanical and wear resistance properties of UHMWPE/quartz composites reached a maximum at 0.5 phr of vinyl tri-ethyloxyl silane. Disease conditions, like osteoarthritis and degenerative joint disease, that cause the breakdown of cartilage in the joints are the main reason for having a hip or knee replacement. Total joint substitution helps to treat these debilitating diseases and improve the quality of life pf patients (Ambrosio, 2009). Two important components of prosthesis for total hip substitution are the femoral and acetabular parts. Femoral parts are made up of Co-Cr alloy, Ti alloy, aluminum, or zirconium. Typically, UMHWPE material is used to fabricate the acetabular part. When compared with the hip joint, the knee joint is considered as more complicated in geometry as well as movement mechanics. In general, the prosthesis of total knee joint substitution has femoral, tibial, and/or patellar components. The tibial and patella surfaces are made up of UHMWPE. The femur end commonly uses metal implants (Ambrosio, 2009; Davim, 2012). Fig. 6.10 shows the prosthesis for total hip/knee joint replacement. UHMWPE fibers embedded in an ethylene butene copolymer matrix composite material was found as a potential candidate for tendon and ligament prostheses (Kazanci et al., 2001; Kazanci et al., 2002a,b; Ratner et al., 2005). Kazanci et al. studied the fatigue test under cyclic loading for a filament wound flat strip of this composite material. Fatigue tests were carried out at room temperature under tension tension sinusoidal load at R 5 0.1 and a frequency of 1 Hz. In the fatigue

6.6 Polyethylene Matrix

FIGURE 6.10 Prosthesis for total hip/knee joint replacement (Davim, 2012).

test, three different copolymer compositions and two different winding angles were used to study the effects of branching density in the matrix and reinforcement angle on the fatigue response of the composites. The short-term fatigue test at high-stress levels exhibited improved fatigue resistance for a copolymer with a lower branching density and smaller reinforcement angle. However, the long-term fatigue test at moderate stress levels was managed by the fatigue rate of degradation, which reduced with branching density and winding angle (Kazanci et al., 2002b). Tribological characterization of biomedical composites plays a vital role in the existence of orthopedic implants in total joint replacements. Kalin et al. studied the tribological characterization of a UHMWPE-based composite with 5%, 10%, and 30% HA reinforcement under different loading conditions. It was revealed that the wear resistance of the UHMWPE/HA composites enhanced with increasing reinforcement content (Kalin et al., 2002; Ambrosio, 2009). Fang et al. manufactured UHMWPE/HA composites by mechanically mixing nanosized HA with UHMWPE in a ball mill and then compression molded the composite into solid blocks. The prepared blocks were allowed to swollen in paraffin oil to improve the UHMWPE chain mobility and reinforcement/matrix interface adhesion before ending hot-press. It was found that ball milling and swelling treatment enhanced the mechanical properties of the HA/UHMWPE composites compared to that of pure UHMWPE (Fang et al., 2005). Mirsalehi et al. studied the effect of the addition of 10 50 wt.% HA nanoparticles on the mechanical properties of UHMWPE, which was synthesized via a sol gel method. It was observed that composites with 50 wt.% exhibited higher young’s modulus and yield strength than that of pure UHMWPE (Mirsalehi et al., 2016). Bovine bone hydroxyapatite (BHA) reinforced UHMWPE composites were developed using a heat pressing formation technique. The incorporation of BHA filler particles enhanced the hardness and creep modulus of the composites as well as the biotribological properties

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for biomedical applications (Wang et al., 2009). UHMWPE/nanohydroxyapatite composites were prepared by vacuum hot-pressing and the composites were irradiated with gamma rays in a vacuum and then molten heat treated in a vacuum immediately after irradiation. The friction coefficient and wear rate in deionized water lubrication were diminished by the addition of nanohydroxyapatite (nHA). Whereas, the friction coefficient was improved and wear rate was decreased by gamma irradiation. The results showed that UHMWPE/nHA composites irradiated at a dose of 150 kGy could be used as bearing materials in artificial joints due to its superior wear resistance over that of pure UHMWPE (Xiong et al., 2009). The enhanced tribological properties in polyamide-6(PA-6)/UHMWPE composite can be achieved by incorporation of tiny UHMWPE particles, which play the role of lubricating the agent as well as PA-6 prevent the UHMWPE particles being transferred into the counterpart. Liu et al. studied the effect of contact pressure, sliding distance, and sliding speed on the wear properties of a polyamide-6 (PA-6)/UHMWPE composite. Contact pressure was found to be the main significant parameter in the wear rate of the PA-6/UHMWPE composite, followed by sliding distance and sliding speed (Liu et al., 2001). Liu et al. investigated the effect of contact pressure on the wear rate of PA-6, UHMWPE, and PA-6/ UHMWPE under dry and lubricated sliding conditions. The contact pressure as well as lubricating condition, have strong influences on the wear loss of the materials. The materials showed greater wear loss under elevated contact pressure under the dry sliding condition compared to that of the lubricated condition. The study observed that UHMWPE established the maximum dry sliding wear rate, under both 1 and 2.5 MPa of contact pressure, followed by PA-6/UHMWPE and PA-6, but PA-6 showed the uppermost wear rate in the lubricated sliding condition, followed by UHMWPE and PA-6/UHMWPE (Liu et al., 2006). Dangsheng studied the influences of carbon fiber content on the tribological characterization of UHMWPE in different concentrations, which is useful as an artificial joint acetabular material. The hardness of the carbon fiber-reinforced UHMWPE composites increased with reinforcement content. The composites wear volume loss was reduced with carbon fiber content under dry and distilled water lubricating conditions. The friction coefficients of the carbon fiberreinforced UHMWPE composites were greater than that of pure UHMWPE under the dry sliding condition and the friction coefficients were lower than that of pure UHMWPE under the distilled water lubrication condition (Dangsheng, 2005). In a study of the tribological characterization of UHMWPE/MWCNT composites, Kanagaraj et al. concluded that the wear resistance of UHMWPE was enhanced by the incorporation of CNTs. The addition of CNTs decreased the wear volume and wear coefficient of the composites due to the improved interfacial strength between the CNTs and the UHMWPE and superior load transfer effect to the CNTs from the UHMWPE. The wear coefficient of the composites decreased with an enhancement of the sliding distance in a linear model (Kanagaraj et al., 2010). A composite consisting of 80% UHMWPE and 20% HDPE reinforced with

6.6 Polyethylene Matrix

MWCNTs varying between 0.2 and 2 wt.% was studied by Xue et al. (2006). The composites showed a significant decrease in wear rate with an increase of both pretreated CNTs in boiling nitric acid and untreated CNT content. The incorporation of 0.5 wt.% of CNTs into a UHMWPE/HDPE blend caused a 50% reduction in the wear rate. CNT-reinforced UHMWPE/HDPE blends offered excellent wear properties compared to UHMWPE/HDPE blends without fillers and UHMWPE alone. The composites with untreated CNTs showed better wear performance compared to that of composites with pretreated CNTS due to the higher creep resistance characteristics of composites with untreated CNTs. The influences of applied pressure on the wear volume of PP, UHMWPE, and PP/UHMWPE blends were studied by Hashmi et al. (2001). Maximum wear in PP and minimum wear in UHMWPE were noticed. The wear rate of PP was more susceptible to pressure than UHMWPE. The frictional heat amplified temperature of the contact surface with sliding distance and makes PP softer, and thereby distorts and may loose its structural integrity. The incorporation of a small weight fraction of UHMWPE in PP enhanced the wear resistance of the PP to a considerable amount and controlled the rise in temperature at the interfacial region by reducing the frictional heat. Liu et al. studied the antiwear characterization and wear mechanisms of UHMWPE and UHMWPE/PP. It was observed that the antiwear properties of UHMWPE were improved by the addition of PP. The friction coefficient and wear rate of the UHMWPE/PP blend was much inferior compared to that of the pure UHMWPE during sliding. The rod-shaped debris from the UHMWPE/PP composites presented between the two contact surfaces acted as a lubricant and helped to decrease the friction coefficient and wear to a significantly lower level (Liu et al., 2004). Self-reinforced polymer composites, also referred to as single polymer composites, are used in a wide range of commercial applications (Alcock and Peijs, 2011; Gao et al., 2012). A homocomposite of an UHMWPE matrix and an UHMWPE reinforcing phase was manufactured and studied by Mosleh et al. (1998). In this process, UHMWPE powder was filled with short chopped UHMWPE fibers with fiber volume fractions between 25% and 75% or continuous UHMWPE fabric portion with fiber volume fraction of 60%, in a layered structure and then heated under pressure to consolidate the shape. Deng and Shalaby observed that the mechanical properties of self-reinforced UHMWPE composites were better compared to pure UHMWPE, however, the wear properties of these composites were found to be the same as pure UHMWPE (Deng and Shalaby, 1997). Chang et al. studied the wear performance of bulk-oriented and fiber-reinforced UHMWPE and reported that the failure of consolidated fiber oriented composites was due to poor interaction between the UHMWPE fibers and the bulk matrix and suggested that superior fiber-matrix interaction would improve the wear behavior of self-reinforced composites (Chang et al., 2000). Self-reinforced UHMWPE composites have been suggested for load-bearing biomedical applications. The utilization of UHMWPE homocomposites, along with

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their tribological performance in an articulation surface for knee joint prosthesis were investigated (Suh et al., 1998). UHMWPE fibers with high mechanical properties produced by gel spinning reinforced with HDPE and LDPE matrices were widely studied (Alcock and Peijs, 2011). Composite materials made up of UHMWPE fibers embedded in an ethylene butene copolymer matrix through filament winding were produced and characterized for elastic, viscoelastic, and fatigue behavior. Composites with UHMWPE fiber volume fractions of 65% have the potential to be used in biomedical applications (Kazanci et al., 2001, 2002b). Creep and wear performance evaluations of ethylene butene copolymers reinforced by UHMWPE fibers were carried out by Jacobs et al. by testing in a ball-on-prism tribometer against steel balls (Jacobs et al., 2002). It was found that the creep resistance of the pure matrix was considerably reduced with increasing branches of the copolymer. The UHMWPE fiber-reinforced copolymers revealed the same wear rate and creep resistance as the solid polymer matrix. HDPE/tricalcium phosphate/UHMWPE nanocomposites were prepared and characterized for their mechanical and biological capability as a substance for bone tissue substitution. The tensile properties of the HDPE/UHMWPE blends were affected by the incorporation of nanosized tricalcium phosphate with enhancments in yield strength, young’s modulus, and a reduction in elongation at break. The addition of tricalcium phosphate nanopowder enhanced osteoblast activity, osteoinduction, and osteoconduction processes. Biological tests showed that the composites were biocompatible and had no toxicity (Abadi et al., 2010). Corona and silane surface treatments of UHMWPE fibers enhanced the mechanical properties of dental fiber-reinforced composites due to good interfacial bond formation between the reinforcement and matrix resin. The fiber surface roughness, hardness, and elastic modulus were enhanced with 5s corona discharge, but were reduced with an additional increase in exposure time (Bahramian et al., 2015). UHMWPE-based composites with graphene oxide were prepared by liquid-phase ultrasonication dispersion and then by hot-pressing. The incorporation of graphene oxide sheets of up to 1 wt.% into UHMWPE enhanced the mechanical and biocompatibility properties, making these composites a potential candidate for artificial joints in the human body (Chen et al., 2012). Wang et al. studied the biotribological behavior of UHMWPE composites containing Ti in a hip joint simulator (Wang et al., 2007). The incorporation of titanium particles into the UHMWPE matrix showed advantages, such as enhanced wear resistance of UHMWPE under simulated body fluid (SBF-9) lubrication. UHMWPE/titanium particle composite cups showed decreased wear rates against 316L steel ball heads. An utmost drop in wear rates of 50% with 20 wt.% content of titanium particles was observed. Abrasive wear and fatigue wear were the major wear mechanisms of UHMWPE/titanium particle composites in the hip joint simulator. Titanium particles over 12 wt.% in the composites led to high wear debris sizes.

6.7 Polypropylene Matrix

6.7 POLYPROPYLENE MATRIX Polypropylene is manufactured by addition polymerization of propylene monomers. Chemically, propylene can be described as 2-methyl ethylene and has an additional CH3 group compared to ethylene. The CH3 group is important as it can be arranged in different spatial conformations in the macromolecules and thereby result in products with differing properties. Broadly, the resulting polypropylene products can be classified as isotactic polypropylene, syndiotactic polypropylene, and atactic polypropylene. In the first case, the CH3 groups are arranged on the same side of the main chain of the polymer. When the CH3 groups are symmetrically arranged on the two sides of the main chain it is termed syndiotactic. If the CH3 groups are randomly distributed in a spatial relationship to the main chain it is termed atactic. Among the three categories, atactic polypropylene has little value due to its amorphous nature even though it has a slight rubbery nature. Isotactic polypropylene has a high melting point due to its high crystallinity as well is as it being stiff. Most commercial polymers are made up of isotactic polymers about 90% 95%. The crystallinity will be fairly high, which contributes to improved softening point, stiffness, tensile strength, modulus, and hardness. Polypropylenes have higher values of Mw/Mn 5 5.6 11.9 compared to polyethylenes and typical molecular weights are depicted as Mn 5 38,000 60,000 and Mw 5 220,000 700,000. A breakthrough invention for producing polypropylene was made by Natta in 1954 (Natta and Corradini, 1967). He used a modified Ziegler process for producing high molecular weight polypropylene. The material was later commercialized under the trade name Moplen by Montecatini in 1957. Several other processes were introduced thereafter, such as the Spheripol process (1983); the Valtec process (1988); and the Himont process (1990) (Maier and Calafut, 1998). The molecular weight and molecular weight distribution of polypropylene affects its properties, especially the rheological and mechanical properties. Compared to polyethylene, polypropylene deviates sharply from Newtonian to non-Newtonian behavior as depicted by rheological investigations. The physical properties of polypropylene are also different to those of polyethylene even though they have similar structures. For example, the density of polypropylene is around 0.90 g cm23, which is lower than that of polyethylene, however, it has a higher Tg and Tm. The most important property of polypropylene enabling its use as a biomedical material is its high melting point as it allows for autoclave sterilization. Another added advantage of PP is its almost similar chemical resistance to PE, which is essential for any biomedical application. However, PP is more susceptible to oxidation, chemical degradation, and crosslinking (irradiation, violet light, and other physical means) than polyethylene. Other properties, such as better creep resistance and higher environmental stress cracking resistance compared to polyethylene are useful in biomedical applications. The typical structure of polypropylene is given in Fig. 6.11. The inherent properties of polypropylene are given in Table 6.3.

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Polypropylene (PP)

[CH2–CH(CH3)]

H C

Isotactic Syndiotactic atactic

H H C H H C H

H C CH3 H C CH3 H C CH3

H C H H C H H C H

H C CH3 CH3 C H CH3 C H

H C H H C H H C CH3

FIGURE 6.11 Typical structure of polypropylene.

Polypropylene can be processed by injection molding, extrusion, blow molding, compression molding, and thermoforming techniques. Therefore, it can be converted into different shapes, such as blown film, flat film, sheets, tubes, packaging films, tapes, etc., which is highly advantageous for biomedical applications. Another advantage of PP processing is that no predrying is necessary except for hygroscopic additives. However, stabilizers and antioxidants are needed for specific purposes. It has an exceptionally high flex rate, excellent wear resistance, high temperature resistance, and low cost. Fiber applications, such as suture, braided ligament, skin and abdominal patches, and sewing rings are worth mentioning.

6.7.1 FINGER JOINT IMPLANTS The incorporation of PP into silicone rubber in order to improve its properties for use in finger joint replacements has been reported by Ziraki et al. (2016). Silicone rubber has unique properties; however, it often fails due to its poor mechanical properties. In order to improve the mechanical properties, silica nanoparticles and PP fibers were incorporated. The tensile properties showed an improvement from 5.6 to 6.21 MPa in comparison with a 2 wt.% silica/silicone composite. A drop in strength in SBF was lower when incorporated with PP fibers rather than with nanoparticles alone. Voids and fiber degradation were noted after the composites were soaked in SBF (Fig. 6.12).

6.7 Polypropylene Matrix

Table 6.3 Properties of Polypropylene Unit

Value

g cm23 %

0.90 0.915 0.01 0.035 1.47 1.51 16.3

Physical Properties Density Water absorption Refractive index, nD20 Solubility parameter

MPa1/2

Mechanical Properties Bulk modulus Tensile strength Elongation at break Young's modulus Fracture toughness Hardness Compressive strength Poisson's ratio Shear modulus

GPa MPa % GPa MPam1/2 MPa MPa GPa

1.6 2.5 21 40 100 300 1 1.6 1.7 2.1 60 100 30 45 0.4 0.45 0.4 0.6

Thermal Properties Melting temperature (Tm) Glass transition temperature (Tg) Service temperature in air without mechanical loading (short term) Service temperature in air without mechanical loading (long term)



C C  C

160 180 -30 to -3 140



100



C

6.7.2 BONE CEMENT Bone grafts or synthetic materials are used to treat bone defects. The technique called autograft is cumbersome as the bone has to be transferred from another site within the patient’s body, thus leading to limitations such as site morbidity and availability. Instead allograft tissue can be employed; however, it also has some disadvantages. The use of bone from another person might transfer diseases associated with the selected bone and it will also be difficult to reshape the bone into the desired shape to fit in exactly as the defective bone. Therefore, the use of polymeric materials such as bone cement is gaining attention. Polypropylenebased materials have been utilized as bone cement. An injectable paste composed of a polypropylene fumarate and calcium phosphate composite were characterized (Peter et al., 1999). An injectable polypropylene fumarate calcium phosphate paste was prepared with appreciable properties, such as compressive strength, compressive modulus, gel point, and curing

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FIGURE 6.12 SEM micrographs of silicone rubber composites; (A) 2 wt.% silica and (B) 2 wt.% PP after being soaked in SBF. Reprinted with permission from Elsevier.

6.7 Polypropylene Matrix

behavior. These properties are suitable for clinical orthopedic applications and the mechanical properties of the cured composites are suitable for trabecular bone replacement. Bacakova et al. prepared materials useful for bone tissue engineering (Bacakova et al., 2007). Among them a terpolymer of polytetrafluoroethylene, polyvinyl difluoride, and polypropylene mixed with 4 wt.% of single- or multiwalled carbon nanotubes was reported. Several biomedical studies have conducted on such materials, for instance, seeding was performed with human osteoblastlike MG 63 cells and it was found that CNT-containing materials showed well spreading of the cells and contained distinct beta-actin filament bundles, whereas the cells on the pure terpolymer were rounded and clustered into aggregates. The single-walled carbon nanotube-filled terpolymer showed a large concentration of the components of focal adhesion plaques such as vinculin and talin. This was exhibited by an enzyme-linked immunosorbent assay on the cells by about 56% and 36%, respectively, compared to the pure terpolymer. Osteogenic differentiation was performed by measuring the concentration of osteocalcin. It was found that the concentration of osteocalcin was lower in cells on the terpolymer containing multiwalled nanotubes. The reason is probably due to the more active proliferation of these cells (on day 7, they reached a 4.5 times higher population density than cells on the unmodified terpolymer). The concentration of ICAM-1, a marker of immune activation, in MG 63 cells showed that the addition of both single- and multiwalled nanotubes into the terpolymer had no effect at all. The polypropylene-containing terpolymer material with CNTs showed exemplary results, such as good support for the adhesion and growth of bone-derived cells. These materials were proposed to be considered for the fabrication of bone implants as well as for applications in bone tissue engineering. Chan et al. fabricated binary and hybrid composites based on PP with hexagonal boron nitride (hBN) and nHA (Chan et al., 2015). The prepared composites were tested in order to utilize them for human bone replacements. The composites were prepared through a melt mixing technique and the specimens were prepared by an injection molding technique. The nHA translated its biocompatibility properties to the composites along with improvements in the mechanical properties by virtue of boron nitride particle incorporation. The mechanical properties, such as elastic modulus, showed improvements with respect to hBN content. Cytotoxicity studies, cell cultivation, and MTT assay results showed the attachment and proliferation of osteoblasts on binary and ternary composites. Tjong et al. studied the design, fabrication, and characterization of the microstructure, physicochemical properties, and biocompatibility of PP reinforced with carbon nanofiber (CNF) and HANR fillers (Tjong et al., 2014). The composites were evaluated toward the development of good mechanical behavior, thermal stability, and improved biocompatibility in order to use the materials as craniofacial implants in orthopedics. The composites were fabricated by varying the loading of CNF by up to 2% and making hybrids with 20% loading of HANR using an extrusion technique. The elastic modulus and tensile strength of PP showed

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FIGURE 6.13 Scanning electron micrographs of cultured osteoblasts of PP/2% CNF 20% HANR hybrid composite for 2 days (A) and 4 days (B). (C) High magnified SEM image showing long filopodias. (Abbreviations: CNF, carbon nanofiber; HANR, hydroxyapatite nanorod; KV, kilo volt; PP, polypropylene; WD, working distance). Reprinted with permission from DOVE medical press.

improvement with the addition of CNF, while tensile ductility and impact toughness were not diminished. The addition of a HANR filler clearly showed improvement in tensile properties indicating enhanced filler-matrix interaction. Among the composites, the hybrid with a 2% loading of CNF and 20% HANR showed the maximum mechanical properties, such as tensile strength and stiffness. The thermal stability of the composites also showed manifold improvement due to the incorporation of both the fillers. It was shown that CNFs act as effective nucleating agents as evidenced by DSC measurements. Biocompatibility studies show that CNF nanofillers enhance the cell adhesion and viability of osteoblasts on PP. Among the composites, the PP/2% CNF/20% HANR composite showed good biocompatibility, which was further established by MTT assay results as shown in Fig. 6.13.

6.7.3 SCAFFOLDS In tissue engineering scaffolds are extremely important. They should have appropriate surface chemistry and properties to support cell attachment, proliferation, migration, and growth. The scaffolds should be biocompatible and should act as a

6.7 Polypropylene Matrix

template for cell growth. Additionally, they should act as an aid in the segregation of cells and support the production, organization, and maintenance of any extra cellular matrix. In order to facilitate cell mitigation and nutrient distribution, highly interconnected macro and micro porous networks should be present in the scaffolds. Cell migration depends on the physical aspects of the scaffolds. These physical aspects depend on physical structure and chemical and biological agents in order to help the cell differentiation and adhesion with the surface. Polypropylene as such has difficulty showing all these aspects, and therefore, the modification of PP through blending or fiber formation has been reported. Park et al. reported the preparation of polypropylene carbonates/poly(lactic acid) (PLA) composite nanofibers by sol gel electrospinning and studied their surface morphology, mechanical properties, and cell viability with cultured myoblasts (Park et al., 2016). The mechanical properties showed drastic improvement after a heat treatment and the nanofibers were nontoxic to the cells. Therefore, the composite nanofibers were deemed applicable for biomedical devices. Shi et al. fabricated porous ultra-short CNT nanocomposite scaffolds for bone tissue engineering (Shi et al., 2007). They utilized polypropylene fumarate as the polymer, ultra-short CNTs and modified CNTs for preparing the composites. The porosities were precisely controlled by a technique called thermal crosslinking particulate leaching and porosities of 75, 80, 85, and 90 vol.% were achieved. The porous scaffolds were characterized by different techniques, such as microCT, mercury intrusion porosimetry, and SEM, in order to establish the pore structures. Fig. 6.14 shows SEM images of the scaffolds. All the scaffolds showed 100% interconnectivity at the order of 20 mm and had specific porosities. When the scaffold porosity increased the pore connections also become bigger. This is why the mean pore size of 80 90 vol.% is significantly higher than 75 vol.% scaffolds. It has an adverse effect on the compressive mechanical properties and it got decreased in highly porous scaffolds. Therefore, the advantages of high porosity have to be compromised for many applications. The modification of short CNTs, in fact, reinforced the scaffolds, however, it did not reflect in the mechanical properties. It may be due to the sample variations and preparation techniques. The osteoconductivity of the scaffolds showed excellent results under static culture conditions. Therefore, the highly porous nanocomposite scaffolds prepared from polypropylene fumarate and ultra-short CNTs show potential for development of peculiar bone tissue engineering scaffolds.

6.7.4 ANTIMICROBIAL APPLICATIONS Healthcare devices offer good health to patients and research to develop materials that reduce microbial attack of the devices is always interesting. A large number of healthcare-acquired infections (HAIs) occur through surface contact by hands or other body parts or by devices, such as catheters, surgical incisions, or intravenous lines. A study by Curtis showed that 2% 5% of patients who had undergone

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FIGURE 6.14 SEM images of scaffolds made of: (A1 4) polypropylene fumarate; (B1 4) ultra-short tube nanocomposite; and (C1 4) modified ultra-short tube nanocomposite with increasing porogen fractions of 75, 80, 85, and 90 vol.% (from top to bottom). Scale bar represents 500 mm. Reprinted with permission from Elsevier.

surgery has an infection at the wound site (Curtis, 2008). Also, 80% 95% of HAIs of the urinary tract happens to be from urinary catheters (Dohmen, 2006). Antimicrobial materials need to be developed in order to reduce the number of HAIs. In this regard, polypropylene-based antimicrobial devices are noteworthy (Delgado et al., 2011; Essa and Khallaf, 2016). Palza et al. systematically studied the effect of copper nanoparticles on the antimicrobial properties of polypropylene (Palza et al., 2015). They tested the filler agglomeration and copper ion release from the composite. In order to improve the dispersion, several methods were employed. The compatibilization technique reduced the agglomeration to a large

6.7 Polypropylene Matrix

extent and improved the copper ion release to 40% more than the original matrix material, thus, paving the way for designing materials with tailored antimicrobial properties. Abbas et al. prepared superhydrophobic polypropylene coating suitable for biomedical applications with self-cleaning properties (Abbas et al., 2014). In order to prepare the surfaces several parameters were changed, such as the concentration of PP, solvent evaporation rate, loading of TiO2 nanoparticles and the heating rate for PP dissolution, to name a few. The static contact angle of the superhydrophobic surface was 165 degrees which is noteworthy. To establish the antibacterial properties cytotoxicity studies and bacterial anti-sticking effects were employed. The surfaces showed encouragingly positive results for antisticking effect against Staphylococcus aureus bacterial suspension. Zhao et al. fabricated a polypropylene-based nonwoven fabric membrane (PPNWF) with a switchable surface from antibacterial property to hemocompatibility (Zhao et al., 2013). The fabrication started with the synthesis of a cationic carboxy betaine ester monomer, [(2-(methacryboxy) ethyl)]-N,N-dimethylaminoethyl ammonium bromide, and methyl ester (CABA-1-ester), followed by the introduction of the same molecule via plasma pretreatment and UV-induced graft polymerization on the PPNWF surface. Membranes with different grafting densities were fabricated. Several experimental techniques were adopted to measure the properties and ATR-FTIR and gravimetric methods clearly showed that the cationic modified surface could be transformed to a zwitterionic modified surface under mild hydrolysis conditions. Effective antibacterial property against S. aureus was detected through biological tests for the cationic modified surface while the zwitterionic modified surface showed resistance to protein adsorption, platelet adhesion, and activation. The latter also exhibited enhanced clotting time. Thus, the fabricated PP-based membrane is recommended for dual functional biomaterial applications. Another interesting approach to decrease the growth of bacteria on the surface was reported by Aumsuwan et al. (Aumsuwan et al., 2009). PP surfaces were treated with microwave plasma reactions in the presence of maleic anhydride, which ultimately produced acidic groups. The prepared surfaces were again modified with two molecular groups; polyethylene glycol followed by penicillin V and diglycidyl polyethylene glycol followed by gentamicin. The former one was developed to counter the growth of S. aureus and the latter for creating antimicrobial surfaces so that Pseudomonas putida growth could be curtailed. The antimicrobial strength of the surface was measured simultaneously using gram positive and gram negative bacteria by varying the molecular rations on the surface. Spectroscopic and biological tests indicated appreciable results in the reduction of bacterial growth and the polypropylene surfaces were recommended for the formation of tunable antimicrobial devices.

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6.7.5 SUTURES Surgical sutures are extremely important medical devices employed to knit together the body after a surgery or injury from any cause. Different materials are used to manufacture sutures, such as nylon, polypropylene, biodegradable polymers, etc. There are two types of sutures available currently; absorbable and nonabsorbable. PP-based sutures are nonabsorbable due to their nonbiodegradable nature. Prolene is the trademark of the PP suture available on the market. However, there are many reports of improving the properties of PP sutures by grafting smaller molecules or drugs into it. There are several reports on the development of PP-based sutures available in the literature (Chatzimavroudis et al., 2017; Lo´pez-Saucedo et al., 2017; Tummalapalli et al., 2016).

6.8 CONCLUSIONS Polymer application in biomedical devices and instruments is noteworthy and several discoveries or inventions in the field have raised the standard of human living. Among the polymer polyolefins—polyethylene and polypropylene—were systematically reviewed here. The initial part of the chapter provided a general introduction about the properties of polyolefins and their applications. The different fabrication processes for biomedical devices were touched upon. A brief account of the biocompatibility of polyolefins was also provided. Because of the ubiquitous properties of polyethylene and polypropylene, they were considered for fabricating biocompatible devices. They are cheap, mechanically strong, and biocompatible for many applications. They were used for making devices such as scaffolds, bone cement, antimicrobial applications, hip prostates, and sutures, to name a few. The surface modification of these polyolefins by various techniques has led to the creation of a large number of biocompatible matrices. A detailed account of the application of such surface modification is included in the chapter, which covered drug delivery devices, tissue adhesives, bone substituents, etc. Compared to engineering plastics, the suitability of commodity plastics, such as polyethylene and polypropylene, for biomedical applications is exemplary in the long run. The coming years will witness research and development in the exploration of the potential of polyolefins-based biomedical devices accessible to the common man. Automation and robotics play a big role in developing ultra-pure, sterilizable, and designable materials for the development of medical devices. As explained in the chapter, they offer many advantages, such as low cost and high-performance biomedical products and are sustainable for large scale production due to automation. However, they do have some disadvantages, such as the removal of the degradable products from polyethylene and polypropylene from the human body or body fluids and the commercialization of the currently available products due to the performance inferiority compared to biodegradable and speciality polymers.

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CHAPTER

Polymethacrylates

7

Benjamin Pomes1,2,3, Emmanuel Richaud3 and Jean-Franc¸ois Nguyen1,2,4 1

UFR d’Odontologie, Universite´ Paris Diderot, Paris, France 2Service d’Odontologie Groupe Hospitalier Pitie´ Salpeˆtrie`re, Paris, France 3Arts et Metiers ParisTech, Laboratoire de Proce´de´s et Inge´nierie en Me´canique et Mate´riaux (PIMM), CNRS, CNAM, UMR 8006, Paris, France 4 PSL Research University, Chimie ParisTech CNRS, Institut de Recherche de Chimie Paris, Paris, France

7.1 MATERIAL SELECTION FOR MEDICAL APPLICATIONS: REQUIREMENTS FOR SEVERAL KINDS OF MEDICAL APPLICATIONS A biomaterial is defined, according to the Consensus Conference of Chester (1992), as a material intended to interface with biological systems to evaluate, treat, increase, or replace any tissue, organ, or function of the body. According to the American National Institute of Health, a biomaterial is also described as “any substance or combination of substances, other than drugs, synthetic or natural in origin, which can be used for any period of time, which augments or replaces partially or totally any tissue, organ or function of the body, in order to maintain or improve the quality of life of the individual.” Orthodontic brackets and surgical instruments are not included in this definition (Bergmann and Stumpf, 2013). All materials used for replacing human tissues have joint specifications such as biocompatibility and they must also be noncytotoxic, nonallergic, nonimmunogenic, nonthrombogenic, and noncarcinogenic. Their specifications depend on their applications. Biomaterials suitable for dental restoration applications should meet these requirements: • • • • • • • •

Nonirritating for the pulp and periodontal tissues. Low volumetric variation. Thermal insulation to protect the pulp from temperature variations. Esthetic and the stability of the different shades. Possible and simple repair and replacement. Easy handling. Resistance to water degradation and wear. Polishing ability to obtain a surface roughness limiting the adhesion of the plaque and thus to avoid secondary decay and periodontal diseases.

Materials for Biomedical Engineering: Thermoset and Thermoplastic Polymers. DOI: https://doi.org/10.1016/B978-0-12-816874-5.00007-4 © 2019 Elsevier Inc. All rights reserved.

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• •

•

Be compatible with sealing materials (bond, cement). Prevent leakage and saliva contamination and so dentinal hypersensitivity, coronary re-infection of the pulped tooth, secondary carious lesions and marginal dyschromias. Have mechanical properties adapted to the type of restoration and ideally the closest to the dental tissues: • Flexural strength greater than or equal to that of enamel ($380 MPa), and dentin (219 MPa). • Modulus of elasticity in the range of 90
98 GPa for enamel and 18
22 GPa for dentin. • Hardness Vickers in the range of ,300 HVN for enamel. • Toughness in the range of 0.7
0.8 MPa m1/2 for enamel, 2 MPa m1/2 for dentin (Xu et al., 1998).

7.2 CHEMISTRY OF POLYMETHACRYLATES AND THEIR COMPOSITES 7.2.1 MONOMERS 7.2.1.1 Methyl methacrylate Polymethyl methacrylate (PMMA) is the most well-known polymer of the methacrylate family, obtained from the in-chain polymerization of methyl methacrylate. It was developed in the 1930s by Hill and Crawford for Imperial Chemical Industries in England (Perspex), by Ro¨hm and Haas in Germany (Plexiglas), and by Du Pont de Nemours in the United States (Lucite). It displays several interesting properties such as: • • • •

A higher light transmission than glass (92% of visible light). A low density (1.18 g cm23) being about half that of glass. Shatter proof Softer and easier to scratch than glass (however, scratch-resistant coatings may be applied).

Its first applications as aircraft windows took place during World War II. Medical applications came later and include: • • •

•

In cardiology in pacemakers. For ophthalmology as artificial eye lenses for cataract surgery. For prosthetic dentistry in removable total and partial dentures (Fig. 7.1), temporary fixed denture, and restorative dentistry and for orthodontic devices. As bone cement for orthopedic surgery of the hip, knee, and other joints for the fixation of polymer or metallic prosthetic implants on living bone.

7.2 Chemistry of Polymethacrylates and Their Composites

FIGURE 7.1 Removable total denture.

7.2.1.2 Other methacrylates for dental applications In the 1950s, composite restorations were made from PMMA. Significant polymerization induced volume shrinkage and heat release as well as the release of methyl methacrylate monomers resulting in marginal discoloration, pulp reactions, and secondary caries. In 1956, Bowen tried to reduce shrinkage by using a bigger monomer: the bisphenol A surrounded by two glycidyl groups. However, moisture tended to inhibit polymerization. Bisphenol A glycidyl methacrylate (Bis-GMA, Fig. 7.2) was obtained by substituting end groups with methacrylate functions and offered an excellent solution to the problem (Bowen, 1962; Soderholm and Mariotti, 1999). Bis-GMA has several advantages compared to PMMA, such as lower volatility, lower diffusion in dental tissues, and lower polymerization shrinkage because of its larger size, which explains its success. In addition, its tetrafunctional structure (i.e., two double bonds) makes it possible to obtain a crosslink network with better mechanical and physical properties. However, Bis-GMA shows a high viscosity (1200 Pa s) (BarszczewskaRybarek, 2009) because of its large size, and rigidity due to the two aromatic rings and especially its hydroxyl groups creating strong hydrogen intermolecular bonds between them which significantly reduces the mobility and make the molecule more hydrophilic. As it will be seen later, this involves a low conversion degree (DC)—39% in the case of photopolymerization and a limited incorporation of fillers as well as difficulty to handle the material (Sideridou et al., 2002). Consequently, Bis-GMA is always associated with minor monomers such as triethylene glycol dimethacrylate—TEGDMA, with a viscosity of 0.011 Pa s

219

220

CHAPTER 7 Polymethacrylates

H2C MMA O

C C

H 2C

CH3 HEMA

O

OMe

O H2C

bis-GMA

C C

CH3 OH

O

O

CH3 OH

CH3

OH

CH3

TEGDMA

O

O

O

O

O bis-EMA

O

O

CH3

O

O

C

O

O

O H2C CH3

C CH3

CH2

O

CH3 O

O

CH3

CH3

CH3

bis-DMA

CH2 O

CH3

H2C

CH3

CH3

O H2C

CH2

O

O

C

O

O

O

CH2 CH3

FIGURE 7.2 Some methacrylate monomers.

(Ilie and Hickel, 2011; Moszner and Salz, 2001)—or other derivatives such as bisphenol A ethoxyl dimethacrylate (Bis-EMA) (452 g mol21) and bisphenol A propoxyl dimethacrylate Bis-PMA (480 g mol21) and having a lower viscosity than Bis-GMA (because of its lower molar mass and the absence of hydroxyl groups). UDMA monomers were developed in the 1970s by Forster and Walker (1974). Those UDMAs constitute a wide family of molecules differing by their molar mass and their structure (Fig. 7.3). The 1,6-bis(methacryloxy-2-ethoxycarbonylamino)-2,4,4-trimethylhexane is by far the most used and has a very low viscosity (23.1 Pa s) (Barszczewska-Rybarek, 2009).

7.2.1.3 Composition of the matrix Monomers Mixtures of several monomers are currently encountered in the literature for model and commercial materials. Usually, UDMA or Bis-GMA are used as major monomers together with minor monomers such as TEGDMA, HEMA, and bis-EMA for lowering the viscosity (Ru¨ttermann et al., 2010; Rahim et al., 2012; Aljabo et al., 2015; Bhamra et al., 2010; Thomaidis et al., 2013). The use of minor monomers lowers viscosity which aims at incorporating more fillers to improve the mechanical properties, and to increase the conversion degree (Floyd and Dickens, 2006).

7.2 Chemistry of Polymethacrylates and Their Composites

H

O H2C

H

O H2C CH3

CH3

O

H2C

O

CH3

O

O

CH3

O

CH3

CH3

H

H

H

N

N

CH2 O O

O O

O

O

O

N

CH3

CH2 O

H

N

O

O

CH3

O

O

N

CH3

O

CH3

CH3

O

CH3

N

O

O

CH3

O

CH2 CH3

O

O N

O H2C

O

CH3

CH3

FIGURE 7.3 Some urethane dimethacrylate monomers.

Activators and polymerization initiators The initiation of the polymerization reaction requires the creation of radicals coming from, for example: •

•

For chemopolymerization: amines (para-amino methyl acetate, paratoluenesulphonic acids, thioureas, and ascorbic acids) and peroxides (benzoyl peroxide as shown in Scheme 7.1, cumene peroxide, and terbutyl hydroperoxide). For photopolymerization: camphorquinone (Fig. 7.4) in combination with an aromatic amine. The camphorquinone displays an absorption in the range of 400
550 nm with a λmax at 470 nm (Leprince et al., 2013) leading to radical generation, as illustrated in Scheme 7.2.

Polymerization inhibitors Phenolic compounds react with free radicals and are used to avoid the possible spontaneous polymerization occurring during monomer storage.

Coupling agents The coupling agent is an amphiphilic molecule bonding the hydrophilic inorganic filler and the hydrophobic resin. One of the most common molecules is 3-(Trimethoxysilyl)propyl methacrylate (Fig. 7.5).

221

222

CHAPTER 7 Polymethacrylates

O

O

O

C O O C

O

C O° °O C

°

+ CO2

SCHEME 7.1 Initiation by benzoyl peroxide.

(A)

(B)

CH3

COOH

N

O

CH3 H3C

O

FIGURE 7.4 (A) 4,4 N trimethylaniline; and (B) camphorquinone used for initiating photopolymerization.

Amine coinitiator

* Intersystem 3

1 O

° O

°° N

*

CH3

° O

crossing

Exciplex state

3 ° O-

°N+ CH3

Activated singlet state

* ° OH

°° N CH2°

2 Initiator radicals

Activated triplet state

SCHEME 7.2 Activation of camphorquinone.

CH3 O

H2C

OMe OMe OMe

O

FIGURE 7.5 3-(Trimethoxysilyl)propyl methacrylate.

7.2.2 DENTAL COMPOSITES Composite material can be defined as a combination of two or more immiscible materials of different chemical natures, leading to better properties than those of the individual components used alone. They are constituted by a matrix and dispersed reinforcements. In the case of an organic matrix, the reinforcement is an inorganic solid (glass, ceramic, metal) in the form of fibers, particles, or flakes. The properties of the composite material depend on the filler volume fraction, shape factor (or length:diameter ratio), and orientation (Kardos, 1993). The matrix allows the transmission of the mechanical stresses to the reinforcement, the

7.2 Chemistry of Polymethacrylates and Their Composites

protection of this later against the external environment, and determines the conditions of use and processing. Their composition has evolved since their introduction in odontology more than 50 years ago. Their clinical success would not have been possible without an understanding of the adhesion phenomena allowing their adhesion to dental tissues: enamel (Buonocore, 1955) and dentin (Nakabayashi et al., 1991). The inorganic fillers may be silicas (SiO2) in crystalline form such as quartz, or in amorphous form such as borosilicate glass, or heavy metal (Sn, Ba) glasses. Their shape can be angular (obtained by grinding), rounded (obtained by melting), or square with rounded corners (Raskin et al., 2006). Organomineral fillers are crushed prepolymerized composites and then added to the monomer/filler mixture, which makes it possible to reduce the shrinkage and to adapt the viscosity of the composite. Organoorganic fillers (trimethylolpropane trimethacrylate) and ceramics with grafted methacrylate groups (OrMoCers) (Raskin et al., 2006; Raskin, 2011) can also be used. Composites can be classified according to the particle size and distribution of the fillers, the viscosity and the mode of polymerization, as dicussed next.

7.2.2.1 Particle size and distribution of fillers Ferracane (2011) proposed to classify dental composites according to the fillers size (Fig. 7.6). 1. Macrofill: The first macrofill conventional composites presented 10
50 μm fillers, providing excellent mechanical properties, but difficulty in polishing and degradation at the surface by abrasion. 2. Microfill: microfill composites with 40
50 nm fillers were developed to overcome these disadvantages. They have good polishing properties, but low mechanical properties. 3. Hybrid: Hybrid composites are a mix of the particle sizes of the two previous families and, therefore, were used as a compromise between the mechanical, optical, and polishing properties. 4. Midfill: The tendency was then to reduce the size of the fillers to result in the hybrid midfill composites with charges of 1
10 μm and 40 nm. 5. Minifill: The evolution continued with the decrease in the size of the fillers with the minifill composites, with 0.6
1 μm and 40 nm filler, from which microhybrid composites were used for restorations in the anterior and posterior sectors. 6. Nanofill: The most recent innovations concern the development of nanofilled composites with 5
100 nm nanoparticules. However, this filler family has declined because of the difficulty of incorporating nanoparticles, the presence of numerous defects linked to the exponential increase of the matrix/fillers interface, and their high viscosity. To solve these problems, the nanofillers were partially sintered into nanoparticle aggregates (cluster) ranging from 5 to 75 nm in diameter (O’Brien, 2008).

223

224

CHAPTER 7 Polymethacrylates

1

Macrofill (10–50 μm)

2

Microfill (40–50 nm)

3

Hybrid (10–50 μm + 40 nm)

4

Midfill (1–10 μm + 40nm)

6

5 Minifill (0.6–1 μm + 40 nm)

Microhybrid

Nanohybrid

FIGURE 7.6 Classification of dental resin composites.

Nanofill (5–100 nm)

Nanofill-cluster

7.2 Chemistry of Polymethacrylates and Their Composites

The size of filler particles incorporated into the resinous matrix of commercial dental composites tends to decrease over the years (Ilie and Hickel, 2011; Ferracane, 2011). Currently, microhybrid composites and nanohybrid composites (i.e., microhybrides with nanoparticles) are commercially available and display relatively close mechanical properties. Fiber-reinforced composites constitute a relatively new class of dental composites and are drawing increasing interest (Ballo and Na¨rhi, 2017).

7.2.2.2 Viscosity According to Einstein’s formula, spherical fillers lead to a viscosity increase given by: η 5 η0 Uð1 1 2:5 3 φÞ

(7.1)

η and η0 being, respectively, the viscosity of filled and unfilled resin, and φ the filler ratio (in volume). Adding c. 50%
60% fillers by weight (i.e., 80% in volume) leads to viscosity almost 30-times higher than for base resin (Papakonstantinou et al., 2013) and can reach more than 2000 Pa s. Reactive mixtures for dental resin composites cover a wide range of viscosities which allow them to meet the requirements of numerous clinical indications. Fluid composites with low viscosity have 0.4
3 μm filler, a Vf of 42%
53%, are packaged as syringes (Fig. 7.7), and indicated for cervical restorations or for low-tissue losses. The high viscosity condensable composites have Vf of 66%
70% and are indicated for site tissue losses 1 and 2 (Sakagushi and Powers, 2012).

7.2.2.3 Polymerization mode The composites can also be classified according to their mode of polymerization: photopolymerization (Fig. 7.8), chemopolymerization, dual photo- and chemopolymerization and more recently high-pressure high-temperature polymerization

FIGURE 7.7 Resin composite in syringe.

FIGURE 7.8 Polymerization lamp.

225

226

CHAPTER 7 Polymethacrylates

FIGURE 7.9 Polymer infiltrated ceramic network block suitable for CADCAM.

for industrial composite blocks suitable for CADCAM applications, which are discussed further next (Fig. 7.9). Photopolymerization is usually performed using lamps with 150
600 mW cm22 irradiance, for 10
60 seconds durations. Photopolymerization of dental resin composites is performed using LED or halogen light curing units with, respectively, an emission peak in the 450
470 nm (Issa et al., 2016) and 450
520 nm (Bala et al., 2005) wavelength ranges. The emission spectra and characteristics of some commercial lamps can be found in (Haenel et al., 2015). CAD/CAM applications suitable for the manufacture of dental restorations are currently being developed (Van Noort, 2012; Miyazaki et al., 2009) because they obtain a constant quality. Chairside CAD/CAM can be used for prosthesis in dental surgery. In addition, blocks for CADCAM applications are manufactured industrially and so are more homogeneous and have fewer defects than handled materials. Two types of blocks are currently commercially available: ceramic blocks and composite blocks. Ceramic blocks have superior mechanical properties and their chemical inertia which give them good biocompatibility. However, they are difficult to machine and cannot tolerate plastic deformations, which leads to the risk of fractures at the fine edges during the machining of the dental prosthesis. Thus, reoperation is more delicate. Composite blocks have lower mechanical properties, lower wear resistance and are less biocompatibile due to residual monomer release from incomplete polymerization. Nevertheless, they are easier to set on dental tissues and reoperation is easier. They also have better machinability and polishability. Indeed, the use of composite resin blocks designs for CAD/CAM significantly reduces the machining time and tool wear (Mainjot et al., 2016). New technologies have been developed on polymerization and the methods for producing the blocks in order to increase the mechanical properties of the composites and to increase their strength, longevity, and biocompatibility. Conventional thermopolymerization and photopolymerization have the disadvantage of being incomplete resulting in a low degree of conversion (56%
67%)

7.2 Chemistry of Polymethacrylates and Their Composites

(Ferracane et al., 1997). Moreover, it induces internal stresses in the composite resulting from the shrinkage and the differential polymerization between the superficial part close to the source of irradiation and the deeper part (Ferracane, 2005). To improve the mechanical properties of a composite, the polymerization mode can be enhanced. A previous study has shown that high temperature (180 C) and high pressure (250 MPa) polymerization allowed a significant increase in the mechanical properties of commercial composites compared to conventional photopolymerization (Nguyen et al., 2012). Dispersed filler composite blocks are synthesized by thermopolymerizing (under high pressure or not) a mixture of mixed fillers and monomers (Mainjot et al., 2016; Nguyen et al., 2013). PICN blocks present a particular microstructure as they are synthesized from a sintered glass-ceramic network with a φf greater than 73.8% in the form of a block, secondarily infiltrated by monomers, and then thermopolymerized under high pressure. Their fundamental characteristic is that they consist of two continuous networks imbricated in one another: • •

A sintered glass-ceramic inorganic network with open porosity. An organic network constituted by the crosslinking of a dimethacrylic monomer inside the inorganic network.

The PICN microstructure allows for a higher fillers ratio compared to classical dispersion, and results in higher mechanical properties (Nguyen et al., 2013).

7.2.3 CHALLENGES IN IMPROVING PROPERTIES Manufacturers try to improve dental composites performance by modifying the formulation of monomers and photoinitiators. Composites with other monomers such as siloxane, oxirane, or silorane have been developed in order to attenuate the shrinkage, but do not provide significant improvement in the mechanical properties (Lien and Vandewalle, 2010). Alternative photoinitiation systems based on mono-acyl phosphine oxide (MAPO), a bioacyl phosphorine oxide with a better production efficiency of free radicals than camphorquinone, allowing an increase in conversion degree, mechanical properties, and better polymerization in depth. These systems also improve biocompatibility because no tertiary amines are needed to generate free radicals. However, their absorption spectrum corresponds less to commercially available photopolymerization lamps (Leprince et al., 2013). Improving composites is a complex process. Several papers have been aimed at comparing the performances of various reactive mixtures (Papakonstantinou et al., 2013; Fonseca et al., 2017). Indeed, changing the nature or ratio of a given component can induce undesired side effects (sometimes minor). The overall possible effects are summarized in Figs. 7.10
7.12.

227

228

CHAPTER 7 Polymethacrylates

↑ Viscosity

↓ Conversion degree

+++ –

↑ Monomer release ↓ Internal stress

↑ φfiller

Mechanical properties

+

↓ Shrinkage – +++ Biocompatibility ↓ φresin

↓ Monomer release

+++

FIGURE 7.10 Effect of increasing the filler ratio (φf). ++

↓ Shrinkage

↓Internal stress

Mechanical properties

+

↑Major monomer

–

–

↓ φfiller

↑ Viscosity

↓Conversion degree

↓Monomer release

↑Monomer release

+ ––

–

Biocompatibility ++

FIGURE 7.11 Effect increasing the concentration in the major monomer.

7.3 METHODS FOR MATERIAL SYNTHESIS 7.3.1 RADICAL POLYMERIZATION REACTION OF PMMA (DIFUNCTIONNAL MONOMER) 7.3.1.1 Mechanistic aspects The anionic polymerization of MMA can be made in the presence of YCl3/lithium amide of indoline/nBli (Ihara et al., 2007). Despite the interesting features of anionic MMA—polydispersity index close to 1, and possibility to get block

7.3 Methods for Material Synthesis

–––

↑ Internal stress

↑↑↑Shrinkage

–––

Mechanical properties

+ ↑ Minor monomers

+

↑ φfiller –––

↑ Conversion degree

↓Viscosity

↓ Monomer release

↑↑↑ Monomer release

–––––

+

+

Biocompatibility

FIGURE 7.12 Effect increasing minor monomers.

H2C

CH3

A

A° O

OMe H2C

A° O

C C

CH3

CH2 ° CH3 O

A

OMe

CH2 ° CH3 C

OMe

O

C

OMe

SCHEME 7.3 Chain initiation of MMA.

copolymers (Baskaran and Mu¨ller, 2007)—most of the industrial and medical PMMA grades are obtained from radical polymerization (O’Brien, 2008; Powers and Sakagushi, 2006). The first step (initiation) corresponds to the creation of radicals by photo- or thermochemical processes (see Schemes 7.1 and 7.2, respectively): A 1 heat or UV-A

The free radicals break the CQC bonds of the monomers to form the first elements of the increasing polymer chain (Scheme 7.3). Then, during the propagation phase, polymers are formed by the successive addition of monomers (Scheme 7.4). The propagation reaction which corresponds to the opening of CQC double bonds is, in essence, exothermic. The heat of polymerization is, thus, given by: ΔHpolym 5 ΔHC5C
ΔHðC2CÞmonomer
ΔHðC2CÞmonomer2monomer

(7.2)

In PMMA, the hindering effect of acetyl groups make the third term quite low compared to other polymers so that the heat of polymerization is lower than for

229

230

CHAPTER 7 Polymethacrylates

H2C

CH2 ° CH3 C A O

C

A

CH2

OMe

O

CH3 C C

CH2

OMe O

O

C C

° CH3 C C

CH3 OMe A

CH2

OMe

O

CH3 C C

OMe

CH2

n

O

° CH3 C C

OMe

SCHEME 7.4 Chain propagation of PMMA.

other olefins (Roberts, 1950). However, the temperature in the bulk of a PMMA made bone cement polymerizing at room temperature reach about 80 C (Khandaker and Meng, 2015). Finally, the termination phase closes the reaction by coupling two reactive polymers, or a reactive polymer with a reactive monomer, resulting in the formation of a stable covalent bond (Scheme 7.5). One of the most remarkable properties of PMMA is that it is soluble in its monomer. In the applications of PMMA as bone cement, PMMA powder mixed with an initator (benzoil peroxide) is hence mixed with a MMA monomer containing N,N dimethyl p-toluidine (Asgharzadeh Shirazi et al., 2017). The PMMA is, hence, dissolved in monomer which polymerizes to give a glassy solid.

7.3.1.2 Kinetic aspects The polymerization mechanism of MMA can be represented in a simple way as (Cardenas and O’Driscoll, 1976): Initiation: Propagation: Termination:

I-2A A 1 M-A-M A-Mi-M 1 M-A-Mi11-M A-Mi-M 1 A-Mj-M -A-Mi1j12-A

kd ki kp kt1

It is assumed that the termination rate constants for combination and chain transfer to monomers are insignificant compared to the termination rate constant for disproportionation. Under the assumption of classical chemical kinetics (steady state hypothesis on A , and on the overall concentration in radical species [R ]), it can be shown that: d½AM  5 2kd ½A 2 kp ½AM ½M  2 kt ½M½R  dt

(7.3)

d½R  5 2kd ½A 2 kt ½R 2 dt

(7.4)

rpolymerization 5 2

d ½M 2kd kp ½A½M 5 dt kt

(7.5)

A

A

CH2 O

CH3 C C

n

OMe

CH2

° CH3 C O

C

OMe

H3C O

C° C

CH3

CH2 OMe

C O

C

CH2 O

CH2

m

OMe

CH2 O

Chain termination of PMMA.

C C

CH2

n

OMe

O

CH3

CH3

C

C

C

C

OMe O

CH3

CH2

OMe

C O

C

CH2

m

A

OMe

A

A

SCHEME 7.5

CH3

CH3 C C

CH n

OMe O

C C

CH3 OMe

H3C O

H C C

CH3

CH2 OMe

C O

C

CH2

OMe

m

A

232

CHAPTER 7 Polymethacrylates

Table 7.1 Approximate Values of Kinetic Parameters for Polymerization (Initiation by AIBN) kd kp kt

s21 I mol21 s21 I mol21 s21

1015exp(261,000/RT) 2.7 3 106exp(210,600/RT) 108exp(21400/RT)

FIGURE 7.13 Monomer conversion versus polymerization time.

It can also be shown that the solution of those differential equations fairly represent the early stages of the polymerization reaction. Some values of kinetic parameters are given in Table 7.1 (Begum et al., 2012). During polymerization, the molar mass increases, which lowers the mobility of radicals and the termination rate. At a certain stage, it results in observed autoacceleration (i.e., the “Thromsdorff effect”), as illustrated in Fig. 7.13. Cardenas and O’Driscoll (1976) proposed to model autoacceleration by using a termination rate constant for nonentangled growing chains and another for entangled ones. Later, Simon and colleagues (Begum and Simon, 2011) proposed a more-refined theory taking into account the role of free volume and its consequences on the diffusion rate of radicals. These developments, however, are out of the scope of this chapter. The most interesting research deals with the copolymerization of several chemically different monomers which is relevant with reactive mixtures presented in Section 7.2.1.3. The theoretical treatment of copolymerization was proposed by Mayo and Lewis (1944). To summarize, the growing chain can be terminated either by an A or B site, which reacts either with a free A or B monomer:

7.3 Methods for Material Synthesis

---A 1 A - ---AA ---A 1 B - ---AB ---B 1 A - ---BA ---B 1 B - ---BB

kAA kAB kBA kBB

The reactivity ratio kAA/kAB and kBA/kBB together with the composition of the reactive mixture, thus, give a prediction of the polymer composition and microstructure (i.e., random, alternating, or block copolymer). This theory was not, however, applied to methacrylates copolymers to the best of our knowledge.

7.3.2 POLYMERIZATION OF METHACRYLATE NETWORKS 7.3.2.1 Mechanistic aspects Networks are used by using a tetrafunctional monomer (i.e., having two double bonds) such as dimethacrylates presented in Section 7.2.1.2. The processes is described in Scheme 7.6 for MMA: This tetrafunctional behavior makes various kinds of intramolecular cyclization reactions possible together with intermolecular crosslinking as presented in Schemes 7.7 and 7.8 (Elliott et al., 2001). °

A A°

A

°

SCHEME 7.6 Initiation step for network polymerization. A

° A

A

°

A

°

° n

n

SCHEME 7.7 Mechanism of crosslinking.

A n -8

A

° n

SCHEME 7.8 Mechanism of primary cyclization.

°

233

234

CHAPTER 7 Polymethacrylates

7.3.2.2 Polymerization kinetics The polymerization of networks by radical polymerization is often divided into four successive steps as schematized in Fig. 7.14 (Pascault et al. 2002a). The first step (pregel step) corresponds to the consumption of inhibitors and their reaction with monomers and the first propagations reaction. The second step (“gel step”) corresponds to the appearance of the first insoluble compounds. Growing chains react either with a monomer or by intramolecular primary cyclization. It results in the formation of crosslinked compact molecules very often called microgels, but better defined as crosslinked microparticles. The gel point is defined by a conversion degree at 1% in theory, but is observed in practice to be around 5% due to cyclizations. Despite the decrease in molecular mobility due to the continuous growth of polymer chains. Monomers can still diffuse and react at the periphery of the microgels in formation. During the third step, the microgels connect to form macrogels which result in an increase in viscosity, so that the mobility of the polymers and monomers is reduced. At the end of this phase, the viscosity is so increased that termination reactions involving macromolecular radicals are inhibited, leading to a sudden autoacceleration of the polymerization rate. The fourth step (“glassy step”) corresponds to the vitrification where polymerization media turn to a glassy state, that is, that polymerization is frozen by lack of macromolecular mobility which explains why the conversion never reaches 100% (Duˇsek, 1996). The polymerization reaction can be completed only by an adequate postcuring step at temperatures higher than the glass transition of the vitrified network.

FIGURE 7.14 Kinetics aspects of network formation.

7.3 Methods for Material Synthesis

7.3.3 PARAMETERS INFLUENCING POLYMERIZATION 7.3.3.1 Intrinsic factors Polymerization intrinsic factors depend essentially on the chemical composition and method of manufacture of the composite: 1. The viscosity of monomers is an important parameter for the polymerization kinetics and conversion of the dimethacrylate polymers by affecting the mobility and the reactivity of the monomers (Barszczewska-Rybarek, 2009). 2. Fillers decrease the conversion by increasing the overall viscosity of the composite and, locally, by reducing the mobility of the monomers around the fillers. For example, adding 50% fillers leads to a strong increase in viscosity (almost 30-times higher), but only a minor decrease in conversion degree (about 63% vs 61%) (Papakonstantinou et al., 2013). Furthermore, fillers can alter the photopolymerization by dispersing the photons superficially (Leprince et al., 2013). 3. Higher concentrations of initiator increases the conversion degree. Nevertheless, in the case of photopolymerization, when the initiator concentration exceeds an optimum value, the conversion degree decreases due to the excessive absorption of photons in the irradiated surface area and, thus, a decrease in the photon transmission in depth (Musanje et al., 2009). 4. Optical properties affect the photon transmission and, therefore, influence the conversion degree and the polymerization depth during photopolymerization. The photon transmission will be reduced in an opaque composite with a darker and more saturated hue, which increases the difference in degree of polymerization between the surface and at depth (Musanje and Darvell, 2006; Shortall et al., 1995). This is, for example, illustrated in (Aljabo et al., 2015) where 40%-filled composites displayed a conversion degree of about 70% at the surface versus c. 40% at a 4 mm depth.

7.3.3.2 Extrinsic factors Temperature and pressure conditions influence polymerization. Higher temperature promotes molecular mobility. Higher pressure decreases molecular mobility (Murli and Song, 2010), but paradoxically has some beneficial effects on the reactivity (Schettino et al., 2008). In the case of photopolymerization, the photon source affects the polymerization by its emission spectrum, irradiation time, irradiation distance, and polymerization protocol (Musanje et al., 2009). The effect of curing protocol on the main properties (glass transition, modulus, conversion degree) is developed in Dewaele et al. (2009).

7.3.4 POLYMERIZATION SHRINKAGE AND ITS CONSEQUENCES The polymerization shrinkage of dental resins composite is inherent to polymerization reactions. This is due to the replacement of the Van der Waals bonds between

235

236

CHAPTER 7 Polymethacrylates

the monomers by covalent bonds and a decrease in the free volume (Kleverlaan and Feilzer, 2005). The shrinkage is about 1.5%
5% by volume (Floyd and Dickens, 2006; Ferracane, 2005) and depends on the concentration of the CQC of the monomers, the volume fraction of the composite, and the conversion degree. The polymerization shrinkage is associated with stresses at the interface between the dental tissues and the composite restoration, and induces: • •

Stresses on dental structures with fracture risks in enamel and dentine (Ferracane, 2005; Park and Ferracane, 2006). Stresses at the joint between the restorative material and the dental tissues resulting in leakage and postoperative sensitivities, marginal discolorations, bacterial contamination, and secondary caries.

This polymerization shrinkage also provokes contraction in the composite resin inducing internal stresses in the material (Ferracane, 2005).

7.4 PHYSICOCHEMICAL, BIOLOGICAL AND MECHANICAL PROPERTIES 7.4.1 STRUCTURE
PROPERTIES RELATIONSHIPS AND LINK WITH CLINICAL APPLICATIONS 7.4.1.1 Glass transition temperature and other transitions The commonality of polymethacrylates obtained by radical polymerization is that they are amorphous materials. The main transition is the glass transition Tg separating the glassy and the rubbery regimes. 1. The glass transition of linear polymer (here PMMA) increases with molar mass, as described by the Fox-Flory’s equation (Fox and Flory, 1950): Tg 5 TgN 2

KFF Mn

(7.6)

Some values of TgN and K are given for PMMA in Table 7.2. Table 7.2 Flox-Flory Parameters for PMMA (Cardenas and O’Driscoll, 1976; Lu and Jiang, 1991) a-PMMA a-PMMA i-PMMA s-PMMA

TgN (K)

KFF (K kg mol21)

387 388 318 405

2.105 21.104 11.104 20.104

7.4 Physicochemical, Biological and Mechanical Properties

2. The glass transition of thermoset networks increases with crosslinking density as, for example, illustrated in polymethylmecrylate networks crosslinked with ethylene glycol methacrylate (Gilormini et al., 2017) or other tetrafunctional acrylates (Loshaek, 1955). The most general equation linking Tg increase with the conversion degree of monomer was proposed by Pascault and Di Benedetto (Pascault and Williams, 1990): Tg 2 Tg0 λx 5 TgN 2 Tg0 1 2 ð1 2 λÞx

(7.7)

where subscripts “0” and “N” correspond to totally unreacted and totally reacted materials, λ is the ratio of heat capacity jump a Tg of cured and uncured materials ΔCpN/ΔCp0. The glass transition of a fully cured network (TgN) is given by DiMarzio’s equation (DiMarzio, 1964): TgN 5

Tgl 1 2 ðKDM Fn0 Þ

(7.8)

where KDM is the DiMarzio’s constant equal to 2.91 for tridimensional networks such as epoxies (Bellenger et al., 1987), n0 is the crosslink density (mol kg21) 5 2/Mm if network is fully cured (Mm is the mass of monomer), Tgl is the glass transition of a “virtual” linear polymer (n0 5 0), F is the flex parameter (kg mol21) related to the molar mass per rotatable bond. The calculation of parameters of DiMarzio’s law is illustrated in Bellenger et al. (1987) and Ernault et al. (2017). Some results are given in Table 7.3. In the case of UDMA, a good agreement is found with values obtained for materials being almost fully cured (Chi Phan et al., 2015). However, Eqs. (7.7)
(7.8) show that Tg decreases dramatically if networks are not totally crosslinked, which is very often the case in networks cured at room temperature in dentistry. This is, for example, illustrated in the case of photocured Bis-GMA-TEGDMA (Table 7.4) (Stansbury, 2012). The consequences of undercuring on mechanical properties at room temperature are illustrated, for example, by Ferracane et al. (1998). The main results are summarized in Table 7.5. Table 7.3 Theoretical Maximum Glass Transition Temperature of Fully Cured Methacrylate Networks Used for Dental Applications BisGMA UDMA BisEMA TEGDMA

Tgl (K)

F (g mol21)

n0 (mol g21)

TgN (K)

327.4 320.6 320.7 268.4

18.3 15.4 17.2 14.7

0.0039 0.00425 0.0037 0.00699

413.2 399 394 382.9

237

238

CHAPTER 7 Polymethacrylates

Table 7.4 Characteristics of Photocured Bis-GMA-TEGDMA Networks Curing time (s) Conversion degree (%) Approximative Tg ( C)

25 36.3 15

45 47.5 25

60 55.1 50

15 68.3 90

Table 7.5 Effect of Curing Degree in a Bis-GMA-TEGDMA (50/50) Matrix Reinforced with 62% Fillers Conversion Degree (%)

E (GPa)

KIC (MPa m1/2)

σf (MPa)

Hardness (kg mm22)

55 60 61 64 66

6.38 8.94 10.35 11.41 14.27

1.29 1.61 1.79 1.89 2.19

88.7 109.2 115.2 117.1 155.4

63.5 73.2 77.0 86.3 93.4

β transition corresponds to the activation of local mobility involving the group of atoms belonging to a monomer. It usually corresponds to a decrease in modulus: ΔEβ 5 1300 MPa. In the case of PMMA, a relative jump of 20% compared to “modulus at 0 K” (i.e., deduced from Eqs. 7.9
7.11) is reported (Gilbert et al., 1986). This transition is, however, not documented, to the best of our knowledge, for dental composites.

7.4.1.2 Short deformation properties Typical stress
strain curves of PMMA at several temperatures and strain rates are schematized in Fig. 7.15 (Moy et al., 2011). In dimethacrylate networks (Foroutan et al., 2011), stress
strain curves usually do not display any “hook” typical of a plastic deformation, which is why they are quite often characterized by their values of elastic modulus and ultimate strength. The elastic behavior of a polymer originates from its cohesive energy which continuously decreases due to the thermal expansion increasing the interchain distance, and the activation of motions decreasing elastic modulus by ΔEi at some given temperatures (e.g., β or γ transitions) (Gilbert et al., 1986):
 X T 2 ΔEi E 5 E0  1 2 α Tg

(7.9)

According to the linear elasticity theory, Young’s (E), bulk (K), and shear (G) modulus values at any temperature are interrelated by Eqs. (7.10)
(7.11): E 5 3UKUð1 2 2vÞ

(7.10)

E 5 2ð1 1 vÞUG

(7.11)

where v is the Poisson’s ratio (see later).

7.4 Physicochemical, Biological and Mechanical Properties

FIGURE 7.15 Typical stress
strain curves in PMMA.

Table 7.6 Estimation of Cohesive Energy and Compression Modulus TEGDMA UDMA Bis-GMA Bis-EMA PMMA

Ecoh (J mol21)

Vm (cm3 cm21)

CED (MPa)

K (GPa)

89,640 169,530 225,930 185,930 33,830

184.9 342.9 339.4 393.4 81.9

485 494 666 473 413

5.3 5.4 7.3 5.2 4.5

Their subglassy values K0 and E0 at very low temperature are shown to be correlated the cohesive energy density (CED) according to the formula (Pascault et al., 2002b): K0 B11UCED CED 5

Ecoh Vm

(7.12) (7.13)

where Ecoh is the cohesive energy (J mol21), Vm is the molar volume (cm3 mol21), CED is the density of cohesive energy (MPa1/2). Ecoh and Vm can be calculated according to the incremental method based on additive group’s contribution proposed by Van Krevelen and Te Nijenhuis (2009), as shown in Table 7.6. If the Poisson’s ratio is on the order of 0.3
0.4 (see later), the Young modulus of the matrix is, thus, expected to be close to 5 GPa. However, experimental values are usually below these maximum values for several reasons, two of these being: •

Firstly, the undercured characteristics of the networks. A good example of modulus increase with conversion degree is given in Stansbury (2012) (see

239

240

CHAPTER 7 Polymethacrylates

•

also Table 7.5). The gelation is observed at a conversion degree about 7% at which materials are a visco-elastic solid having an elastic modulus c. 100 Pa. When the conversion degree reaches c. 20%, the modulus is close to 10 MPa. It exceeds 100 MPa when the conversion degree is more than 0.5. Secondly, the existence of several thermomechanical transitions increasing the mobility of groups, monomers, and later units made of several monomers (e.g., the β transition presented in Section 7.4.1.1 and decreasing, step-by-step, the modulus.

At room temperature, that is, presumably in the subglassy domain (between Tβ and Tg), Young’s modulus values are, thus, on the order of 1 GPa for some lightcured unfilled dimethacrylates (Sideridou et al., 2003; Bindu et al., 2013). In the case of highly filled commercial materials, several equations describe the effect of fillers on mechanical properties (Atai et al., 2012). One of the most wellknown is the Halpin-Tsai equation in the case of spherical fillers (Pal, 2005): 2   3 Ed 5 4 2 Em 2 2 5 Er 5 1 1 φ   2 2 Ed 1 3

(7.14)

Em

Despite the modulus of “pure” fully cured matrix is in the order of 5 GPa (see Table 7.7), it is not surprising that the elastic modulus of commercial dental materials increases linearly with filler content (see, e.g., Masouras et al., 2008) and can reach values c. 10 GPa (Ferracane et al., 1998; Papadogiannis et al., 2015; Jager et al., 2016b). It can be observed that the filler ratio influences the viscosity of the reactive mixture and later its conversion degree (see Fig. 7.10 and Ferracane et al., 1998) so that predicting the value of composites modulus remains intricate. Experimentally, as illustrated in the case of PMMA (Mott et al., 2008), the modulus at very low temperatures can be estimated from ultrasonic measurements from a relationship between the longitudinal and the shear wave velocities (VL and VT) of the sample immersed in water and the density: sffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi Eu 1 2 νu  VL 5 ρ ð1 1 ν u Þð1 2 2ν u Þ sffiffiffiffiffiffi Gu VT 5 ρ

(7.15)

(7.16)

Table 7.7 Poisson’s Ratio of Polymethacrylates Used for Dental Applications Resin

Filler

BisGMA 1 TEGDMA BisGMA 1 TEGDMA BisGMA 1 UDMA 1 BisEMA BisGMA 1 TEGDMA

40% 66% 60% 47%

Poisson’s Ratio Colloidal silica 0.01
0.09 μm Zircone 1 silica 0.01
3.5 μm Zircone 1 silica 0.01
3.5 μm Zircone 1 silica 0.01
6 μm

0.372 0.302 0.308 0.393

7.4 Physicochemical, Biological and Mechanical Properties

The Poisson’s ratio (ν) in a composite can also be estimated from the volume fraction of fillers (φ), the Poisson’s ratio of matrix, and filler ν m and ν f (Halpin and Kardos, 1976): ν 5 ð1 2 φÞν m 1 φν f

(7.17)

Its value for matrices (0.3
0.35) increases at about 0.5 when T reaches Tg, that is, when the polymer turns from a glassy to rubbery state at which it is incompressible (Mott et al., 2008). In the case of filled dental composites (Chunga et al., 2004), values ranging from 0.3 to 0.4 are recorded at room temperature (Table 7.7). Higher values are observed for commercial composites. According to several authors, hardness and elastic modulus are well-correlated (Thomaidis et al., 2013; Pal, 2005). Li et al. (2009a) propose, for example, a linear correlation: EB0:15 3 Knoop Microhardness

(7.18)

E being expressed here in GPa. The positive effect of curing on the microhardness is illustrated in Haenel et al. (2015) whereas Li et al (Atai et al., 2012) also show a decrease in hardness with polymer thickness. Ultimate flexural strength has to display a value at least equal to 50 MPa for clinical requirements (as mentioned in Bindu et al., 2013). Typical values for dental composites are given in Table 7.8 (see Barszczewska-Rybarek, 2009). According to Eyring’s theory, plasticity originates in the jump of segments. This phenomenon is thermally activated (with a ΔH energy corresponding to the potential barrier) and facilitated by the activation volume vflow and the external _ stress σ. The yield stress σY is, thus, linked to the strain rate γ:   σy ΔH kT 5 2 Uln γ_ 0 γ_ vflow vflow 2

(7.19)

This equation is close to the experimental observations by Kambour, according to which:   σY 5 CU Tg 2 T 1 σY0

(7.20)

Table 7.8 Flexural Strength, Elastic Modulus, and Brinell Hardness for Some Unfilled Polymethacrylates Used for Dental Applications (BarszczewskaRybarek, 2009) Resin

σ f (MPa)

E (MPa)

HB (N mm22)

Poly(bis GMA) Poly(TEGDMA) Poly(UDMA) Poly(bisGMA-co-TEGDMA) Poly(bisGMA-co-TEGDMA-co-UDMA)

115 85 140 95 105

3800 3900 3500 4100 2800

75 135 165 90 190

241

242

CHAPTER 7 Polymethacrylates

C ranges from 0.5 to 1 MPa K21 (Cooke et al., 1998; Li and Strachan, 2011). A decrease in glass transition, thus, results in a decrease in yield stress. At temperatures above the glass transition, thermoset networks are in a rubbery state. Elastic behavior is given by the Flory approach according to which the Young’s modulus is proportional to the concentration in elastically active chains n0 (Mark, 1984): E5

3ρRT 5 3n0 RT MC

(7.21)

where ρ is the density, MC the average molar mass between crosslinks, R the gas constant, and T the absolute temperature. Values of rubbery modulus (measured at 175 C) of several bis-GMA 1 HEMA networks close to 25 MPa are given in Park et al. (2009). This means that the average molar mass between crosslink nodes is c. 400
500 g mol21 (i.e., n0 c. 2 mol kg21). This is the expected order of magnitude in these materials (see Table 7.3) since the HEMA comonomer contributes to an increase of the molar mass between crosslinks. The correlation between rubbery modulus and conversion degree is illustrated in the case UDMA by Sadoun and colleagues (Chi Phan et al., 2015). Identically to glassy modulus, it increases with filler content (Munhoz et al., 2017) as described by Guth (1945): E 5 E0 :ð1 1 2:5φ 1 14:1φ2 Þ

(7.22)

where 14.1 3 φ expresses the filler
filler interaction effect on elasticity and is particularly relevant for highly filled matrices such as dental composites. Even if Young’s modulus on the rubbery plateau is not itself helpful data for practitioners, it is noteworthy that it allows an estimation of the concentration in elastically active chains, expected to decrease during hydrolytic degradation (see Section 7.5.3). 2

7.4.1.3 Ultimate properties The toughness expresses the ability of a material to absorb energy and plastically deform without fracturing. According to the Griffith’s equation, the stress intensity factor in an infinite plate with a crack of 2a length is: KI 5 yUσUða:πÞ1=2

(7.23)

that is, that sample fails either if the stress σ, or the size of the crack a, exceed a critical value. PMMA toughness can be easily studied using a common tensile test sample. In the case of dental composites, various methods are proposed, for example, using notched disks (Watanabe et al., 2008) allowing to study failure in mode I or II. The values of toughness are shown to depend on the load rate, but stay close to 1.5 MPa m1/2 (Wada, 1992). These values are actually very close to unreinforced poly(UDMA) (Phan et al., 2014) and, more unexpectedly, in filler reinforced composites (Atai et al., 2012; Guth, 1945; Ornaghi et al., 2014).

7.4 Physicochemical, Biological and Mechanical Properties

Reversely, the toughness can reach about 2.5 MPa m1/2 for resins reinforced with 7.5% short glass fibers (Bocalon et al., 2016). The presence of rubbery fillers has a positive effect on toughness (Mante et al., 2010; Omran Alhareb et al., 2017). The impact strength is typically measured using a Charpy impact test on notched or unnotched samples. The typical value for PMMA is about 5 kJ m22 with possible improvements by reinforcing with various kinds of rubbery particles (NBR (Omran Alhareb et al., 2017) or poly(methyl methacrylate-b-butyl acrylateb-methyl methacrylate) (MAM) (Lalande et al., 2006). In unfilled poly(UDMA), poly(Bis-GMA), poly(TEGDMA), and their mixtures the value is higher and can reach c. 9 kJ m22 (Barszczewska-Rybarek, 2009). Lastly, it is noteworthy that the combined effect of fillers and low monomer viscosity lead to porosities (Balthazard et al., 2014) which are quite detrimental to the ultimate mechanical properties. It must also be highlighted that the difference between the thermal coefficient dilatation of the polymer filler induces stresses at the polymer/filler interface (Ferracane, 2005) which becomes an area of weakness where a crack will easily propagate and reduces the toughness of the material.

7.4.2 BIOCOMPATIBILITY The biocompatibility of dental resins may affect both the patient and the dentist. In the case of samples immersed in water, it can be observed that a part of the mass is lost presumably because of the migration of low molecular mass compounds (Sideridou and Karabela, 2011). Those phenomena are usually quantified by measuring the soluble fraction, being the relative mass decrease of a composite resin before immersion and after immersion and complete drying. If the results clearly depend on the curing process, it seems that this soluble part can represent from 0.1% to 1% by weight of the polymer mass in common light-cured dimethacrylate matrices (Ru¨ttermann et al., 2010; Sideridou et al., 2003). This quantity can even be higher if low molecular mass compounds produced from the partial hydrolytic or enzymatic degradation of networks, which will be addressed in the following of this chapter. As expected with uncompletely cured networks, a great part of soluble (eluted from resins) compounds contains unreacted monomers and photoinitiators (Munhoz et al., 2017; Ferracane, 2006). The migration of a chemical out of a polymer is, in great part, controlled by its molar mass (expressing the architecture) of compounds so that it is quite likely that dimers or trimers diffusion is slow enough to be neglected in the first approach. Quite interestingly, it was observed that the fraction of UDMA and Bis-GMA extracted from UDMA-TEGDMA or Bis-GMA/TEGDMA was systematically higher than their fraction in the polymerization mixture. Since their higher molar mass makes them less mobile than TEGDMA monomers, this suggests that BisGMA and UDMA are not randomly polymerized with TEGDMA (Floyd and Dickens, 2006) (see Section 7.3.1.2).

243

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CHAPTER 7 Polymethacrylates

The toxicity of several kinds of chemicals involved in the formation of dental composites was addressed by Thonemann et al. (2002). According to TC50 measurements performed on several kinds of cells, Bis-GMA seems by far to display the highest cytotoxicity compared to UDMA and MMA. Wear debris are produced from attrition and abrasion of resin composite dental restorations (Heintze, 2006). Ingestion of filler particles could result in potential harm to the liver, kidney, or intestine (Gatti and Rivasi, 2002; Gatti, 2004). Nevertheless there is no scientific evidence that swallowed particles induce a significant health risk for patients (Heintze, 2006). Besides this, polishing, shaping, and grinding composites result in particle dust (,5 μm particles including nanoparticles ,100 nm) which can be inhaled and penetrate the lungs (Van Landuyt et al., 2012), provoking cell toxicity in human bronchial epithelial cells (Cokic et al., 2016)—for exposure under relatively high particle concentrations compared to standard use. The fiber-reinforced composites were recently introduced and seem to offer a good combination of reinforcement of mechanical properties and low toxicity (Ballo and Na¨rhi, 2017).

7.5 LONG-TERM BEHAVIOR Let us recall that there are two kinds of ageing phenomena: •

•

Physical ageing, where the polymer backbone remains unmodified, whereas the free volume is changed by physical relaxation, the ingress of an external penetrant, or the loss of an adjuvant (typically a plasticizer). Chemical ageing where the polymer backbone or its lateral groups undergo chain scissions or crosslinking induced by any kind of chemical reaction.

In the specific case of polymeric dental materials, the main mechanisms were listed by Ferracane (2006). The most relevant ones for acrylates will be addressed next.

7.5.1 AGING BY PHYSICAL RELAXATION This kind of mechanism is common to every glassy polymer. When cooled from elevated temperatures, the specific volume decreases, but the decrease rate is below the Tg of the polymer. Lower than this temperature, the polymer can be considered as in an “out-of-equilibrium” state. The primary thermodynamic properties (volume, enthalpy, etc.) decreases slowly during the storage at an ageing temperature (Ta) lower than Tg (Fig. 7.16). Meanwhile, the necessary enthalpy for initiating rubbery phase motions is increased when reheating above Tg, generating the enthalpy overshoot commonly

7.5 Long-Term Behavior

Ageing range

Truly glassy state

Equilibrium line Tβ

Temperature

Tg

FIGURE 7.16 Mechanism of physical relaxation ageing. Tg is the glass transition temperature, Tβ the highest temperature of subglassy transition, and v is the specific volume (Struik, 1978).

observed by DSC (Diaz-Calleja et al., 1987) and a slight increase of Tg toward a “fictive temperature” (Tf) given by: Tf 5 Tg 2

Δh ΔcP

(7.24)

The enthalpy overshoot tends toward an asymptotical value given by:   ΔhN 5 ΔcP U Tg
Ta

(7.25)

where ΔcP 5 cPl
cPg is the heat capacity jump at Tg, and Ta is the temperature at which physical ageing occurs. From a practical point of view, the distance to equilibrium increases with decreasing temperature, but the rate decreases very strongly when temperature decreases, so that this ageing mechanism is typically significant in the temperature range [Tg; Tg 2 60 C] (Pethrick and Davis, 1998). Physical ageing is described by a distribution of relaxation times:  β

t φðtÞ 5 exp 2 τ

(7.26)

The function seems to depend on the experimental technique (e.g., dynamic modulus, dilatometry, calorimetry) used for measuring (Pe´rez et al., 1991). τ is the relaxation time that can be predicted by two models (Hodge, 1994): 1. Tool
Narayanaswamy
Moynihan model predicts the relaxation time from the value of “fictive” temperature (Narayanaswamy, 1971): τ 5 τ 0 exp





xΔh ð1 2 xÞΔh 1 RT R:Tf

(7.27)

2. Kovacs
Aklonis
Hutchinson
Ramos (KAHR) (Grassia and Simon, 2012; Grassia and D’Amore, 2011):

θδ τ 5 τ R exp½θðTR 2 T Þexp 2 ð1 2 xÞ Δη

(7.28)

245

246

CHAPTER 7 Polymethacrylates

where τ R is the relaxation time at the reference temperature TR, θ is linked to activation energy, x is an adjustable parameter ranging from 0 to 1, η is parameter specific to the value under study (i.e., depending on heat capacity [cP] for enthalpic relaxation, dilatation coefficient α for volumetric relaxation). The full description of these models is out of the scope of this chapter. Since ageing by physical relaxation results in a decrease of free volume quantity, all physical and mechanical properties related to free volume are changed, for example: • • • •

Yield stress is increased, possibly because the activation volume vflow (see Eyring’s equation) is linked to the residual free volume. Creep resistance (Hutchinson and Bucknall, 1980) Toughness (Arnold, 1995), Loss tangent is decreased in the domain of the subglassy transition (c. 320K in the case of PMMA) (Diaz-Calleja et al., 1987; Etienne et al., 2007) (see Fig. 7.17).

A final consequence is worth being investigated. It seems clear that physical ageing relaxation induces free volume collapse, i.e., a decrease in the size of nanovoids present in the polymer (Pethrick and Davis, 1998) with possible consequences on the water ingress in the composite.

7.5.2 HUMID AGEING The penetration of water is an environmental factor that can drastically limit the performance of polymers. Two subcases can be distinguished: • •

Water diffuses into the polymer, but does not change the polymer’s architecture. Water diffuses and reacts with the polymer, generating scission of lateral or skeletal bonds (Section 7.5.3).

7.5.2.1 Water solubility The mechanism of the polymer
water interaction can be, firstly, described by the shape of the sorption isotherm where the ratio of water in the polymer
water mixture is plotted by the function of water activity (or partial pressure), as illustrated in Fig. 7.18. Several shapes of sorption isotherms are described for the sorption of gases into polymers. In the case of water penetration, three main cases exist: 1. Henry’s law is the simplest theory for water dissolution. It is associated to a very dilute solution behavior in which dissolved water molecules are few and far between. It assumes that maximum (equilibrium) water uptake is directly proportional to the external water partial pressure: C 5 sUP

(7.29)

7.5 Long-Term Behavior

(A) 0.14 0.13

α relaxation Aging time: (1) 4 h (2) 24 h Physical aging (3) 48 h (1) (4) 72 h (2) (5) one week

tan (ϕ)

0.12 β relaxation

0.11

(3)

0.10 0.09 (5)

0.08 (4)

280

300

320

340

360

380

400

Temperature (K) (B) ΔCp = Cp(T)-Cp(320K) (J (g K)−1)

1.2 1.0 0.8

Aging time: (1) 4 h (2) 24 h (3) 48 h (4) 72 h (5) one week

(5) (4)

(3)

0.6 0.4

(2)

0.2

(1)

0.0

340

350

360

370 380 390 Temperature (K)

400

410

420

FIGURE 7.17 The effect of physical ageing on PMMA at 363 K after thermal annealing on: (A) loss factor; and (B) specific heat. Reused with permission of Elsevier.

where C is solubility expressed, for example, in cc(STP)/cc(polymer). s is the solubility coefficient (expressed in mol L21 Pa21) expected to obey Van’t Hoff law:
 ΔHS sðT Þ 5 s0 exp 2 RT

(7.30)

P is the water partial pressure (expressed in Pa) obeying Clapeyron’s law:
 ΔHvap PðT Þ 5 P0 exp 2 RT

(7.31)

247

248

CHAPTER 7 Polymethacrylates

FIGURE 7.18 Shape of sorption isotherms in polymers.

The apparent activation energy for water equilibrium concentration is, thus, given by: ES 5 2 ΔHS
ΔHvap

(7.32)

21

ΔHvap is close to 43 kJ mol . In PMMA, ΔHS B 243 kJ mol21 so that water uptake does not depend on temperature as in the first approach (Barrie and Machin, 1971). 2. An isotherm which is linear at low activity and displays an upturn at higher water pressures (such as B isotherm in Fig. 7.18) is associated to a type III isotherm in the BET classification and is sometimes called a Flory Huggins isotherm. This upturn has two explanations: a. Due to clustering of solvent molecules. b. Due to plasticization of the polymer matrix induced by solvent sorption. The mathematical description of Flory Huggins isotherms is given by: ln P=P0 5 lnð1 2 φp Þ 1 φp 1 χφ2p

(7.33)

where φp is the volume fraction of the polymer in the water
polymer mixture. is the Flory parameter describing the polymer
water affinity, expressed by: χ5

2 Vmwater   δpolymer 2δwater RT

(7.34)

where Vmwater is the molar volume of water, δpolymer, and δwater are, respectively, the solubility parameters of polymer and water, R the ideal gas constant and T the absolute temperature. In the case of PMMA-water association, both solubility parameters are known: δPMMA 5 19.0 MPa1/2 and δwater 5 47.9 MPa1/2. However, the resulting χ parameter seems to be an overestimation (5.91 at 35 C) compared to the value deduced from sorption isotherms (3.48). One possible explanation

7.5 Long-Term Behavior

is that this estimation of χ from the solubility parameter values is not refined enough for taking into account the various kinds of interactions (e.g., dispersive, dipole-dipole, hydrogens, etc.). Another explanation is the presence of water clusters, which can be observed by dielectric measurements at high water uptake (Garden and Pethrick, 2017). The clustering function was defined by Zimm and Lundberg (1956) as: fZL 5



  @a1 =φ1 G11 5 2 1 2 φ1  21 ν1 @a1 T;P

(7.35)

When fZL is below a value of
1, no clustering occurs. Means cluster size (MCS) is given by: MCS 5 1 1 φ1 UG11 =v11

(7.36)

It was, hence, shown (Davis and Elabd, 2013) that water molecules associate to form “dimers” when water external partial pressure (or activity) exceeds 0.2, which is typically the case when PMMA is immerged in water. 3. The Langmuir isotherm describes the equilibrium between the absorption (rate constant k1) and desorption (rate constant 5 k21) of molecules on a surface; the rate being propositional to water external partial pressure (P) and the concentration of sorbed water (c), as described by: k1 UPUð1
cÞ 5 k21 Uc

(7.37)

which can be reformulated under the general form: c5

Ap 1 1 Bp

(7.38)

In some cases, the sorption of a penetrant can be modeled by the dual sorption theory, which is the combination of Henry and Langmuir sorptions: c 5 sp 1

Ap 1 1 Bp

(7.39)

Some cases corresponding to this dual sorption mode are presented in Vieth et al. (1976). There are several values for water-uptake in methacrylates immerged in water (Sideridou et al., 2003, 2008; Delpino Gonzales et al., 2016; Smith and Schmitz, 1988). Some of these are provided in Table 7.9. The water affinity of a given polymer can be expressed as the number of water molecules absorbed per monomeric unit (Morel et al., 1985): H5

X wm M 5 ni Hi 1800

(7.40)

where M is the molar mass of the repetitive unit, ni is the number of groups able to bind with Hi molecules of water.

249

250

CHAPTER 7 Polymethacrylates

Table 7.9 Water Equilibrium Mass Uptake in Several Methacrylate Polymers PMMA BisGMA TEGDMA UDMA BisEMA D3MA



23 C 37 C 37 C 37 C 37 C 37 C

w/w

Reference

1.90% 3.57%
3.86% 5.74% 2 6.25% 2.39%
3.10% 1.92%
2.11% 0.65%
0.66%

Smith and Schmitz (1988) Sideridou et al. (2008) Sideridou et al. (2008) Sideridou et al. (2008) Sideridou et al. (2008) Sideridou et al. (2008)

Hi can be estimated from several polymers or chemicals (ideally with only one kind of functional group). For example, data for PMMA (Table 7.8) suggest Hester B 0.1. However, this simple theory can fail for several reasons: •

• •

The contribution of hydroxyl group (HOH) might be weaker than in linear polymers such as PVOH because of the possibility of intramolecular hydrogen bonds (Kalachandra and Kusy, 1991) as observed when OH are hydrogenbonded with heteroatoms (nitrogen atoms at the crosslink node for epoxy/ diamine networks (Morel et al., 1985), oxygen for Bis-GMA. Water can have a very specific interaction with other water molecules (clustering). This simple theory considers the presence of polar and apolar groups, but not their molecular arrangements and the subsequent possibility to create “complexes” with water molecules.

Lastly, Kerby et al. (2009) also observed that the water uptake value is correlated with water contact angle: hydrophilic samples display a lower contact angle than hydrophobic ones.

7.5.2.2 Water diffusion The diffusion of water in a polymer is expected to be described by the second Fick’s law, provided that diffusivity does not depend on the water concentration: @c @2 c 5 Dw  2 @t @r

(7.41)

This equation was solved analytically by Crank (1975), for example, in the case of an “infinite” plate having a 2e thickness:
 N mðtÞ2m0 8 X 1 D∙ð2n11Þ2 ∙π2 ∙t 512 2 ∙ ∙exp 2 mN 2 m0 π n50 ð2n11Þ2 4e2

(7.42)

It is easy to verify that Eq. (7.42) describes fairly the sorption curves displayed in Fig. 7.19A. Increasing the number of terms of the index n allows a better description of the curve in the earliest sorption times.

7.5 Long-Term Behavior

FIGURE 7.19 The general shape of mass uptake in a polymer in the presence of water in the case of a: (A) Fick Diffusion model; and (B) Langmuir diffusion model.

This equation can be simplified into: •

•

at low penetrant uptake (m/mN , 0,6):

rffiffiffiffiffiffiffi ΔmðtÞ 4 D∙t 5 ∙ ΔmN e π

(7.43)


 ΔmðtÞ 8 Dπ2 5 1 2 2 ∙exp 2 2 t e ΔmN π

(7.44)

at high penetrant uptake:

In other words, if mass uptake varies linearly with the square root of time at low mass uptake values, it means that diffusion obeys Fick’s law. The initial slope can, thus, be used to estimate the apparent diffusivity. It seems that the water diffusion into methacrylate-based polymers obeys Fick’s law (Sideridou and Karabela, 2011; Barrie and Machin, 1971; Dhanpal et al., 2009;

251

252

CHAPTER 7 Polymethacrylates

Table 7.10 Values of Diffusion Coefficient in Several Resins (Sideridou and Karabela, 2011; Barrie and Machin, 1971) T (K) PMMA

BisGMA BisEMA UDMA TEGDMA D3MA

313.7 323.4 333 343.2 310 310 310 310 310

D (cm2 s21)

Reference

28

5.2 3 10 8.9 3 1028 1.5 3 1027 2.4 3 1027 1.1 3 1027 0.74 3 1027 0.69 3 1027 0.15 3 1027 0.62 3 1027

Barrie and Machin (1971) Barrie and Machin (1971) Barrie and Machin (1971) Barrie and Machin (1971) Sideridou and Karabela (2011) Sideridou and Karabela (2011) Sideridou and Karabela (2011) Sideridou and Karabela (2011) Sideridou and Karabela (2011)

Sideridou et al., 2004). Some values are given in Table 7.10 that show differences between experiments performed in sorption mode compared to desorption mode (Sideridou and Karabela, 2011; Dhanpal et al., 2009; Sideridou et al., 2004). Since samples actually differ from infinite plates (length [L] and width [l] are finite values), Shen and Springer (1976) proposed a correction for taking into account the geometry of samples: Dre´ el 5 

Dapp 2e 11 2e L 1 l

2

(7.45)

D is expected to obey to Arrhenius law:
 ED DðT Þ 5 D0 :exp 2 RT

(7.46)

In PMMA, ED would be on the order of 45 kJ mol21 (Barrie and Machin, 1971), which seems to be quite common with other values reported for amorphous polymers in their glassy state (Li et al., 2009b). Values slightly lower (30
35 kJ mol21) are observed in some matrices of dental composites (Dhanpal et al., 2009). In some cases, the sorption curves display two plateaus which are attributed to the existence of polymeric sites inducing strong interactions with water molecules. Water molecules present in the polymer matrix divide into two parts: • •

Free water which diffuses. Bound water being in strong interaction with some specific sites of the polymer.

This Langmuir-type diffusion was mathematically described by Carter and Kibler (1979) by denoting: • •

n(t) the concentration in “free” water. N(t) the concentration in “bound” water.

7.5 Long-Term Behavior

• •

γ the probability that “free” water becomes “bound” water. β the probability that “bound” water becomes “free” water. The equilibrium is described by: γUnN 5 βUNN

(7.47)

The system is described via a system of differential equations: @n @N @2 n 1 5 D 2 @t @t @x

(7.48)

@N 5 γn 2 βN @t

(7.49)

By denoting: κ5

π2 D ð2eÞ2

(7.50)

Its solution can be approximated: •

•

at short absorption times:
 pffiffiffiffiffi NðtÞ 4 β 5 3=2   κt N β1γ π

(7.51)

NðtÞ γ 512 expð 2 βtÞ N β1γ

(7.52)

at high absorption times:

The “pseudo plateau” value can be approximated by: Npseudo equilibre β 5 N β1γ

(7.53)

However, such a process has not been reported, to the best of our knowledge, for the case of any methacrylate-based materials. The rate at which water diffuses is influenced by the penetrant architecture and polymer free volume, as proposed by Cohen and Turnbull (1959):   D 5 D0 Uexp 2b=Vf

(7.54)

where, b is related to penetrant size and Vf to polymer “empty” space allowing penetrant jumps. “Free volume” theory is, for example, well-illustrated in the case of epoxy networks when correlating the measured value of diffusivity (determined from the classical gravimetric method) with the volume of nanoholes (expected to be linked to free volume) (Frank and Wiggins, 2013). In the case of methacrylate-based polymers, it was observed that the diffusivity values in networks-based tetrafunctional methacrylates was lower than in analogous methacrylates, presumably because of the highly crosslinked nature of the networks and the lower free-volume content (Kalachandra and Kusy, 1991).

253

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CHAPTER 7 Polymethacrylates

Theories linking diffusivity with the capability of polymer segments to facilitate diffusion are also illustrated in polymers having a subglassy transition (also named β transition) where Halary observed a good correlation between the occurrence of this transition and the diffusivity of water (Halary, 2000). It seems, also, that the ageing by physical relaxation (see Section 7.5.1) can influence the overall-water ageing process, since the decrease in free volume is accompanied by an overall lower penetration of water, as illustrated by Siu-Wai Kong (1986) in the case of epoxies resins. Contrarily to theories considering diffusion being mainly influenced by the amount of free volume, Verdu and colleagues (Merdas et al., 2002) observed that for a family of epoxy resins, diffusivity was inversely correlated with solubility. Since this later mainly originated from polar groups (such as isopropanol in epoxies) inducing a strong interaction with water, they suggested that these later decrease the diffusivity. The following mechanism was proposed: [P1. . .W]-P1 1 W WW 1 P2-[P2. . .W]

Dissociation of water/polymer complex Jump of a water molecule from P1 to P2 Formation of a new water/polymer complex

This theory is verified in part by Dhanpal et al. (2009) for some dental composites, where the differences are few.

7.5.2.3 Consequences of physical ageing on mechanical properties Water is a small molecule characterized by a low Tg (Tg 5 136K for water; Angell et al., 1978). It is, thus, not surprising that the main consequence of water penetration in acrylate polymers is a decrease of glass transition temperature (Smith and Schmitz, 1988), which can be tentatively described by several models, such as: 1. The DiMarzio’s equation is the simplest law for describing the Tg depletion for a polymer in presence of a plasticizer: 1 w1 w2 5 1 Tg Tg1 Tg2

(7.55)

2. A thermodynamic theory was proposed by Couchmann and Karasz (ten Brinke et al., 1983): lnTg 5

w1 ΔCp1 lnTg1 1 w2 ΔCp2 lnTg2 w1 ΔCp1 1 w2 ΔCp2

(7.56)

where ΔCp is change of heat capacity at Tg for each component of the mixture equal to ΔCP1 5 1.94 J (g K)21 for PMMA and ΔCP2 5 0.318 J (g K)21 for water. 3. Ellis and Karasz (1984): Tg 5

w1 ΔCp1 Tg1 1 w2 ΔCp2 Tg2 w1 ΔCp1 1 w2 ΔCp2

(7.57)

7.5 Long-Term Behavior

4. Kelley and Bueche (1961): Tg 5

ΔαTg2 φ2 1 α1 Tg1 φ1 Δαφ2 1 α1 φ1

(7.58)

where α1 is the diluent coefficient of cubic expansion, and Δα is the change of coefficient of cubic expansion for the polymer at Tg for which it is considered as an universal value close to 4.8 3 1024 K21. The effect of water absorption in PMMA is illustrated on the depletion of elasticity modulus in the glassy state from 3700 to 3000 MPa when water fraction increases from 0% to 1.4% (Delpino Gonzales et al., 2016). Tg 5

f w1 ΔCp1 Tg1 1 w2 ΔCp2 Tg2 f w1 ΔCp1 1 w2 ΔCp2

(7.59)

where f is considered as the fraction of water active for reducing Tg (i.e., 1
f is the fraction of water absorbed in microvoids). In fully cured networks, water plasticization results in a decrease in compressive strength and microhardness (Ferracane et al., 1998; Kalliyana Krishnan et al., 1997). A good illustration is given in the case of Bis-GMA 1 HEMA networks in Park et al. (2009), the results of which are schematized in Fig. 7.20. According to the Kambour’s theory, yield stress depends on the value of Tg (see Section 7.4.1). A decrease in glass transition results in a decrease in yield stress as illustrated by Leger et al. (2013): according to these authors, changes in mechanical properties induced by solvent ingress leading to a given temperature decrease (e.g., 20 C) are equivalent to the mechanical properties decrease induced by the same temperature increase (i.e. 20 C).

FIGURE 7.20 Schematic thermomechanical behavior of virgin (full line) and water-aged (dashed line) bis-GMA 1 HEMA resin (in the case of physical ageing without chemical reactions).

255

256

CHAPTER 7 Polymethacrylates

It is worth noting that physical ageing is, in principle, characterized by an equilibrium state, i.e., physical properties are expected to plateau after a longexposure time. However, since dental composites are sometimes undercured, the decrease in Tg and subsequent macromolecular mobility increase may also induce an increase in conversion degree and later in some mechanical properties (Ferracane et al., 1998; Malacarne-Zanon et al., 2009).

7.5.2.4 Role of the interface A supplementary level of complexification is observed in composites. It was also observed that the polymer/fillers interface displays usually weaker properties than the “bulky” polymer matrix, as illustrated in numerous papers dealing with Tg measurements of organic matrices at the boundary with the interface (Tillman et al., 2002; Joliff et al., 2013). It is, thus, observed that water solubility and diffusivity in composites are higher than those predicted by a simple mixture law (i.e., assuming that water does not diffuse into the filler) as observed by Chateauminois et al. (1994) in the case of reinforced epoxies. This was also illustrated by Kalachandra (1989) in the case of PMMA filled with barium-based particles (Table 7.11). The detrimental effects of the interface can be partially attenuated by the use of proper coupling agents (such as 4-methacryloxyethyl trimellitic anhydride). Other examples are reported for the case of dimethacrylates used in dentistry where: •

•

Water solubility reaches 1.2%
1.7% even for composites with c. 75% weight fillers (Wei et al., 2011), that is, higher than expected from the water-uptake value for pure resin (see Table 7.8). The diffusivity is observed to be higher in composites than in the pure matrix (Dhanpal et al., 2009). Table 7.11 Water Sorption Parameter at 37 C in PMMA and Its Composites

Material PMMA PMMA 1 barium sulfate PMMA 1 barium sulfate 1 coupling agent PMMA 1 barium glass silance coated PMMA 1 barium glass silance coated 1 coupling agent

% Polymer

Diffusion Coefficient ( 3 108 cm2 s21)

100 31 30

% Water Uptake Calulated

Experimental

2.26 33.18 10.08

0.71 0.68

2.28 1.23 1.27

25

10.55

0.57

1.23

23

9.42

0.52

1.04

7.5 Long-Term Behavior

The water solubility and/or diffusivity are higher than in the pure matrix, thus, needing a much more complex modeling (Joliff et al., 2014). In terms of mechanical properties, these observations are consistent with the results obtained by Jager et al. (2016a) who found highly filled matrices to be the most sensitive to water permeation.

7.5.2.5 Effect of penetrant composition mixture It seems that, in a first approach, water has almost the same effect than various artificial salivas, as illustrated by Al-Mulla et al. (1989). Ethanol is also more soluble in dimethacrylates than water (Malacarne-Zanon et al., 2009), consistent with its lower solubility parameter (26.5 MPa1/2) than water. The effect of water and more acidic media (coke, orange juice) was compared by Rahim et al. (2012) showing that: • • •

Diffusion mechanism remains Fickian in every case. The water diffusion coefficient remains about constant. The maximum fluid uptake and the fraction of solubilized compounds increases. This last result can be explained due to better leaching of unreacted monomers and soluble materials and/or a possible hydrolysis reaction, which will be next.

7.5.3 CHEMICAL AGEING BY HYDROLYSIS A supplementary effect of water penetration is the possible hydrolysis of ester groups. In the case of biomaterials, these reactions were proposed: 1. In the case of PMMA where chain scission occur on side chains (Ayre et al., 2014; Ali et al., 2015; Semen and Lando, 1969; Du et al., 2006) (Scheme 7.9). The hydrolytic degradation is, thus, expected to firstly modify the polarity of the material (since each elementary reaction converts one ester group into a carboxylic acid group). 2. In the case of TEGDMA in the presence of enzymes such as cholesterol esterase (Ferracane, 2006; Finer and Santerre, 2004) (Scheme 7.10). In ideal networks (no dangling chains or loops), one has: n 5 n0
3s

CH3

CH3 CH2 C O

C

(7.60)

n

OMe

SCHEME 7.9 Mechanism of PMMA hydrolysis.

CH2 C O

C

n

OH

+ CH3OH

257

258

CHAPTER 7 Polymethacrylates

CH3

CH3 CH2 C O

C

H2O n

O

O

O

O

C

CE

O

C CH2

n

CH2 C O

C

HO

n

OH O

O

OH

CH3

SCHEME 7.10 Mechanism of TEGDMA hydrolysis.

FIGURE 7.21 Typical kinetic curves for various ageing processes involving water.

But, this relation is no longer valid when networks differ from the ideal (since each chain scission generates two dangling chains). The complete mathematical treatment is proposed in Richaud et al. (2014). In networks such as Bis-GMA, the ratio of ester bond needed for the “degelation” of the network (i.e., the total disappearance of elastically active chains and the possible solubilization of the material in a good solvent) is given by (Gilormini et al., 2014): 1 xd 5 1 2 qffiffiffiffiffiffiffiffiffiffiffi 1 1 L2

(7.61)

L being the number of reactive units (here hydrolysable ester groups) per elastic chain. Since carboxylic acids are efficient catalysts of the hydrolysis reaction, the hydrolysis degradation can display a certain autoaccelerated behavior (Richaud et al., 2014). The kinetic aspects of ageing are depicted in Fig. 7.21.

7.5 Long-Term Behavior

Lastly, it must be mentioned that the kinetics of water ageing may be controlled by the rate of water diffusion from the edge to the bulk (El Yagoubi et al., 2012). This means that the general equation for water consumption in any layer of the polymer is given by (Jacques et al., 2002): @w @2 w 5 Dw  2 2 kh we @t @x

(7.62)

where Dw is the water diffusivity, kh is the rate constant for hydrolysis, e and w are, respectively, the concentrations in water and in hydrolysable groups (namely ester). This equation can be adapted for several practical cases (e.g., the existence of several kinds of reactive groups, or the possibility of noncatalyzed and catalyzed hydrolysis), but in the simplest case, the thickness of degraded layer can be approximated by: rffiffiffiffiffiffiffiffiffiffiffiffiffi Dw ws zdegraded 5 kh

(7.63)

In other words, the final properties of the polymer correspond to the average of the undegraded bulk and the degraded edges.

7.5.4 CHEMICAL AGEING BY RADIOLYSIS Radiolytic processes are involved in the sterilization of some polymers before implantation (Barton et al., 2013). Results obtained on PMMA unambiguously show the decrease of molar mass, which is ascribed to the mechanisms shown in Scheme 7.11. The number of scissions induced by an irradiation dose D per unit mass is given by: s 5 GðsÞD

(7.64)

where G(s) denotes the radiochemical yield of PMMA (in mol J21). The radiochemical yield G(s) should be close to 1
2 3 1027 mol J21 in PMMA (Schnabel, 1978; Thominette et al., 1989; Babu et al., 1984; Charlesby and Moore, 1964). Even if it is not of clinical interest, the case of the radiolysis of networks obtained from dimethacrylate remains an open question since the reactivity of esters (i.e., the radiochemical yield for chain scission) present in dimethacrylate might differ from the reactivity of esters present in PMMA (Gilormini et al., 2017). CH3

CH3 CH2 O

C C

CH2 OMe

C O

C

CH3 CH2

OMe

O

C ° C°

CH3 CH2

OMe

SCHEME 7.11 The mechanisms of chain scission in PMMA.

C O

C

OMe

CH3

CH3 CH2

C

CH2

°C O

C

OMe

259

260

CHAPTER 7 Polymethacrylates

7.5.5 CREEP AND FATIGUE Let us first recall that there are two kinds of fatigue experiments: 1. Those made on unnotched samples: they lead to Wohler curves where mechanical stress is plotted versus the number of cycles. The Wohler curves of bulk- and reinforced-PMMA are compared in Baleani et al. (2003). The report shows that filler would have a detrimental effect on fatigue resistance. 2. Those made on notched samples (fatigue crack propagation): they lead to the plot of crack growth rate (da/dN) versus the intensity factor (ΔK 5 Δσaπ1/2f 5 Kmax 2 Kmin, Kmax, and Kmin are defined as the maximum and minimum stress intensity experienced by the crack and f being a geometric factor). The classical curve for fatigue crack propagation is depicted in Fig. 7.22. The left side of the curve leads to the minimal value of stress intensity factor (also named fatigue crack inception or threshold ΔKi or ΔKth) leading to crack propagation. Higher values of ΔKi correspond to tougher samples. ΔKi is slightly higher than 0.1 MPa m1/2 in PMMA (Ramsteiner and Armbrust, 2001) and c. five-times higher for composites used in dental restoration (Shah et al., 2009). The right side of the curve is a vertical asymptote reached at a stress intensity value KC close to the sample toughness (see Section 7.4.1) and linked to the critical size of a defect leading to brittle failure: 1 aC 5  π

rffiffiffiffiffiffiffi KC Δσ

FIGURE 7.22 Schematic results of fatigue crack propagation experiment.

(7.65)

7.5 Long-Term Behavior

In the intermediary domain, a linear dependence between crack growth rate and stress intensity factor is observed, which is described by the Paris law (Paris et al., 1961): da 5 CΔK m dN

(7.66)

In the case of thermoplastic polymers, the positive effect of average molar mass is illustrated in Skibo et al. (1977). The most probable explanation is that longer polymer chains produce effective entanglements and better fatigue resistance of the crazes. Fig. 7.23 illustrates a comparison of crack growth rate on PMMA aged in air and that grown in a Ringer solution (Ayre et al., 2014) and illustrates the complexities of the effect of ageing on fatigue properties. m would be on the order of 5 for hydrolytically aged dental composites made of Bis-GMA, UDMA, Bis-EMA, and a small amount of TEGDMA, together with 60% of silica and zirconia fillers. Moreover, SEM observations show that the cracks propagate at the particle matrix interface and cause possible matrix-filler debonding (Shah et al., 2009). From a practical point of view, it is noteworthy that polymers are bad thermal conductive materials and that a high frequency for fatigue test can induce significant self-heating. Temperature can for example exceed 100 C in a PMMA submitted to a 50 Hz cyclic stress, meaning that sample reaches is rubbery domain and sample fails at a very low number of cycles (Justice and Schultz, 1980).

–5.0

Log da/dn (m cycle–1)

–5.5 –6.0 –6.5 –7.0 –7.5 –8.0

Cemex air Cemex Ringers Palacos R air Palacos R Ringers

–8.5 –9.0 –9.5 –10.0 2.4

2.5

2.6

2.7

2.8

2.9

3

3.1

3.2

Log ΔK (Pa√m)

FIGURE 7.23 Crack growth rate of a PMMA bone cement aged in air or in Ringer’s solution Reused with permission of Elsevier.

261

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CHAPTER 7 Polymethacrylates

7.6 CONCLUSION AND PROSPECTS FOR THE FUTURE OF THESE MATERIALS PMMA and methacrylates polymers are easily obtained by radical in-chain polymerization which makes composites easy to use for surgeons and dentists as biomaterials and matrices. In this chapter, we presented various methods for synthesis and the structureproperties relationships aimed for designing materials. However, they are sensitive, in particular, to water resulting that the initial properties cannot be maintained in vivo. The main degradation mechanisms expected to occur at moderate temperature, during use or sterilization by irradiation were presented and might be helpful to avoid complications induced by longterm ageing. One of the scientific challenges remains to take into account the presence of fillers complicating the lifetime prediction. In the future, researches should pursue improving the dispersion of fillers and the development of materials that offer a good compromise between low viscosity and improved mechanical properties. The properties could also be improved by new polymerization strategies involving, for example, enhanced hydrostatic pressure to limit the presence of porosities which are known to dramatically decrease the mechanical properties of composites and increasing composites’ water uptake (Soles et al., 2000).

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Thermoset polymethacrylate-based materials for dental applications

8

Muhammad Hassan1, Mehmood Asghar2, Shahab Ud Din3 and Muhammad Sohail Zafar4,5 1

College of Dentistry, University of Lahore, Lahore, Pakistan 2National University of Medical Sciences (NUMS), The Mall, Rawalpindi, Pakistan 3Shaheed Zulfiqar Ali Bhutto Medical University/Pakistan Institute of Medical Sciences, Islamabad, Pakistan 4Department of Restorative Dentistry, College of Dentistry, Taibah University, Al Madinah, Saudi Arabia 5 Department of Dental Materials, Islamic International Dental College, Riphah International University, Islamabad, Pakistan

8.1 INTRODUCTION The idea of replacing missing teeth with the help of synthetic or natural materials is not uncommon for mankind. Various prostheses have been used for a long period of time for the replacement of lost natural teeth (Johnson, 1959). Before the 17th century, materials such as wood, ivory, and the bones of hippopotami or whales were used for the fabrication of denture bases that were carved to fit the spaces in edentulous regions. The artificial teeth were attached to the denture bases with the help of metallic wires. However, these materials were hardly available and were quiet expensive. As a result, these were only affordable by rich patients (Khindria et al., 2009).

8.1.1 GOLD In 1775, a French dentist Etienne Bourdet first used gold to make denture bases that contained small holes similar to the sockets of a tooth. These denture bases did not rest accurately on the alveolar ridge and the tissue surfaces of these dentures were made in the form of a cup. However, the high cost and poor esthetics of these denture bases prevented their widespread use in dentistry (Murray and Darvell, 1993).

Materials for Biomedical Engineering: Thermoset and Thermoplastic Polymers. DOI: https://doi.org/10.1016/B978-0-12-816874-5.00008-6 © 2019 Elsevier Inc. All rights reserved.

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8.1.2 PORCELAIN By the 18th century, newer materials had been introduced for the fabrication of dentures. In 1774, De Chemant prepared the first set of dentures made with porcelain (Murray and Darvell, 1993). However, the brittleness, heavy weight, and absence of natural appearance of the porcelain dentures were major problems that limited its widespread usage as a denture base material (DBM) (Khindria et al., 2009).

8.1.3 VULCANITE In 1839, Charles Goodyear discovered vulcanized rubber which was formed by the reaction of a natural rubber with sulfur (Alla et al., 2015) (Fig. 8.1). The introduction of vulcanite as a DBM led to a significant reduction in the cost of dentures, making them affordable to the general population. Vulcanite, due to its favorable properties, such as dimensional stability, comfort, light weight, ease of preparation, and economy, was able to successfully replace previously used DBMs (Murray and Darvell, 1993). However, its properties were still far from ideal, for instance, its poor esthetics and lack of chemical bonding with the porcelain teeth restricted its long term use (Johnson, 1959).

8.1.4 ALUMINUM In 1867, Bean invented the casting machine and performed a casting of the first denture base made from an aluminum alloy (Tandon et al., 2010). Similarly, in

FIGURE 8.1 Structure of vulcanized rubber polymer with disulfide crosslinks. Adapted from Alla, R.K., Raghavendra Swamy, K., Vyas, R., Konakanchi, A., 2015. Conventional and contemporary polymers for the fabrication of denture prosthesis: Part I
Overview, composition and properties. Int. J. Appl. Dent. Sci. 1, 82
89.

8.1 Introduction

1888, Carroll presented a method of casting aluminum denture bases under pressure. Despite the light weight and accuracy of fit of these dentures, the high cost of their fabrication and difficulty in relining the aluminum denture bases discouraged their use in dentistry (Khindria et al., 2009).

8.1.5 CELLULOID In 1870, Circa introduced celluloid, which was based on a polymer that was made by plasticizing cellulose nitrate and camphor. One of the favorable properties of celluloid was its ability to be stained pink with the help of pigments, so as to match the color of gums and oral mucosa. Although celluloid was deemed as an excellent replacement for vulcanite rubber, it was observed that the material soon lost its pink color resulting in the absorption of stains from food products and drinks. In addition, many patients complained about the residual taste of camphor while wearing these dentures. As a result, the popularity of this material soon faded and its use was gradually discontinued (Rueggeberg, 2002).

8.1.6 BAKELITE In 1909, phenol formaldehyde, also known as Bakelite, was introduced by Leo Bakeland, which was commercially used for dental purposes in 1924. Although the material possessed excellent esthetics, its inherent brittleness, potential for staining as well as difficulty to fabricate and repair were major factors that prevented its frequent use for denture fabrication (Khindria et al., 2009).

8.1.7 POLYVINYL CHLORIDE In 1930, polyvinyl chloride (PVC), which is a copolymer of acetate and vinyl chloride (Fig. 8.2), was introduced for denture fabrication. However, the material did not gain sufficient popularity due to its inherent mechanical weakness caused by the development of residual stresses during fabrication, in addition to its tendency to discolor after exposure to hot foods and liquids (Young, 2010).

Based on polymerization method

Heat cured

Chemically cured

Light cured

Microwave cured

FIGURE 8.2 Classification PMMA denture base resins based on method of polymerization.

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8.1.8 BASE METAL ALLOYS The use of base metal alloys (such as nickel-chromium and cobalt-chromium) for denture fabrication dates back to 1907. However, the first use of metal alloys for the fabrication of dentures was documented in 1937 (Khindria et al., 2009). These materials gained popularity as materials of choice for the fabrication of dentures due to their strength, light weight, and low cost. However, dentures fabricated from these alloys were difficult to repair, had a potential for causing allergic and cytotoxic reactions in susceptible individuals, and were prone to corrosion and tarnishing, which resulted in the poor esthetics of the prostheses. As a result, research for the invention of a perfect DBM continued (Khindria et al., 2009).

8.2 POLY(METHYL METHACRYLATE) AS A DENTURE BASE Poly(methyl methacrylate) (PMMA), which is a colorless and odorless polymer of acrylic acid, was first introduced by Rom and Has in the form of transparent sheets, while its powder form was introduced in 1937 by Nemours (Tandon et al., 2010). Later on (1937), Walter Wright first introduced PMMA as a DBM and by 1946 it became one of the most commonly used materials for the fabrication of dentures. During this time, about 95% of dentures worldwide were being produced from PMMA (Peyton, 1975; Nogueira et al., 1999) due to its favorable characteristics, such as its ease of processing and pigmenting (Johnson, 1959), adequate mechanical properties, economy, and relatively low toxicity (Khindria et al., 2009). Despite being widely used for the fabrication of denture bases, PMMA does not fulfill all the requirements of an ideal DBM with reference to its physical and mechanical properties. This is due to its susceptibility to fracture under cyclic loading (Narva et al., 2005) and its tendency to absorb water, which can lead to a reduction in its mechanical properties (Matinlinna et al., 2004; Takahashi et al., 2013). Attempts at overcoming the inherent drawbacks of PMMA through chemical modification or reinforcement with various metallic or fibrous inserts have met with varied results. Among them, various forms of GFs are most commonly being used for the reinforcement of PMMA dentures as a means of improving their durability and strength (Garoushi et al., 2009) (see Section 8.6.3.4 for details).

8.2.1 CLASSIFICATION OF PMMA RESINS 8.2.1.1 According to the ISO standards The ISO standard 20795-1 (2013) for denture base resins classifies denture base resins into several categories, which are discussed here and summarized in Table 8.1.

8.2 Poly(Methyl Methacrylate) as a Denture Base

Table 8.1 Classification of PMMA Resins According to ISO 20795-1 2013 Type

Class

Description

1 1 2 2 3 4 5

1 2 1 2

-

Heat processing (powder and liquid) Heat processing (plastic cakes) Autopolymerizing (powder and liquid) Autopolymerized (powder and liquid, pour type) Thermoplastic Light activated Microwave curing

8.2.1.2 According to method of polymerization Based on their method of polymerization, PMMA resins can be classified as heat, chemical, light, and microwave cured polymers (Fig. 8.2).

Heat cured PMMA Heat cured resins consist of a benzoyl peroxide (BPO) initiator that is activated with the application of thermal energy. The BPO initiator dissociates into carbon dioxide (CO2) and free radicals are generated when the temperature is raised to about 60 C. The thermal energy required for the activation of the initiator is commonly provided using a water bath (Alla et al., 2015). The typical composition of PMMA denture base resins is given in Table 8.1. Polymerization stages of heat cured PMMA The formation of PMMA involves the addition polymerization of multiple methyl methacrylate (MMA) molecules. The polymerization reactions start from an active center where MMA molecules are added sequentially to the end of a growing chain. The steps involved in the addition polymerization of PMMA denture base resins are: Initiation and activation To initiate an addition polymerization reaction, a source of free radicals is required that are provided by an initiator. The initiator, usually BPO, dissociates upon the provision of thermal, chemical, or light energy and generates free radicals in a process called activation. These free radicals react with the CQC of the methyl methacrylate monomer, resulting in the formation of a single C
C bond and another free radical at the advancing end of the chain is created. This stage is physically characterized by the mixing of PMMA beads with the MMA monomer, and the mixed polymer appears as a grainy or sandy mass. Propagation During this stage, the free radical that forms at the growing end of the polymer chain reacts with another monomer molecule, resulting in the extension of the polymer chain as well as the formation of another free radical that reacts with another MMA molecule. This chain reaction continues as long as monomer molecules are available to react with the free radicals at the growing side of the polymer chain. Since the polymer chains are actively growing in this

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phase, the polymer physically becomes sticky or stringy during this stage. In some cases, the transfer of free radicals from the growing chain to a dead polymer also occurs. This results in the termination of the growing chain and the formation of free radicals alongside the polymer somewhere in the middle of the chain. The growth of a new polymer chain then starts at the site of the attachment of the free radical, resulting in the branching of the polymer molecule. Termination The propagation of the polymerization reaction continues until the supply of MMA molecules is exhausted. However, the termination of the polymerization of reaction occurs more often as a result of the formation of a dead chain, which can occur due to: (1) the transfer of a growing chain to a dead chain; (2) direct coupling between two free radical polymer chains; (3) the exchange of hydrogen (H) atoms from one chain to another; (4) the reaction of a growing polymer chain with the initiator or an impurity (McCabe and Walls, 2013). The use of a correct water/powder ratio, which is 3
3.5/1 v/v or 2.5/1 w/v, is essential during mixing to ensure that no faults occur during the processing of the resin. If too little liquid is used for mixing, all the polymer beads will not be completely wet, resulting in a polymer that will have a granular texture. In cases where excess liquid is used, a higher level of polymerization shrinkage will be observed. Upon mixing of the powder and liquid, the resultant mass passes through five distinctive physical stages, namely the sandy, stringy, doughy, rubbery, and the stiff stage. The sandy stage This stage appears immediately after the mixing of the polymer and monomer components. During this stage, almost no chemical interaction between the polymer and the monomer occurs, and the mixture appears as grainy or coarse. The stringy stage During this stage, the individual polymer beads are attacked by the monomer molecules. The smaller polymer beads get completely dissolved within the monomer and their chains get dispersed within the liquid phase. At the same time, the polymer chains in the larger beads start to uncoil and thus increase the viscosity of the mixture. This stage is characterized by the stickiness or stringy appearance of the mix when it is touched or drawn apart (Skinner, 1951). The doughy stage When an increased number of polymer chains enter the monomer solution, the mix enters into a doughy stage. A large number of swollen, but undissolved polymer chains still remain in the solution. At this stage, the mixed mass does not adhere to the sides of mixing vessel and it loses its stringiness. Clinically, this is the ideal stage for packing the mixed mass into a dental flask for curing (Skinner, 1951). The rubbery stage At this stage, almost all the monomers have been converted into polymer, or have been consumed either as a result of absorption by the growing polymer chains or due to their evaporation. The mass becomes rubbery and rebounds when compressed or extended. At this stage, the PMMA can no longer be used for compression molding. The stiff stage When the mixture is allowed to stand for an extended duration, continued chemical reaction as well as evaporation of the monomer results in the

8.2 Poly(Methyl Methacrylate) as a Denture Base

mixture, which then becomes stiff and dry in appearance and tends to resist mechanical deformation.

The compression molding technique To fabricate a denture base using the compression molding or conventional technique, a mold must first be prepared in gypsum. This involves the careful selection and alignment of the artificial teeth in the correct occlusal scheme in a wax pattern that is made over the trial denture base. This is done by making an impression, registration of the jaw relations, and articulation along with the setup of teeth. The completed trial base is invested into the master cast. Once the gypsum inside the flask has set, the wax is removed from the flask by immersing it in boiling water, followed by thorough cleaning of the master cast and the application of a thin layer of a separating medium, such as sodium alginate. The separating medium is then left in open air to dry and after some time, a thin layer of calcium alginate forms over the gypsum mold. The separating medium prevents the accidental bonding of the plaster and the prepared resin. Mixing of the powder and liquid is carried out in the correct proportions and the mixture is allowed to stand until it becomes doughy. The time required by the mixture to attain the doughy stage is termed as the dough-forming time and it is dependent on several factors, such as the particle size and the molecular weight of the polymers, the incorporation of a plasticizer or crosslinking agent, the ambient temperature, and the polymer/monomer ratio used for mixing (Alla et al., 2015). The pressed denture flask is then heated in a water bath or an oven under a precisely controlled temperature. In order to ensure that the acrylic is properly cured and to prevent uncured monomer being left behind, the temperature is gradually increased. A high level of porosity can result in cured resin if the flask is immersed directly into boiling water, because the monomer boils at 100.3 C and evaporates before the polymerization process is complete (McCabe and Walls, 2013). Compression molding is the most commonly used method for the processing of heat cured acrylic resins due to the relative ease and low cost of processing involved in this technique (Nogueira et al., 1999).

The injection molding technique Denture bases can also be fabricated using the injection molding technique in which a specially designed denture flask is used. In this process, half of the flask is filled with freshly mixed dental stone which contains the master cast. Sprue holes and a vent hole, which serve as a medium for the entrance of the DBM and an escape for hot gasses respectively, are attached to the wax pattern before the corresponding half of the denture flask is filled with gypsum and placed over the first half to complete the investment procedure (McCabe and Walls, 2013). Once the stone has set, the flask is opened and the wax pattern is removed with the help of a wax solvent (Anusavice et al., 2012). The heat cured PMMA is

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then added into the flask through injection and the heating process is initiated. Acrylic is continuously added to the flask during heating in order to compensate for the polymerization shrinkage. After polymerization, the denture is removed from the flask and then finished in a similar way to that used in compression molding (Skinner, 1951; Anusavice et al., 2012). Polymerization cycles Different polymerization cycles are used for the curing of heat cured acrylic resins: (1) constant heating of the water bath at 74 C for 8 hours or longer without any terminal boiling treatment; (2) the temperature of the water bath is maintained at 74 C for 8 hours followed by terminal boiling for 1 hour; and (3) heating the denture base at 74 C for 3 hours and then boiling it at 100 C for another hour (Anusavice et al., 2012). Following polymerization, the denture flask should be allowed to cool gradually till it attains room temperature. If rapid cooling is performed, a distortion of the denture base can occur due to a mismatch between the thermal contraction of the PMMA and the investment material. Ideally, the flask should be removed from the water bath and bench cooled for 30 minutes followed by immersion in tap water for 15 minutes before the flask is opened (Anusavice et al., 2012).

Chemically cured PMMA These are also known as autopolymerizing resins and have a slightly different composition and mode of activation in comparison to heat cured resins. This is because unlike with heat cured materials in which the polymerization reaction is activated by the application of thermal energy, chemically cured PMMA resins contain a tertiary amine initiator, usually dimethyl-p-toluidine (Alla et al., 2015), which chemically reacts with BPO and generates free radicals that are required for the initiation of the polymerization reaction (McCabe and Walls, 2013). Once the polymerization process has progressed, the reaction is then similar to that of heat cured resins. Chemically cured PMMA denture base resins do not attain the same degree of polymerization as heat activated materials. As a result, they possess inferior mechanical properties, such as flexural strength (FS), impact strength, and hardness (Tandon et al., 2010). In addition, they also leach out a higher amount of residual monomer and possess a higher degree of water sorption (WS) and solubility when compared with heat cured PMMA resins (Anusavice et al., 2012). An advantage of using chemically cured resins is that the dentures constructed from them show better adaptation and dimensional stability in comparison to heat cured PMMA. This is because cold cured DBMs show lesser thermal shrinkage during polymerization in comparison to heat cured materials (Tandon et al., 2010; Anusavice et al., 2012; Alla et al., 2015) due to incomplete polymerization (Alla et al., 2015). However, these materials show poor mechanical and physical properties, in addition to the yellowish discoloration of the resin over time due to the oxidation of the amine initiator (Alla et al., 2015). As a result, these cold cure resins are not used frequently for the fabrication of dentures. Instead, they are

8.2 Poly(Methyl Methacrylate) as a Denture Base

used for denture repairs and the fabrication of custom trays and removable orthodontic appliances (Anusavice et al., 2012).

Light cured PMMA In light cured PMMA resins, polymerization is initiated with the help of a photosensitive agent, such as camphorquinone, which generates free radicals when it is exposed to a visible light source. The composition of light curing PMMA resins is different to other types of resins. The constituents of light curing resins include: • • • •

Urethane dimethacrylate matrix Micro-fine silica High molecular weight acrylic resin monomers PMMA acrylic beads (fillers)

Light curing resins are supplied in premixed, light-proof pouches. Since light may not be able to penetrate inside the conventional investment media to polymerize the resin, light curing DBMs cannot be polymerized with the conventional investment methods. Instead, light curing resin is carefully molded over the cast followed by positioning the teeth over the cast and finally curing the resin by exposing it to a visible light source for an appropriate duration. The prepared denture is then retrieved from the cast and finished with conventional finishing procedures (Anusavice et al., 2012). Light cured DBMs provide easier fabrication of the denture and can be easily adjusted prior to polymerization. In addition, they tend to have a lesser degree of residual monomer and polymerization shrinkage in comparison to heat and chemically cured materials. However, light cured resins cannot be processed using the conventional polymerization technique and possess a limited depth of cure, which makes their processing difficult, technique sensitive, and expensive. As a result, these materials are not commonly used for the preparation of dentures (Tandon et al., 2010).

Microwave curing PMMA Heat cured PMMA DBMs can also be polymerized with the application of microwave energy. For this type of polymerization, a nonmetallic denture flask must be used, in addition to a denture base resin that has been specifically produced for microwave curing. The energy generated from a microwave is used to propagate the polymerization reaction within the DBM. One of the advantages of microwave cured denture base resins is the speed with which the resin can be cured in a shorter duration. In addition, various studies have also shown that the dimensional accuracy and the physical properties of these resins are similar to those of conventionally heat cured resins. On the other hand, microwave curing PMMA resins cannot be polymerized using conventional techniques and require special nonmetallic flasks for polymerization. Moreover, microwave cured resins demonstrate lesser bond strength to the acrylic teeth

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compared to chemically or heat activated PMMA DBMs. As a result, their use in dentistry is limited. From this discussion, it can be concluded that the properties of the previously mentioned polymers (i.e., heat, chemical, light, and microwave cured PMMA) are quite different from each other essentially due to the differences in their composition and processing methods and they, therefore, have different applications in dentistry.

8.3 PROPERTIES OF PMMA DENTURE BASE RESINS Some of the properties of heat cured PMMA are shown in Table. 8.2.

8.3.1 FLEXURAL STRENGTH FS, also known as transverse strength or modulus of rupture, is the measure of the strength of a bar under a static load that is supported on either end by lower supports. According to the ISO 20795-1 (2013) for denture base resins, a threepoint bending test is required to measure the FS of a specimen (Fig. 8.3). Properties of heat cured PMMA resins are summarized in Table 8.3. FS is a combination of the compressive, tensile, and shear stresses, and measures the resistance of the denture base against fracture caused as a result of bending. A high FS is important in ensuring the success and durability of a denture because the application of uneven forces and stress on the alveolar bone can cause the spontaneous fracture of dentures due to bending. Table 8.2 Properties of Heat Cured PMMA (Alla et al., 2015) Property of PMMA Solubility

Value In hydrocarbons In water

Sorption Modulus of elasticity Proportional limit Compressive strength Tensile strength Elongation (%) Impact strength Hardness Fatigue strength Thermal conductivity Coefficient of thermal expansion Glass transition temperature

0.04 g cm22 0.02 mg cm22 0.69 mg cm22 3.8 3 103 MPa 26 M Pa 76 MPa 48
62 MPa 1%
2% 0.98
1.27 J 18
20 KHN 1.5 3 106 cycles at 17.2 MPa 5.7 3 1024 C cm21 81 3 1026/ C 125 C

8.3 Properties of PMMA Denture Base Resins

Force

Specimen

FIGURE 8.3 Schematic representation of three-point bending apparatus for measurement of flexural strength in denture base materials.

Table 8.3 ISO 20795-Determined Limits for Sorption and Solubility for Denture Base Resins (2013) Type of Denture Base Resin

Max. Limit for Sorption (µg mm23)

Max. Limit for Solubility (µg mm23)

Type 1, 3, 4, 5 Type 2

32 32

1.6 8

Various authors have measured the FS of PMMA denture base resins and provided different results regarding this property. Khan et al. (1987) evaluated the transverse strength of heat cured PMMA and a light cured dental resin. Specimens measuring 75 3 25 3 3 mm (n 5 4) were prepared using a gypsum mold. A three-point bending test was performed using a universal mechanical tester (crosshead speed; 20 mm min21). The results pointed out that the light cured resin (Triad VLC) showed a significantly higher FS (125.23 6 14.70 MPa) than the heat cured (Estron) denture base resin (97.67 6 8.83 MPa). However, they tested a small number of specimens (n 5 4), used a very high crosshead speed (20 mm min21) for fracturing the specimens, and did not mention the distance between the supports of the testing rig, which is very important in measuring the FS of a material. The results of this study were contradicted by Machado et al. (2007) who prepared heat and light cured acrylic specimens measuring 50 3 25 3 2.5 mm (n 5 10) by investing brass blocks in a gypsum mold. Lucitone 199 specimens were then heat cured at 74 C for 9 hours followed by terminal boiling for 30 minutes. The light cured specimens were cured for 5 and 10 minutes respectively. The FS was measured by performing a three-point bending test, where the span length between the lower supports was 40 mm and the crosshead speed was 5 mm min21. The heat cured Lucitone had an improved FS (87.12 6 8.08 MPa) in comparison to the VLC triad resin (57.96 6 7.31 MPa). The results of the FSs of

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the specimens were correlated with the help of scanning electron microscope (SEM) imaging; which showed that the heat cured Lucitone 199 specimens had higher FSs because they underwent significant deformation before fracturing, while the VLC Triad specimens showed high levels of porosities that appeared as crater-like cavities on the surfaces of the specimens. As a result, they demonstrated lower FSs in comparison to heat cured Lucitone 199. Dixon et al. (1991) compared the FSs of dry and wet specimens of a high impact strength resin (Lucitone 199), a rapid polymerizing resin (Accelar 20), and a light polymerized resin (Triad) measuring 65 3 10 3 3 mm (n 5 25). The heat curing specimens were prepared in gypsum molds, while the light cured specimens were fabricated with stainless steel molds and cured for 10 minutes. No information was provided regarding the intensity or wavelength of the light source. Lucitone 199 showed the highest FS (96.26 6 5.76), followed by Accelar 20 (81.98 6 10.90) and Triad (49.96 6 6.98). Additionally, the FS of Lucitone was significantly reduced following immersion for 30 days (81.06 6 5.20), while there was no significant decrease in the FS of Accelar 20 following 30 days of immersion (78.98 6 11.90). Alkhatib et al. (1989) studied the FSs as well as the porosity of three denture base resins with different specimen thicknesses (3, 6, 11.6, and 17.7 mm) and curing cycles. Two of the resins were microwave polymerized, while the third was polymerized conventionally using a water bath. It was also observed that the heat cured specimens tested for all four thicknesses were free of porosity, while only one microwave cured resin was porosity-free for all of the four different tested thicknesses. Porosity in the microwave cured specimens tended to increase when higher energy was used for curing with a shorter duration of time. Pfeiffer et al. (2005) demonstrated the flexural properties of a microwave cured polyurethane based resin (Microbase) and compared it with chemically modified thermoplastic resins (Polyan and Promysan), and heat cured resins (Paladon 65 and Sinomer). Specimens measuring 65 3 10 3 3.3 mm (n 5 5) were used. A three-point bending test was performed on a universal testing machine (UTM) following the guidelines of ISO 1567 (1999), where the span length was 50 mm and the speed of the crosshead was 5 mm min21. The results showed that the conventional heat cured Paladron resin had a significantly higher FS (78.6 6 5.5 MPa) followed by the chemically modified heat cured Sinomer (72.3 6 2.1 MPa) and the microwave cured resin (67.2 6 5.3 MPa). No remarkable difference was reported between the FS of Paladon 65 and the thermoplastic resin Polyan. However, Promysan had a significantly higher FS than Paladon 65. Barbosa et al. (2007) investigated the FS of microwave (Onda-Cryl), heat (Classico), and chemically cured (Jet) acrylic resins, when they were polymerized using different curing cycles. The results indicated that that the highest FS was demonstrated by Onda-Cryl (109.63 6 5.31 MPa), followed by Classico (92.84 6 4.73) and Jet (84.40 6 1.68 MPa). It was observed that the heat cured resin showed a higher FS in comparison to the chemically cured resin, however, the difference was not significant. The results of this study also pointed out that a

8.3 Properties of PMMA Denture Base Resins

longer curing duration for specimens in a water bath seems to enhance water uptake, which ultimately resulted in a reduction in the strength of the DBMs. In addition, autopolymerizing specimens that were cured at higher temperatures had better mechanical properties than those cured at room temperature. Ali et al. (2008) evaluated the mechanical properties of dual cure (Eclipse), heat cured (Meliodent), and autopolymerized resins (Probase Cold) having dimensions of 65 3 10 3 2.5 mm. The Meliodent specimens were fabricated by heating in a water bath, while the cold curing Probase specimens were chemically cured in a pressure pot. The authors of this study pointed out that the FS of the dual cure resins was significantly higher that of than the heat and chemically cured resins. The Meliodent heat cured resin possessed greater FS than that of Probase. In light of the mentioned studies, it can be concluded that while the FS of denture base resins is dependent on various factors, such as the chemical composition of the DBM, the method of polymerization, the duration and temperature of polymerization as well as the preparation, storage, and thickness of the specimens; heat cured PMMA DBMs have a significantly higher FS in comparison to chemically cured materials.

8.3.2 FRACTURE TOUGHNESS Fracture toughness (FT), also known as the critical stress intensity factor, is basically a measure of the ability of a denture base to prevent the propagation of cracks, when notches and inherent material flaws are present (Zappini et al., 2003). ISO 20795-1 (2013) for DBM describes the method for assessing the FT of DBMs, by performing a three-point bending of specimens that contain a notch in the midline. Stafford et al. (1980) compared the FT of heat, autopolymerized, and high impact strength denture base resins, using single edge notch (SEN) and tapered cleavage (TC) methods. The results indicated that regardless of the testing method, the heat cured resins possessed a significantly higher FT (2.06 6 0.17 MN m23/2) as compared to the autopolymerized (1.63 6 0.1 MN m23/2) and high impact strength resins (1.77 MN m23/2). Interestingly, it was observed that the wet specimens had higher FTs in comparison to the dry PMMA specimens. Similarly, Neihart et al. (1988) investigated the FT of several denture base resins using the short rod method, and pointed out that Lucitone heat cured resin possessed a significantly higher FT (1.17 6 0.06 MN m23/2) in comparison to caulk autopolymerizing repair resin (0.99 6 0.03 MN m23/2). The authors of this study did not provide any information regarding the dimensions of the specimens as well as the crack that was produced in the specimens for testing the FT, which are essential in the interpretation and comparison of materials tested in this study. Finoti et al. (2012) compared the FT of a heat cured Lucitone 55 resin and four reline resins. The heat cured resin possessed a remarkable FT in comparison to the reline resins. The FT of the DBMs actually increased following long term immersion, which is also supported by previous studies (Stafford et al., 1980). In

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conclusion, the FT of heat cured DBM is generally greater than cold cured acrylic resins. Therefore, these materials are better able to resist fractures caused by the development and propagation of cracks within their structure.

8.3.3 IMPACT STRENGTH Another property related to FT is impact strength, which is defined as the energy required to fracture a denture base resin under the effect of an impacting force (Zappini et al., 2003; Anusavice et al., 2012). Despite the frequent use of this property due to its ease of testing for DBM, impact strength is not an inherent property of a material as it is dependent upon the dimensions of the specimen, notch depth, impact velocity, and several other factors (Oku, 1988). Robinson and McCabe (1993) measured the impact strength of 12 different types of PMMA DBMs, which contained varying degrees of surface defects. The results of their study indicated that the impact strengths of acrylic DBMs were significantly reduced in the presence of surface defects as small as 16 µm. These surface defects can serve as notched areas where cracks tend to propagate and lead to fracture. The addition of butadiene styrene rubber to PMMA results in significant enhancement in its impact strength (Tandon et al., 2010) and in the formation of “high impact strength” PMMA DBMs. However, the increase in impact strength is also accompanied with a considerable decrease in elastic modulus.

8.3.4 CROSSLINKING Crosslinking agents are commonly added to PMMA DBMs for improving their properties. Crosslinking agents have three beneficial effects on acrylic resins. Firstly, they decrease the susceptibility of these resins to be dissolved in organic solvents (Jagger and Huggett, 1990; Arima et al., 1996). Secondly, they enhance the resistance of PMMA resins against crazing under stress (O’Brien, 2002). Finally, crosslinking agents also help in improving the handling characteristics and mechanical properties of acrylic dentures. Crosslinking agents increase the molecular weight of polymers and prevent the formation of an oxygen inhibition layer, thereby ensuring minimal residual monomer concentration in the set polymer. The most commonly used crosslinking agents in dentistry are ethylene glycol dimethacrylate (EGDMA) and 1,4-Butylene glycol dimethacrylate (BGDMA) (Arima et al., 1996).

8.3.5 SORPTION AND SOLUBILITY Sorption refers to the process by which materials take up water or any fluid in which they are immersed (Latief, 2012). PMMA resins have been shown to absorb water, which is dependent on the polarity of the resin molecules (Miettinen and Vallittu, 1997b). Following immersion, water molecules tend to penetrate the spaces between the polymer chains, and as a result, widen the

8.3 Properties of PMMA Denture Base Resins

existing gap between the polymer chains. According to the guidelines of ISO 20795-1 for DBM (2013), the sorption values for heat cured denture base resins must not exceed 32 µg mm23, while the solubility should not be greater than 1.6 µg mm23 (Table 8.4). Water uptake within PMMA resins has two main effects. Firstly, water molecules act as plasticizers and interfere with the entanglement of the polymer chains. Secondly, the uptake of water tends to cause an expansion of the polymer, thereby affecting its dimensional stability (Takahashi et al., 2013). Solubility refers to the maximum amount of a solute that can be dissolved in a solvent at a given temperature for a given period of time (Bayraktar et al., 2006; Latief, 2012). The ISO standard 20795-1 for denture base resins (2013) requires that type 1, 3, and 5 denture base resins should not have a solubility greater than 1.6 µg mm23, while the solubility of type 2 denture base resins must be lesser than 8 µg mm23 (Table 8.4). Similarly, Cucci et al. (1998) investigated sorption and solubility in two autopolymerizing (Duraliner II and Kooliner) and one heat cured resin (Lucitone 550). It was observed that Duraliner II showed significantly lesser WS (19.48 6 6.94 µg mm23) in comparison to Kooliner (29.99 6 8.15 µg mm23) and Luctione 550 (38.31 6 5.43 µg mm23), while there was no significant difference between the WS of Lucitone and Kooliner DBMs. The solubility of Duraliner was 19.48 6 6.94 µg mm23, while that of Kooliner was 29.99 6 8.15 µg mm23. Tuna et al. (2008) evaluated the WS and solubility of two heat cured and eight autopolymerized denture base resins. The authors pointed out that while all the denture base resins had WS values that were within the limits of ISO 20795-1 (2013), the heat cured resins generally had lesser WS values and solubility than the autopolymerized resins. Apart from the residual monomer that leaches out of denture base resins, Cimpan et al. (2000a) pointed that other water soluble toxic materials, such as formaldehyde, benzoic acid, dibutyl phthalate, phenyl benzoate, phenyl salicylate, and diclohexyl phthalate also tended to leach out of DBMs (discussed in detail in Section 8.3.6). In light of the studies discussed, it can be concluded that the sorption of a DBM is directly related to the amount of residual monomer within the cured resin Table 8.4 Some Properties of E-Glass Fibers Used for the Reinforcement of Dental Prostheses (Vallittu, 1998; Khan et al., 2015) Property

Value 23

Density (g cm ) Tensile strength (GPa) Modulus of elasticity (GPa) Elongation at break (%)

2.54 3.4 73 2.8

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which can be influenced by various factors, such as the type of DBM, polymerization conditions and methods, duration of specimen storage as well as the thickness of the specimens. Since most of these parameters are better controlled in heat cured DBMs, these materials tend to exhibit better physical, mechanical, and biological properties in comparison to chemically cured DBMs.

8.3.6 THERMAL CONDUCTIVITY The thermal conductivity of PMMA resins is quite low at approximately 5.7 3 1024 C cm21 (Alla et al., 2015). Due to the low thermal conductivity of acrylic resins, heat generated during fabrication cannot escape, leading to a considerable rise in temperature and potentially causing surface crazing. Another problem associated with the low thermal conductivity of PMMA is that denture wearers are unable to detect the temperature of the food and drinks. As a result, patients may take an extremely hot drink without having any sensation until the hot fluid burns their soft tissues or scalds the esophagus or throat (Van Noort and Barbour, 2013).

8.3.7 RESIDUAL MONOMER The solubility of a denture resin is actually representative of the amount of the unreacted monomer, plasticizer, or initiator as well as any water-soluble components that leach out when denture base specimens are immersed in water for a week. Therefore, denture bases that contain a high level of residual monomer tend to have a higher degree of solubility (Fletcher et al., 1983; Latief, 2012). Douglas and Bates (1978) evaluated the residual monomer content for heat, chemical, and pour type resins. The results indicated that the heat cured resins possessed a significantly lesser amount of residual monomer when compared with the chemically cured and pour type resins Dogˇan et al. (1995) studied the effects of residual monomer in cured heat and autopolymerized resins on the mechanical properties as well as their sorption and solubility. They observed that there was a direct relationship between the residual monomer and WS, and pointed out that the greater the amount of residual monomer in a resin, the higher the sorption will be. In addition, it was also found that a longer duration of polymerization for heat cured resins and a higher polymerization temperature for autopolymerized resins seemed to reduce the amount of residual monomer, and hence, improved the mechanical properties of these resins. Similar results were achieved by Vallittu et al. (1995a) who concluded that chemically cured resins had a greater amount of residual monomer in comparison to heat cured resins. Bayraktar et al. (2006) found that heat cured denture base resins had significantly lesser degrees of residual MMA monomer in comparison to microwave cured resins.

8.3 Properties of PMMA Denture Base Resins

8.3.8 COLOR STABILITY DBMs are expected to mimic the oral mucosa in shade and color. Therefore, denture base resins not only need to match the color of oral tissues, but they are also required to have excellent color stability, so that their esthetics do not degrade over time (Sagsoz et al., 2014). PMMA denture resins are prone to undergo color changes over time, which can be linked to various factors. A high residual monomer content as well a poor degree of conversion tends to decrease the color stability of PMMA. Similarly, porosity within the polymerized structure caused by overheating or improper fabrication technique can also result in the staining of PMMA (Scotti et al., 1997). Finally, frequent intake of beverages such as tea, coffee, and wine also causes the staining of PMMA resins (Scotti et al., 1997).

8.3.9 RADIOPACITY DBMs are desired to be radiopaque, so that they are visible on a radiograph if a fractured piece of denture is accidentally swallowed (Chandler et al., 1971a; Bloodworth and Render, 1992; Young, 2010). Unfortunately, PMMA DBMs are inherently radiolucent and are quite difficult to visualize on radiographs in cases where they are accidentally inhaled or ingested (Lang et al., 2000). Heavy metallic elements have been incorporated to enhance the radiopacity (Lang et al., 2000) or by alternatively using metallic DBMs (Young, 2010). However, all these materials had their own merits and demerits and none of these efforts provided an adequately radiopaque DBM. Initially, efforts were made to incorporate pins, lead foils, and barium sulfate impregnated yarns into the denture base resins. However, these materials did not provide sufficient radiopacity to the DBMs, as they were unable to chemically or physically bind to the denture base and tended to leach out with the passage of time (Bloodworth and Render, 1992). The incorporation of metal salts provided better results in terms of radiopacity. However, these needed to be added in extremely high concentrations in order to provide the desired results of radiopacity. It was demonstrated that the addition of a concentration greater than 10% of these salts led to a remarkable reduction in the bond strength as well as esthetics of the dentures (Bloodworth and Render, 1992). The addition of glass fillers as a means of enhancing the radiopacity of DBMs has also been investigated (Chandler et al., 1971b; Young, 2010). Probably the most promising results were achieved by Chandler et al. (1971b) by the addition of silanated glass particles into denture base resins. This method not only improved the radiopacity of the dentures, but also provided an optical clarity and enhanced the esthetics of cured denture bases. Furthermore, the problem of the leaching of radiopaque fillers from denture bases which was observed in the case of metallic salts was also eliminated. However, this approach resulted in the reduced mechanical properties of denture base resins. In addition, the polishability and laboratory fabrication of these resins was difficult in comparison to the conventional DBM (Chandler et al., 1971b; Bloodworth and Render, 1992).

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Lang et al. (2000) investigated the effect of adding various concentrations of Triphenylbismuth (TPH) on the radiopacity of heat cured, injection molded acrylic denture base resins. A concentration of 30% w/w of TPH made the acrylic resins sufficiently radiopaque without significantly disturbing its mechanical properties as well as the esthetics of the dentures. In conclusion, an increased radiopacity is an essential requirement for DBMs so that they can be easily detected on a radiograph and can be easily removed in case a broken piece of a denture is swallowed or inhaled into the lungs, which would otherwise lead to a medical emergency.

8.3.10 BIOCOMPATIBILITY AND CYTOTOXICITY Biocompatibility refers to the ability of a material to provide a favorable host response when it is used within the human body (Gautam et al., 2012). A basic requirement for a DBM is that it should be nontoxic and safe for use within the oral cavity. Various concerns regarding the biocompatibility of acrylic resins have been raised primarily due to the residual monomer that is inevitably present within cured denture base resins and which can cause tissue irritation, inflammation, as well as cytotoxicity (Jorge et al., 2003; Lung and Darvell, 2005). Kedjarune et al. (1999) pointed out that the higher the amount of monomer mixed with the polymer, the greater was the cytotoxicity as well as the amount of residual monomer within the cured resin. Similarly, Lamb et al. (1983) observed that polymers that were prepared using a higher polymer/monomer ratio contained a lesser degree of residual monomer in comparison to those that were prepared using a lower polymer/monomer ratio. It was also concluded that resins that were cured using an extended polymerization cycle contained lesser amounts of residual monomer in comparison to those that were cured using a shorter cycle. Tsuchiya et al. (1994) and Cimpan et al. (2000b) observed that there was a significantly greater elution of substances from autopolymerized resins and they possessed greater cytotoxicity when compared with heat and microwave polymerized denture base resins. Similar results were obtained by Sheridan coworkers (1996), who observed that autopolymerizing denture base resins exhibited the greatest cytotoxic effects in comparison to heat and microwave cured DBM. It was also pointed out that the residual monomer and the cytotoxicity of resin materials were reduced considerably following their immersion in water before use. From this discussion, it can be concluded that the cytotoxicity of a DBM depends on the chemical natural of the DBM, the method of polymerization, the polymer/monomer ratio used, and the degree of residual monomer. In general, heat cured DBMs contain relatively lesser amounts of residual monomer and other leachable components after polymerization; they are less cytotoxic in comparison to cold cured polymers. In addition, the cytotoxicity of a DBM can be significantly reduced by increasing the duration of polymerization and by immersing the prepared denture in water prior to use.

8.4 Contemporary Denture Base Materials and Modifications of PMMA

8.4 CONTEMPORARY DENTURE BASE MATERIALS AND MODIFICATIONS OF PMMA Despite the several favorable properties of PMMA as a denture base resin, such as its biocompatibility, esthetics, and ease of processing, polishing as well as repairing, this material is far from ideal. PMMA undergoes volumetric as well as linear polymerization shrinkage, which can lead to significant changes in the dimensional accuracy leading to the fabrication of faulty dentures (Anusavice et al., 2012). Another drawback of PMMA DBMs is their tendency to fracture (Darbar et al., 1994) as a result of their inherently poor flexural and impact strengths, implying that dentures prepared from PMMA tend to fracture easily when they are flexed or bent repeatedly during mastication (Beyli and Von Fraunhofer, 1981) or when they are accidentally dropped. In addition, PMMA resins tend to absorb water, resulting in a remarkable reduction in its mechanical properties due to hydrolysis. Flexural fatigue of PMMA dentures occurs due to the repeated application of a bending force. While the application of such force only on one occasion may not fracture the denture, it can lead to the development of microscopic cracks, which gradually fuse together upon repeated bending and ultimately lead to fracture (Smith, 1961; Johnston et al., 1981). The fracture of acrylic dentures can also occur as a result of an impacting force having such a high magnitude that the dentures can fracture spontaneously. This mode of fracture usually occurs as a result of accidental dropping while cleaning, sneezing, or coughing, or due to a sudden blow to the face (Jagger et al., 1999). While the fracture resistance of acrylic dentures can be improved by increasing the thickness of the denture base, such a practice can lead to excessively thick dentures that can cause gagging, but it can also interfere with the movement of the coronoid process of the mandibles and cause dislodgement of the dentures in addition to preventing the normal functioning and efficiency of the dentures (Zarb et al., 1997; Meng and Latta, 2005; Kiilu, 2008). In an attempt to avoid excessively thickening denture bases, various techniques have been tried by several authors, such as the use of alternate denture bases, the chemical modification of PMMA or the reinforcement of PMMA with fibers or particulate fillers to enhance its strength (Tandon et al., 2010). A brief account of these techniques is given here:

8.4.1 POLYAMIDES Polyamides, also known as nylons, are a group of thermoplastic materials that are produced by a condensation reaction between a dibasic acid and a diamine (Sepu´lveda-Navarro et al., 2011). In contrast to PMMA, which is amorphous, nylon is a crystalline material that is insoluble in various solvents and possesses high heat resistance as well as flexibility (Stafford et al., 1986; Sharma and Shashidhara, 2014). Stafford et al. (1986) compared various physical and

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mechanical properties of nylon and conventional resins. The nylon dentures had significantly higher FSs in comparison to the PMMA DBM due to their enhanced flexibility. In addition, the nylon dentures also showed significantly higher impact strength than the conventional acrylic resins. Ucar et al. (2012) observed that the FS of polyamide denture base polymers was significantly greater than that of acrylic resins. In contrast, the flexural modulus of the nylon DBM was significantly lower in comparison to PMMA. Yunus et al. (2005) compared the mechanical properties of polyamides and conventional PMMA denture base resins, and pointed out that the nylon denture possessed a significantly lower FS in comparison to compression molded or microwave cured PMMA resins. Due to their excellent flexibility, dentures prepared from polyamides have an excellent resistance against flexural fatigue and fracture (Goiato et al., 2010) and prove to be highly useful in cases where patients are unable to tolerate hard denture bases. In addition, they can also be used as alternative materials in individuals who are allergic to PMMA. However, due to the difficulty in processing and repairing of these materials, coupled with their high WS, solubility, and their potential for staining (Stafford et al., 1986; Sagsoz et al., 2014; Wieckiewicz et al., 2014), polyamides are not frequently used for the fabrication of dentures, except for a few special cases (Shamnur et al., 2010).

8.4.2 EPOXY RESINS These polymers are used in liquid form and poured directly into a mold. These materials possess sufficient strength, hardness, and toughness, in addition to having excellent dimensional stability and the ability to bind to metallic components of dentures. However, despite their benefits, epoxy resins are not commonly used due concerns regarding their toxicity, high WS (Stafford and Braden, 1968; Alla et al., 2015), and inability to chemically bind with the acrylic teeth (Smith, 1962).

8.4.3 POLYCARBONATES Polycarbonates are basically chains of bisphenol-A carbonate. Although these materials possess better mechanical properties than the conventional acrylic denture base resins, polycarbonates are not frequently employed for the fabrication of dentures as a result of their difficulty of processing and the high cost involved in their fabrication (McCabe and Wilson, 1974). In addition, concerns have been raised about their toxicity due to their tendency to leach out bisphenol-A (BPA) during their fabrication and in clinical use (McCabe and Wilson, 1974; Suzuki et al., 2000).

8.6 Reinforcement of PMMA Denture Base Materials

8.5 CHEMICAL MODIFICATION OF PMMA In order to overcome the inherent drawbacks of PMMA, several attempts have been made toward their chemical modification (Jagger et al., 1999). Rubber has been commonly incorporated into PMMA so as to fabricate high impact strength acrylic dental resins (Garg and Mai, 1988; Jagger et al., 1999). The addition of rubber into PMMA results in the formation of a semi-interpenetrating network of PMMA and rubber. If a crack develops within the denture, it tends to propagate through the PMMA matrix. However, it will decelerate when it comes into contact with the rubber. This implies that high impact strength resins are capable of absorbing higher impact forces at higher strain rates before fracturing. Rodford (1990) incorporated low molecular weight butadiene styrene rubber into PMMA resins, and observed a considerably greater impact strength of the chemically modified resin. Stafford and Braden (1968) reviewed various denture base resins and pointed out that rubber modified resins had better impact strength and dimensional stability in comparison to other denture resins. Rubber modified high impact strength resins prevent denture fracture occurring after accidental dropping or due to a sudden impact. However, in the case of a fracture caused due to anatomical reasons, their fracture resistance has been shown to be identical to that of conventional PMMA resins (Uzun and Hersek, 2002). Regardless of the success of this technique, chemical modification of PMMA DBM is restricted clinically due to the high costs incurred in the fabrication of these modified resins (Jagger et al., 1999).

8.6 REINFORCEMENT OF PMMA DENTURE BASE MATERIALS The reinforcement of PMMA using various fiber or metallic inserts has also been investigated by various authors as a means of strengthening the resin and to prevent acrylic denture fractures; it has also shown promising results in terms of the enhancement of the mechanical properties of dentures (Ladizesky et al., 1993b; Vallittu, 1995, 1996; Jagger et al., 1999; Butterworth et al., 2002).

8.6.1 REINFORCEMENT WITH METAL WIRES OR MESH Metals have been incorporated into acrylic denture resins in the form of particulate fillers, mesh, or wires as a means of enhancing their strength (Jagger et al., 1999). Vallittu et al. (1995a) incorporated metallic wires and mesh into PMMA and observed the effect of various dimensions and forms on its mechanical properties. The results indicated that the addition of metallic wires enhanced the transverse strength of acrylic resins by up to 85% in comparison to unreinforced

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PMMA, while the addition of a metallic mesh had little reinforcing effect on the acrylic resins. Vojdani and Khaledi (2006) assessed the addition of metallic wires on the flexural strength of denture bases and reported that metallic wire reinforced PMMA specimens had a significantly higher FS than unreinforced resins. Similar results were achieved by Vallittu et al. (1995b), who evaluated the impact strength of conventional as well as metal wire and continuous E-GF reinforced acrylic resins. The results showed that the metal wire significantly improved the impact strength of PMMA in comparison to the unreinforced resins. However, no remarkable differences were observed while comparing metal wire and GF reinforced resins. Despite the enhancements in the mechanical properties of DBM after the incorporation of different types of metallic inserts, various difficulties were encountered with this technique, such as poor adhesion between the wire and the resin and the development of areas of stress concentration within the denture base. Attempts were also made to enhance the adhesion of wires with the PMMA matrix through sandblasting (Vallittu and Lassila, 1992) or chemical bonding (Jacobson et al., 1988; Vallittu, 1993), which proved to be unsuccessful. As a result, the incorporation of metallic wires within acrylic resins has limited value in dentistry due to their difficulty of inclusion as well as their poor adhesion with the PMMA resin matrix.

8.6.2 FIBER REINFORCEMENT The incorporation of various fibers within PMMA has been shown to improve its mechanical properties (Solnit, 1991; Vallittu, 1996; Alla et al., 2013). The addition of fibers into acrylic resins results in the formation of a PMMA
fiber composite which possesses properties that are independent of the individual resin and fibers used (Vallittu, 1996). A PMMA
fiber composite consists of a resin matrix into which fibers have been incorporated. A characteristic feature of these composites is their relatively greater length in comparison to their cross-sectional dimensions. Fiber reinforcement of denture base polymers has proved to be one the best techniques for reinforcing their mechanical properties. Fiber reinforced denture polymers (FRPs) are constituted of a matrix consisting of a resin, into which fibers have been incorporated. The matrix provides protection to the fibers against the harsh oral environment, while the fibers enhance the strength of the resin matrix. Factors such as fiber type, length, orientation, silane treatment, concentration, and resin preimpregnation, determine the mechanical attributes of fiber reinforced dental prostheses (Khan et al., 2015).

8.6.2.1 Effects of fiber length on properties of fiber reinforced denture base resins Depending on length, fibers can be categorized as long/continuous or short fibers. Short fibers are those that do not extend through the span of the prosthesis

8.6 Reinforcement of PMMA Denture Base Materials

(O’Regan and Meenan, 1999), while long or continuous fibers extend throughout the length of the prosthesis (Xu et al., 2003). In the case of short fibers, there exists a critical length (Lc) (Petersen et al., 2006; Garoushi et al., 2007; Chawla, 2012) and fibers shorter than this length tend to decrease instead of enhancing the mechanical properties of denture bases. A Lc of 0.5
1.6 mm has been measured for short GFs. Fibers greater than this length have been shown to mechanically enhance the properties of dental resins (Petersen, 2005).

8.6.2.2 Effect of fiber orientation Fiber orientation within a resin plays a pivotal role in determining its mechanical properties (Dyer et al., 2004; Garoushi et al., 2006; Karbhari and Strassler, 2007; van Heumen et al., 2008; Hyer, 2009). Fibers oriented in one direction are called unidirectional (Vallittu, 1996), and are anisotropic, implying that they impart strength and rigidity only in a single direction (Vallittu, 1996; Butterworth et al., 2002). These unidirectional fibers possess a Krenchel’s factor (Laws, 1971) of 1% or 100% (Gharoushi and Vallitu, 2009). Multidirectional or woven fibers are dispersed throughout a resin in a network or mesh form (Fig 8.4). These fibers are isotropic, and provide strength to the matrix in all directions. However the magnitude of mechanical enhancements provided by these fibers is lesser than that of unidirectional fibers and their Krenchel’s factor is less than 1 (Garoushi et al., 2008) (Fig. 8.5). These fibers are useful if the direction of the force is not known or in cases where there is a lack of space for the placement of unidirectional fibers (Garoushi et al., 2009).

8.6.2.3 Effects of resin impregnation on PMMA resin-based materials An important aspect of the strength of fiber reinforced acrylic resins is the extent of adhesion between the fibers and the matrix. One method of enhancing the adhesion between the fibers and the PMMA is by preimpregnation of fibers, which is achieved by coating them with the resin monomer prior to 0

1

0.5

0.25

0.38

FIGURE. 8.4 Reinforcing efficiency (Krenchel’s Factor) of glass fibers with variable plane orientations. Adapted from Khan, A.S., Azam, M.T., Khan, M., Mian, S.A., Rehman, I.U., 2015. An update on glass fiber dental restorative composites: a systematic review. Mater. Sci. Engi. C, 47, 26
39.

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Glass fibers

A-glass

E-glass

C-glass

AR-glass

R-glass

S-glass

FIGURE 8.5 Classification of different kinds of glass fibers used in dentistry for reinforcement of dental prostheses.

polymerization (Freilich et al., 1998). Various studies have documented an increased wettability and bond strength between the resin and fibers when they are preimpregnated with resin monomer (Vallittu, 1996; Behr et al., 2000; Bae et al., 2001; Chong and Chai, 2003). Preimpregnation of fibers ensures a uniform load distribution between the matrix and the fibers, which leads to enhanced mechanical properties. Moreover, minimal voids within the resin matrix are present, thereby minimizing the probability of crack generation and propagation within the PMMA
fiber composite.

8.6.2.4 The effect of silane treatment on properties of PMMA denture base resins Silane coupling agents promote adhesion between a resin and fibers (Kanie et al., 2000). Fibers that have not undergone any surface treatment with a coupling agent such as silane act as impurities and create voids within the matrix, which decreases the strength of the polymer (Grant and Greener, 1967; Solnit, 1991). Silane treatment of fibers helps in the formation of a chemical bond between acrylic resins and fibers (Solnit, 1991). In addition, these fibers exhibit superior bond strengths in comparison to fibers that have not undergone silane treatment (McDonough et al., 2001; Matinlinna et al., 2004). The most commonly used silane in dentistry is 3trimethoxysilylpropyl methacrylate (MPS) (Lung and Matinlinna, 2012).

8.6.3 DIFFERENT TYPES OF FIBERS USED IN DENTISTRY 8.6.3.1 Carbon fibers The earliest use of carbon fibers (CFs) for PMMA reinforcement, which was in 1970s, reported a twofold increase in transverse strength of PMMA (Schreiber, 1971). Another study demonstrated similar results with an 89% increase in the elastic modulus (Larson et al., 1991) of CF reinforced PMMA. Beneficial properties of CFs include their high flexibility and tensile strength (Manley et al., 1979), coupled with the low molecular weight and low thermal expansion of CF reinforced PMMA composites. However, their clinical use is limited due their poor aesthetics (Mona, 1999; Uzun et al., 1999) and difficulty of incorporation into PMMA.

8.6 Reinforcement of PMMA Denture Base Materials

8.6.3.2 Aramid fibers Aramid or Kevlar fibers, which are chemically constituted of polyamide fibers, result in better impact strength of resin based materials (Mona, 1999). These fibers possess mechanical properties and wettability that is superior to CFs and GFs (Alla et al., 2013). Also, they do not require any pretreatment with coupling agents, as in the case of CFs or GFs. Their use in the fabrication of dental prostheses is, however, limited due to their high cost, unpleasant yellow color (Mona, 1999; Alla et al., 2013), and the tendency of the exposed fibers to cause tissue irritation (Mona, 1999).

8.6.3.3 Polyethylene (UHMWPE) fibers In contrast to CFs and AFs, using white colored fibers of ultra-high molecular weight polyethylene (UHMWPE) does not alter the aesthetics of acrylic dentures (Mona, 1999; Alla et al., 2013). They have been shown to impart ductility and toughness to denture base resins (Clarke et al., 1992; Ladizesky et al., 1992, 1993a,b). UHMWPE fibers can be used in a chopped, woven, or continuous form for the fabrication of dental prostheses. Various studies have shown the effect of orientation and fiber concentration on strength of UHMWPE fiber reinforced materials (Vallittu, 1996; Alla et al., 2013). A concentration of greater than 1% UHMWPE fibers has been shown to significantly increase impact resistance, while concentrations in excess of 3% tend to make the resin highly viscous and unmanageable. Despite the improvements in the mechanical properties of PMMA resins, the complicated technique for the placement of these fibers within PMMA along with the need for surface treatment of these fibers prior to use has limited their widespread clinical use in prosthodontics (Alla et al., 2013).

8.6.3.4 Glass fibers Problems posed in utilizing CF, AF, and UHMWPE fibers led to the search for alternative reinforcing fibers that possessed strength and did not affect the esthetics of dentures. This led to the discovery of glass fibers (GFs), which possess excellent aesthetics and mechanical properties (Uzun et al., 1999; Tacir et al., 2006), and are widely utilized in clinical dentistry currently (Gharoushi and Vallitu, 2009). Their use for fiber reinforcement was documented decades ago (Vallittu, 1996). Different forms and orientations of GFs have been experimented with to reinforce acrylic resins. Since GFs are the most commonly used material for the reinforcement of dentures, and have also been used in this study, they will be discussed here in detail.

Types and composition of glass fibers Several types of GFs are used in dentistry, such as C, D, S, and R-type GFs. A classification of GFs is given in Fig. 8.5. A-glass fibers contain 25% soda and lime. They are mainly used as fillers for plastics owing to their low cost of production in comparison to the other GFs. C-glass fibers are used in cases where an

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excellent chemical resistance is desired. However, these fibers are too weak to be used as insulating materials. D-glass fibers have excellent electrical properties for such applications. On the other hand, their low chemical resistance in comparison to other GFs limits their use. S-glass fibers are used where optimal strength and stiffness is required. However, due to their high cost, labor consuming production process, and short service life span, the use of S-glass fibers is limited. E-type glass is a calcium-aluminum-borosilicate glass that contains a very low alkali content. E-glass fibers have excellent mechanical properties (Table 8.5) and are able to effectively resist degradation in a hydrated environment. Mainly E-glass fibers are used in dentistry owing to their high strength, resistance against various chemicals, as well as lower cost in comparison to the other GFs (Zhang and Matinlinna, 2012; Khan et al., 2015). The different types and compositions of GFs used in dentistry are summarized in Table 8.5.

Properties of glass fiber reinforced denture base resins Various studies have documented the beneficial role of GFs in the improvement of the mechanical properties of resin based materials (Vallittu, 1996; Mona, 1999; Tacir et al., 2006; Garoushi et al., 2007; van Heumen et al., 2008; Alla et al., 2013). Uzun et al. (1999) reported an increased impact strength while reinforced with GFs. Similar results were also achieved by Goguta et al. (2006) and Hari and Mohammed (2011). The effect of the length and orientation of GFs have also been investigated. Several authors have reported an increased impact strength (Uzun et al., 1999; Hari and Mohammed, 2011) and transverse strength of PMMA DBMs when they were reinforced with woven GFs in comparison to unidirectional fibers (Unalan et al., 2010).

Table 8.5 Composition in wt.% of Different E-Glass Fibers (Khan et al., 2015) Component

A-Glass (%)

E-Glass (%)

C-Glass (%)

AR-Glass (%)

R-Glass (%)

S-Glass (%)

SiO2 Al2O3 CaO MgO B2O K2O Na2O Fe2O3 ZrO2 ZnO

71 3 8.5 2.5

15




53
55 14
16 20
24 20
24 6
9 ,1 ,1 ,1
0
0.7

56
58 12 17
22 2
5
0.4 0.1
2 0.2
2 2 2

62 0.8 5.6


14.8
16 0

75.5 0.5 0.5 0.5 20 3 0
1




62
65 20
25

0
1

0.2



8.6 Reinforcement of PMMA Denture Base Materials

A similar conclusion was also made by Unalan et al. (2010) in a study that reported the increased transverse strength of denture bases when using woven GFs. Kim and Watts (2004) also showed increased impact strength of acrylic resin when used with woven E-glass fibers. In contrast, Vallittu et al. (1994) and Vojdani and Khaledi (2006) showed a greater transverse strength while using unidirectional fibers. Vallittu (1999) observed an enhancement in the flexural properties of PMMA containing GFs. It has been shown that the highest mechanical strength is achieved when unidirectional GFs are used, in comparison to woven GFs (Vallittu, 1999; Ellakwa et al., 2002; Vojdani and Khaledi, 2006; Ko¨ro˘glu et al., 2009). Miettinen and Vallittu (1997a) compared the physical properties of unreinforced and GF reinforced denture base resins. The results indicated that there was an insignificant difference between the WS and solubility of both types of resins. Polat et al. (2003) compared the WS and solubility of injection and compression molded and GF reinforced PMMA resins. The results indicated that the WS and solubility of injection molded fiber reinforced resins was lesser in comparison to the compression molded specimens. Durkan et al. (2010) investigated the amount of water absorption in heat and microwave cured fiber reinforced DBMs and demonstrated that the microwave cured GF reinforced DBMs had significantly lesser water absorption. In addition, it was also shown that StickNET woven GFs had lesser WS than unidirectional Stick GFs. In contrast, C ¸ al et al. (1999) showed that the WS and dimensional changes in PMMA resins reinforced with unidirectional or woven GFs were lesser in comparison to unreinforced resins. It was also observed that the WS and dimensional changes in GF reinforced resins decreased with an increase in fiber content. The placement of unidirectional GFs within PMMA is quite difficult and they tend to clump within the acrylic dough. In addition, unidirectional GFs tend to protrude from the prosthesis due to wear, thereby causing tissue irritation. A more convenient method may be to use woven GFs that are relatively easy to place with acrylic resins and provide strength in all dimensions, combined with lesser chances of mucosal irritation in case they protrude out of the denture base. Various commercial GFs are available for the fabrication of reinforced dentures, such as StickNET, FiberKor, Vectris pontic, etc. However, these are quite expensive and significantly increase the cost of fabrication of dentures in comparison to conventional unreinforced dentures. The cost of the production of dentures reinforced with these commercial fibers is, therefore, high and can be reduced using industrial GFs, which are not only cheaper but also easily available and have been shown to provide similar results in terms of enhancement of the physical and mechanical properties of PMMA dentures (Vojvodi´c et al., 2009). Different types of PMMA DBMs (heat, cold, light, and microwave cured) have been used for the reinforcement of acrylic dentures and have been shown to enhance the physical and mechanical properties of these materials. Amongst these, heat activated acrylic resins are frequently used to reinforce dentures due to their superior physical and mechanical properties, relatively low cytotoxicity,

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ease of processing, and lower cost of production, in comparison to chemically cured materials. Although the newer types of PMMA resins, such as light and microwave cured polymers, possess superior mechanical and physical properties compared to conventional heat cured PMMA, they are quite expensive and require specialized equipment for their use in the fabrication of dental prostheses. As a result, they are not commonly used for making acrylic dentures and heat cured PMMA continues to be the most commonly used DBM for the preparation of prosthetic dental appliances.

8.7 CONCLUSION Poly(methyl methacrylate) finds widespread applicability worldwide for the fabrication of artificial teeth and dental prostheses. Despite having several beneficial properties, PMMA is prone to fracture due its weak mechanical properties and susceptibility to hydrolytic degradation. Attempts have been made at imparting favorable characteristics to PMMA DBMs showing varying results. Among these efforts, the addition of glass fibers within PMMA has provided excellent results.

LIST OF ABBREVIATIONS µg a0 AF ANOVA ATR b BPA BPO bt CQC C
C CF CO2 cm21 cps DBM DW EDS Exp-I Exp-II Exp PMMA F FPD

Microgram Precrack length Aramid Fibers Analysis of variance Attenuated total reflectance Width of specimens for measuring flexural strength Bisphenol-A Benzoyl peroxide Width of specimens for measuring fracture toughness Carbon
carbon double bond Carbon
carbon single bond Carbon fibers Carbon dioxide Per centimeter (wave number) Counts per second Denture base materials Distilled water Energy dispersive X-ray spectroscopy StickNET glass fiber reinforced PMMA specimens Industrial glass fiber reinforced PMMA specimens Exp-I and Exp-II PMMA specimens Force required to fracture specimens Fixed partial denture

List of Abbreviations

FRP FS FT FTIR GF GPa H H ht kN kV L IR Lc m1 m2 m3 M Max Min mL mm3 MMA MPa MPS Mt N nA PEMA PMMA PVC RPD RPM SEM SEN SPSS SS t^1/2 TC TPH UHMWPE UTM v/v % Vmax W0 Wt Wd WS Wsl

Fiber reinforced denture polymers Flexural strength Fracture toughness Fourier transform infrared Glass fiber Gigapascal Hydrogen atoms Thickness of specimens for measuring flexural strength Height of specimens for measuring fracture toughness Kilo Newtons Kilovolts Distance between the supports of UTM Infrared Critical length of reinforcing fibers Conditioned mass Mass of specimens after immersion Reconditioned mass Molar concentration Maximum Minute Milliliter Cubic millimeter Methyl methacrylate Megapascal (Trimethoxysilyl)propyl methacrylate Mass uptake (%) of PMMA specimens at time t (sec^1/2) Newton Nano Amperes Poly(ethyl methacrylate) Poly(methyl methacrylate) Polyvinyl chloride Removable partial denture Revolutions per minute Scanning electron microscopy Single edge notch Statistical package for social sciences Stainless steel Square root of time (seconds) Tapered cleavage Triphenyl bismuth Ultra-high molecular weight polyethylene fibers Universal testing machine Volume by volume percentage Wave number Weight of PMMA specimens at time 0 (s^1/2) Weight of PMMA specimens at time t (s^1/2) Weight of PMMA specimens at equilibrium in desorption Water sorption Water solubility

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FURTHER READING Bhola, R., Bhola, S.M., Liang, H., Mishra, B., 2010. Biocompatible denture polymers-a review. Trends Biomater. Artif. Organ 23, 129
136. Zafar, M.S., Ahmed, N., 2014. Nanoindentation and surface roughness profilometry of poly methyl methacrylate denture base materials. Technol. Health Care 22, 573
581.

CHAPTER

Maleic anhydride copolymers as a base for neoglycoconjugate synthesis for lectin binding

9

Nadezhda A. Samoilova1, Maria A. Krayukhina1, Olga S. Novikova2, Leonid M. Likhosherstov2 and Vladimir E. Piskarev1 1

A. N. Nesmeyanov Institute of Organoelement Compounds, Russian Academy of Sciences, Moscow, Russian Federation 2N. D. Zelinsky Institute of Organic Chemistry, Russian Academy of Sciences, Moscow, Russian Federation

9.1 INTRODUCTION Lectins (agglutinins) are unique biomacromolecules—carbohydrate-binding proteins of nonimmune origin (Sharon, 2007; Wu et al., 2008; Liener et al., 1986; Spain and Cameron, 2011). Although known for nearly a century, it was only during the past two decade lectins have become the focus of intense scientific interest. Lectinology can be divided into three major parts: (1) studies on lectins; (2) application of lectins as tools; and (3) lectins sensing and isolation. Studies on lectins are carried out to characterize the structure of lectin molecules, conformational and functional properties, their carbohydrate-binding specificity, and biological roles. The second part entails the application of lectins with well-defined specificity for different preparative or analytical purposes (as an excellent tool to study of glycoconjugation) both on cell surfaces or in solution. Three-dimensional structures and amino acid sequences of several hundreds of lectins have been established. Almost all, both without and with different ligands, have been elucidated mostly by high-resolution X-ray crystallography, generally up to 1.8 2.2 E (see 3D Lectin database: www.cermav.cnrs.fr/lectines). These studies provide many details about the interactions between lectins and sugars, and reveal novel protein folds and diverse quaternary structures (Sharon, 2007). Lectins are classified into six groups according to their specificities to monosaccharides: (1) Gal-specific lectins; (2) GalNAc-specific lectins; (3) Man and/or Glc-specific lectins recognizing complex N-linked oligosaccharides; (4) GlcNAc, and/or (5) Galβ1-4GlcNAcβ1-linked specific lectins, LFuc-specific lectins (subgroups based on the numbers and location of LFucα linkage); and (6) Sialic acid (SA) specific lectins (subgroups based on the recognized linkage of SA) (Wu et al., 2008).

Materials for Biomedical Engineering: Thermoset and Thermoplastic Polymers. DOI: https://doi.org/10.1016/B978-0-12-816874-5.00009-8 © 2019 Elsevier Inc. All rights reserved.

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Lectins and antibodies are the two main entities used for glycan profiling (lectin microarrays and lectin-based biosensors) (Dan et al., 2016). Lectins participate in a great number of vital biological processes like cell adhesion, pathogen infection, immunity, fertilization, and many others. Lectins that initiate cell adhesion are called selectins. Interactions of selectins with cell-surface glycoconjugates mediate tethering and rolling adhesion of leukocytes and platelets on vascular surfaces. Recent studies have helped elucidate the molecular details of selectin ligand interactions, the biosynthetic pathways for constructing selectin ligands, and the biophysical and cell biological features that modulate selectindependent rolling under flow (McEver, 2002). The effect of concanavalin A (Con A), fucose-binding protein (FBP), Ricinus communis agglutinin (RCA), and wheat germ agglutinin (WGA) on fertilization of the ascidian Phallusia mammillata was investigated (Honegger, 1982). Con A, FBA, and RCA had no influence on fertilization and did not bind to the chorian or sperm. So, for example, a significant feature of the lectins characterized in urodele egg jelly is their specificity. They are all inhibited by D-glucose derivatives (Hedrick, 1986). Lectin-based biosensors were used for the determination of carbohydrates, pathogenic bacteria and toxins, and cancer cells. Lectins have been widely used for the construction of electrochemical and optical biosensors by exploiting the specific binding affinity to carbohydrates (Wang and Anzai, 2015). The large superfamily of proteins which recognize a diverse range of ligands and defined by the presence of at least one are C-type lectin-like domain (CTLD) are C-type lectin receptors (CLRs) (Dambuza and Brown, 2015). Particular interest is represented by a single extracellular CTLD-containing receptor clusters “Dectin-1 and Dectin-2,” which is associated with the signal adapters or possess integral intracellular signaling domains. CLRs have traditionally been associated with the recognition of fungi, but recent research has shown unexpected and diverse functions. Their new role in antimicrobial defense, homeostasis, autoimmune reactions, and allergies and their functions in the recognition and response to dead and cancer cells has been described (Dambuza and Brown, 2015). The first step in microbial pathogenesis is binding to the host cell which is often mediated by human cellular carbohydrate - bacterial lectin (adhesin) interaction. A new strategy based on the construction of a network of interactions of lectin-glycans (LGI) to identify potential receptors of human binding to pathogenic adhesins with lectin activity was described (Ielasi et al., 2016). Linking the results from the screening of glycan arrays of these adhesines with the human glycoprotein database by constructing an LGI network is a new and perspective approach. This strategy has been used to detect human receptors for pathogenic fungi Candida albicans (Als1p and Als3p adhesins) and C. glabrata (Epa1, Epa6, and Epa7 adhesins) that cause candidiasis and for a virulent strain of E. coli (FimH adhesion). The LGI network approach allows us to profile and prioritize potential lectin-binding receptors in the host based on the experimental data. New potential targets were predicted and confirmed experimentally for a number of selected adhesines. This strategy has also been used to predict the interaction of lectins

9.1 Introduction

with enveloped glycoproteins of human pathogenic viruses and in identifying anti-HIV activity of FimH adhesin (Ielasi et al., 2016). Lectin microarrays have proved to be useful in studying multiple lectin glycan interactions in a single experiment and, with the advances made in the field, hold promise for enabling glycomic profiling of cancers in a fast and efficient manner (Syed et al., 2016). A bead-based lectin matrix was developed to increase the sensitivity of glycosylation profiling. Lectins are chemically bonded to microspheres coated with a fluorescent dye, and glycan-lectin recognition is carried out in three dimensions. The effectiveness of this platform was evaluated, and the detection limit (LOD) of R. communis lectin (RCA120) was 50 pg mL21 (1 pM) of asialofetuine, providing the lectin bead-based microchip with the highest sensitivity among the registered lectin microchips (Wang et al., 2014a). In addition, a multiplex analysis was carried out which made it possible to simultaneously detect several carbohydrate epitopes in one reaction vessel. Increased (α-1,6) core fucosylation and (α-2,6) sialylation patterns were observed under the glycosylation patterns of hepatocellular carcinoma associated immunoglobulin G analyze, these results may provide significant clinical evidence for disease diagnosis (Wang et al., 2014a). In one study (Nakajima et al., 2015), glycoproteins were isolated from formalin-fixed, paraffin-embedded tumor specimens and normal epithelium from 53 consecutive curatively resected stage I III colorectal cancer cases and then to obtain lectin glycan interaction (LGI) values was subjected to lectin microarray. Clinicopathological factors associated with distant recurrence were also identified. LGI values associated with long-term relapse were confirmed in 55 cases of stage II colorectal cancer by resection. LGI values for lectin Agaricus bisporus (ABA) excreted in cancer tissues were statistically associated with long-term relapse. ABA staining with lectin showed strikingly intense signals in the cytoplasm and apical surfaces of cancer cells, while weak staining was observed in the case of normal epithelium. This ABA tumor to normal LGI ratio may be a new prognostic biomarker of long-term relapse of cured, resected colorectal cancer. The clinical benefits of novel predictive markers for distant recurrence with colorectal cancer using lectin microarray analysis of cell surface glycan modifications was evaluated (Nakajima et al., 2015). Lectins are present in plants (high availability), microorganisms, animals, and humans. In plants these proteins more often used as models for the study of lectin-carbohydrate interaction. Plant lectins are divided into different taxonomic families: Legumes, Cereals, Amaryllidaceae, Moraceae, Amaranthaceae, Euphorbiaceae and Loranthaceae, Solanaceae, Labiatae, and Urticaceae. There is a need for the creation of new, simple, rapid, and inexpensive analytical tools for sensing, analysis, and purification of lectins. Lectin carbohydrate interactions are often rather weak (Ka 5 102 103 M21) (Liener et al., 1986), but can be dramatically enhanced using multivalent carbohydrates manifesting a so-called cluster effect (Pieters, 2009). Most soluble lectins

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have a multimeric structure, thus, enabling interaction with a few carbohydrate ligands and substantially increasing binding power. The cluster effect may be achieved by use of neoglycoconjugates—synthetic or semisynthetic (macro) molecules bearing a number of identical carbohydrate ligands. Neoglycoconjugates are usually prepared either by polymerization of glycomonomers or functionalization of different polymeric matrixes with free carbohydrates, or, more often, with glycosynthons. Controlled/living radical polymerization techniques (nitroxide-mediated polymerization, atom transfer radical polymerization, reversible addition-fragmentation chain transfer, and radical and ring-opening polymerization) in combination with other routes, such as click reactions, as well as various types of block copolymerization, and a wide range of other approaches for macromolecular architectures creation were used to prepare neoglycoconjugates based on synthetic polymers (Godula and Bertozzi, 2010; Chen et al., 2007; Wang et al., 2014b; Kiessling et al., 2006; Sun et al., 2014; Zhang et al., 2013, 2014; Lele et al., 2005; Yilmaz and Becer, 2014; Pati et al., 2012; Parry et al., 2013; Bovin, 1998; Richards et al., 2016). Bovin’s review (Bovin, 1998) describes the synthesis, physicochemical characteristics, and application for studying carbohydrate-binding proteins of polyacrylamide type neoglycoconjugates. Godula and Bertozzi (2010) developed a poly(acryloyl hydrazide) scaffold to which we conjugated a variety of reducing glycans ranging in structure from simple mono- and disaccharides to considerably more complex human-milk and blood oligosaccharides. The conjugation proceeds in a stereoselective manner, providing glycopolymers with pendant glycans accommodated mostly in their cyclic β-glycosidic form. Fluorescent statistical glycopolymers were synthesized via reversible additionfragmentation chain-transfer polymerization (RAFT) and successfully employed in lectin-mediated bacterial binding studies. The resultant glycopolymers contained three different monomers—N-(2-hydroxyethyl) acrylamide (HEAA), N-(2-aminoethyl) methacrylamide (AEMA) and N-(2-glyconamidoethyl)methacrylamides—possessing different pendant sugars. Low predictable degrees of polymerization were observed among the products. After polymerization, the glycopolymers were further modified by different succinimidyl ester fluorophores targeting the primary amine groups on AEMA (Wang et al., 2014b; Guo et al., 2015). To explore the effect of polymer structure on their self-assembled aggregates and characteristics, the study was devoted to developing a series of amphiphilic block and random phenylboronic acid-based glycopolymers by RAFT polymerization. Amphiphilic triblock copolymers containing conjugated polyfluorene as the middle block and glycopolymer as the side blocks have been designed and prepared, which could assemble into fluorescent nanoparticles with the sugar isomers on the surface. Sun et al. (2014) Multiblock glycopolymers made of di(ethylene glycol) ethyl ether monomers and carbohydrates (mannose or glucose) were synthesized, and these highly monodisperse glycopolymers were then used in different systems for binding and inhibition of DC-SIGN, the protein important for HIV infection (Zhang et al., 2013). A technique was developed to synthesize protein 2 polymer conjugates by initiating atom transfer radical

9.1 Introduction

polymerization of monomethoxy poly(ethylene glycol) 2 methacrylate from 2-bromoisobutyramide derivatives of chymotrypsin (a protein 2 initiator). Polymerization initiated from the monosubstituted protein 2 initiator resulted in the conjugate containing a single, near-monodispersed polymer chain per protein molecule with a polydispersity index 1.05 (Lele et al., 2005). The synthesis of high molecular weight, water-soluble, O-glycopolypeptide polymers by the ringopening polymerization of their corresponding N-carboxyanhydride (NCA) was reported. The per-acetylated-O-glycosylated lysine-NCA monomers, synthesized using stable glycosyl donors and protected amino acid, was polymerized using amine initiators (Pati et al., 2012). A polymerizable version of the Tn-antigen glycan was prepared and converted into well-defined glycopolymers by RAFT polymerization. The polymers were then conjugated to gold nanoparticles, yielding “multicopy-multivalent” nanoscale glycoconjugates (Parry et al., 2013). The combination of copper-mediated living radical polymerization (Cu(0)-LRP) with thiol halogen, thiol epoxy, and copper-catalyzed alkyne azide coupling (CuAAC) click chemistry was employed to provide a new route to multiblock sequence-controlled glycopolymers. Multiblock poly(glycidyl acrylate)-co-(acrylic acid 3-trimethylsilanyl-prop-2-ynyl ester) (poly(GA)-co-(TMSPA)) were obtained by Cu(0)-LRP in DMSO at ambient temperature via iterative monomer addition whereby the sequence of the multiblocks is attained in a designed way. Thiol halogen and thiol epoxy reaction of poly(GA) have been exploited, which suggested a preference for the reaction of the halogen rather than the epoxide for the thiol with triethyl amine as catalyst. The obtained multiblock poly(GA)-co(TMSPA) was then used for CuAAC and sequential thiol halogen (epoxy) reactions to build functional glycopolymers in a defined sequence. Zhang et al. (2014) The glycan microarray-oriented and density-controlled glyco-macroligand microarray based on endpoint immobilization of glycopolymer that was accompanied with boronic acid (BA) ligands in different sizes as detachable “temporary molecular spacers” was prepared. In brief, firstly an O-cyanate chain-end functionalized lactose-containing glycopolymer was precomplexed with polyacrylamide-BA, lysozyme-BA, and bovine serum albumin (BSA)-BA conjugates as macromolecular spacers and, then secondly, immobilized onto an amine-functionalized glass slide (Narla and Sun, 2012). Recently, spacered mono- and diantennary glycopolymers were used to study the interaction with Con A. Preparations of the highly ordered monoantennary, homofunctional diantennary, and heterofunctional diantennary neoglycopolymers of α-D-mannose and β-D-glucose residues were achieved via ring-opening metathesis polymerization (Loka et al., 2015) and it was also established that alkyl spacers enhance lectin binding of glycopeptides based on poly(glutamic acid) (Mildner and Menzel, 2014). Despite of the wide range of glycopolymers that are accessible via polymerization of sugar-containing monomers, there are some limitations associated with these polymerization strategies. Recent significant achievements in using reactive glycosynthons present great opportunities for extending of the field of

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neoglycoconjugates synthesis, and a great potential for further development in this area exists (Garcia et al., 2011). For lectin detection and biosensing, metal-labeled glycoconjugates are often used. Gold and silver nanoparticles (NPs) are intensively employed in fundamental science and applied fields because of a number of unique physicochemical properties. Due to their spectral characteristics, especially plasmon resonance, (Chow, 2010; Sau and Goia, 2011; Welles, 2010; Pereira et al., 2014) gold and silver NPs may be applied as signal transducers in biospecific sensors changing their spectral characteristics upon the binding of a ligand with the analyte (Liu et al., 2011; Jans and Huo, 2012; Xing et al., 2014; Dykman and Khlebtsov, 2012). Polymeric neoglycoconjugates bearing gold and silver NPs are called glyconanoparticles (GNPs). GNPs may bear varying amounts of carbohydrate ligands, and principally modulate cellular surfaces (Fuente and Penade´s, 2006). These GNPs are widely used for lectin (agglutinin) binding studies. Gold GNPs, for example, may be used in lectin dot-blot assays or binding studies imitating erythrocytes in agglutination (Piskarev et al., 2003). Gold nanoparticles have an intense red-violet color, thus, enabling their direct visual detection. Gold GNPs most often are prepared from the appropriate carbohydrate derivatives and citratestabilized gold NPs (Marradi et al., 2013) and sometimes are not completely stable. Most often tiol-containing carbohydrate derivatives are used, and GNPs preparation involves a number of synthetic steps (Barrientos et al., 2003; Halkes et al., 2005; Reynolds et al., 2013; Toyoshima and Miura, 2009). For example, thiol-derivatives of neoglycoconjugates based on lactose, maltose, or glucose have been prepared for attachment to gold surfaces and synthesis of the disulfides—linkers involve glycosidation of conveniently protected oligosaccharide derivatives with 11-thioacetate undecanol or 11-thioacetate-3,6,9,-trioxaundecanol by means of the trichloroacetimidate method (Barrientos et al., 2003). A method has been established for the preparation of GNPs from free oligosaccharides (Halkes et al., 2005). The crucial part of this method is the two-step reaction sequence for the introduction of a thiol-spacer in the free oligosaccharides. Via reductive amination, trityl-protected cysteamine was introduced into the glycan. After removal of the trityl group, thiol-spacered oligosaccharides were used for the GNPs preparation. According to Toyoshima and Miura (2009), glycopolymersubstituted gold nanoparticles were prepared via living radical polymerization with a reversible addition-fragmentation chain-transfer reagent. Polyacrylamide derivatives with α 2 mannose (α-Man) and N-acetyl-β-glucosamine were synthesized and hydrogenated to obtain a thiol-terminated polymer. The thiol-terminated glycopolymers were mixed with gold nanoparticles to yield the polymersubstituted gold nanoparticles with various diameters, which aggregated upon the addition of saccharide-recognition proteins. We will present two types of glycomaterials based on maleic anhydride copolymers: (1) water-soluble gold- or silver-containing GNPs for lectin detection (Scheme 9.1); and (2) crosslinked neoglycoconjugates for lectin sorption (Scheme 9.2). A simple approach for their synthesis entails a single stage—covalent

SCHEME 9.1 Schematic representation of lectin sensor preparation.

SCHEME 9.2 Schematic representation of lectin sorbent preparation.

9.1 Introduction

coupling of free, or amino spacered carbohydrate, with soluble, or crosslinked maleic anhydride copolymers through reactive anhydride functions without use of condensing agents. For lectin sensing, a color label (nanosized gold or silver) was synthesized directly into the water-soluble colloidal neoglycoconjugate matrix, and the separate stage of preparation of nanometal protecting shell with use of citrate, alkyl thiol, or phosphine, etc., was excluded. Microspherical crosslinked particles of maleic anhydride copolymers bearing carbohydrate derivatives were prepared for lectin sorption studies. The maleic anhydride copolymers used had a few advantages, such as commercial availability, or possibility of simple synthesis according to known procedures, and well-defined structural features (strict alternation of co-monomer units). Copolymers of maleic anhydride are widely used in technology, for example, copolymers of maleic anhydride with styrene (SMA) normally in impact-modified and optional glass fiber-filled variants is available as a crystal-clear granule that can be used in a wide variety of applications. Alternatively, SMA is applied using its transparency in combination with other transparent materials to heat-boost other polymers. It was shown that maleic anhydride copolymers had no toxicity. Cytotoxicity of maleic anhydride copolymers with styrene, vinyl acetate, and methyl methacrylate was evaluated by using a mouse fibroblast cell line (L929). Karakus et al. (2013) Most importantly, copolymers of maleic anhydride may be easily converted into copolymers of maleic acid (in water), and resulting copolymers of maleic acid may be transferred again into an active anhydride form (under thermal effect on dry maleic acid copolymers). Anhydride groups of the copolymers (soluble or crosslinked) enable a smooth reaction either with hydroxyl or amino groups of the appropriate carbohydrates, or carbohydrate derivatives (Auzely-Velty et al., 2002). For example, the complexes containing glucose oxidase were synthesized by the covalent immobilization of enzymes in aqueous medium (in the absence of toxic chemicals and solvents) on the shell of the alternate maleic acid copolymer covering a core made of nanogold or its alloy (Samoilova et al., 2014). Maleic acid copolymers were used for the creation of interpolyelectrolyte complex shells for the stabilization and functionalization of different nano- and microparticles. (Samoilova et al., 2010; Krayukhina et al., 2010; Samoilova et al., 2014) The biomedical application of maleic acid (anhydride) copolymers is based on their structure and activity. The anhydride cycle offers the possibility to obtain conjugates with different drugs by chemical reactions at mild conditions. These copolymers can be used for different purposes: as supports for bioactive molecules (conjugates, films, solid dispersions, micro/nanoparticles), in drug controlled release systems, and also as components of biomaterials (in dental applications or in tissue engineering) (Popescu et al., 2011). Conjugates of maleic anhydride copolymers often have a pH-sensitive solubility (better solubility at pH . 7), so can be used to release biologically active compounds in certain segments of the gastrointestinal tract (intestine). Drug delivery systems based on maleic acid copolymers have been described. Bacu et al., (2002) presented their results regarding the synthesis of new derivatives of phenothiazine with potential pharmacological properties, and for

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obtaining controlled release systems by their reaction with maleic anhydride copolymers. One of the most studied polymers with per se activity in numerous in vitro and in vivo experiments is the copolymer of maleic anhydride with divinyl ether (DIVEMA). Preliminary results have shown DIVEMA as active against Friend leukemia virus, adenocarcinoma 755, Lewis lung carcinoma, and Dunning ascites leukemia (Breslow, 1976). It has been shown that antitumor activity of DIVEMA can be explained by its capacity, and the copolymer can also be used as an adjuvant in chemotherapy (Pearson et al., 1974). It has been shown that the prophylactic treatment with DIVEMA protects the body from the viral infection. DIVEMA has been active against over 20 viruses, including cancer-inducing viruses, Friend leukemia, Rauscher leukemia, Moloney sarcoma, and others. The effect of DIVEMA against Rauscher leukemia was studied on both immunosuppressed and normal mice and in both cases it was found out that DIVEMA caused a dramatic reduction in the virus titer, and these are not the only results of stimulating host immune response. The interferon induction activity of DIVEMA was responsible for prophylactic action. The antifungal activity of DIVEMA was demonstrated against Cryptococcus neoformans and antibacterial activity against both gram-positive (Remington and Merigan, 1970) and gram-negative (Giron et al., 1972) bacteria. Conjugates of poly(styrene-alt-maleic anhydride) with therapeutic agents such as amlodipine, amantadine hydrochloride, zonisamide, gabapentin, and mesalamine were prepared by the formation of the amide bonds of the amino groups of drugs with the polymeric anhydride groups (Khazaei et al., 2013). The amounts of covalently conjugated drugs were determined by a 1H NMR method, and the release rate was studied in vitro at pH 1.3 and temperature 37 C. The different models to obtain drug-release data were examined in kinetic studies under in vitro conditions and the obtained data were well-fitted to the Korsmeyer-Peppas equation, revealing a dominant Fickian diffusion mechanism for drug release. Other pharmaceutical applications of maleic anhydride copolymers (especially with alkyl vinyl ethers in half-ester form) in drug release entail the preparation and use of composites where the polymer does not interact chemically with the drug. These composites include micro/nanoparticles, micelles, films for tablet coating, or solid dispersions. Maleic acid (anhydride) copolymers may also be used as components in analysis kits, biomaterials, and tissue engineering (Popescu et al., 2011). In this chapter, we described how water-soluble silver or gold GNPs were prepared from neoglycoconjugates (synthetic glycoconjugates (GC)). GCs, in turn, were obtained from maleic anhydride copolymers, and either free N-acetylD-glucosamine (GlcNAc), or glycosynthons (glycyl-spacered ligands)—β-Nglycyl-N-acetyl-D-Glc (N-Gly-GlcNAc) and N-Gly-β-D-Gal-β1-4Glc (N-Gly-lactose) (Scheme 9.3A GC-1, GC-2 and GC-3), respectively. Crosslinked neoglycoconjugates also contained glycyl-spacered ligands (Scheme 9.3A CLGC-1 and CLGC-2). Glycyl-spacered glycosynthons N-Gly-GlcNAc and N-Gly-lactose were prepared using uncomplicated, two-stage synthesizes and with 70% 80% yield.

9.1 Introduction

SCHEME 9.3 (A) Schematic representation of water-soluble polymeric neoglycoconjugates.

To estimate the specificity and activity of synthesized neoglycoconjugates, their interaction with lectins from different taxonomic families was studied. Taxonomic families included β-D-GlcNAc-specific lectins from potato (Solanum tuberosum agglutinin, STA—Solanaceae family of lectins) wheat (Wheat germ agglutinin, WGA—Cereal family of lectins), pokeroot (Phytolacca americana agglutinin, PAA (PWM)—Cereal family of lectins), β-D-Gal-specific lectins from castor beans (R. communis agglutinin, RCA 120—Euphorbiaceae and Loranthaceae family of lectins), and arachis (Arachis hypogaea agglutinin, AHA (PNA)—Legumes family of lectins) and were studied using the dot-blot technique and UV-VIS spectroscopy. Crosslinked neoglycoconjugates were used as models for elucidation of lectin-carbohydrate binding, for lectins isolation, and sorption.

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SCHEME 9.3 (Continued) (B) Schematic representation of crosslinked neoglycoconjgates.

9.2 Experimental

9.2 EXPERIMENTAL 9.2.1 MATERIALS Poly(ethylene-alt-maleic anhydride) (EMA) with an average molecular weight M 5 25 kDa was purchased from Monsanto (USA), while poly(N-vinyl-pyrrolidonealt-maleic anhydride) (VMA, M 5 40 kDa) was prepared earlier (Conix and Smets, 1955). Upon dissolving in water, copolymers turned into poly(ethylene-alt-maleic acid) (EM) and poly(N-vinyl-pyrrolidone-alt-maleic acid) (VM). Crosslinked spherical granulated maleic anhydride copolymers were obtained earlier (Samoilova et al., 1981). Two copolymers in equimolar amount, VMA and poly (styrene-alt-maleic anhydride), were crosslinked with 10% (mol.) 4,4v-diaminodiphenyl oxide. The solutions of the two copolymers and crosslinking agent in DMF were vortexed for 1 2 minutes at room temperature and then poured into silicon liquid. Emulsion was intensively stirred for 3 hours and incubated for 16 hours at room temperature; supernatant was decanted, and the pellets were separated by filtration, washed with ether, (CH3CO)2O DMF (1/1, v/v), dry acetone, ether, and dried under vacuum. Lectins from wheat germ (WGA), potato (STA), pokeroot (PAA), arachis (AHA), and castor beans (RCA 120), as well as BSA, AgNO3, HAuCl4, NaBH4, D-GlcNAc, and lactose were Sigma-Aldrich products. All other reagents and salts were of analytical grade (Sigma-Aldrich) and used without further purification. Milli-Q-purified water was used. PVDF transfer membranes were obtained from “Immobilon-P” (Millipore, USA). Dissociation constants of lectin-neoglycoconjugate complex for both types of colloidal glycoparticles were calculated using equilibrium data of plasmon resonance maximum shift in UV-VIS absorbance spectra, or ion-exchange fractionation of lectins and lectin-neoglicoconjugates (λ 280 nm, Toyoperl DEAE-650S (Tosoh, Japan)).

9.2.2 INSTRUMENTATION UV-visible absorption spectra were obtained using UVIKON-922 spectrophotometer (Germany). pH values were determined using Fisher Scientific 300 403.1 pH-meter (USA). Transmission electron microscopy (TEM) micrographs were performed with LEO 912 AB microscope (Omega, Karl Zeiss; Germany) operated at an accelerating voltage 100 kV. For TEM observations, a drop of colloid solution was placed onto a Formvar-coated copper grid and then evaporated. The particle size distribution was obtained from a count of 200 300 individual particles. Dynamic light scattering (DLS) experiments were performed on a PhotoCor Complex instrument (PhotoCor, Russia) equipped with an automated goniometer, a PhotoCor-PC2 pseudocorrelative system for the photon count, a PhotoCor-FC one-plate real time multitemporal correlator, and a Uniphase 1135 P helium-neon

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laser (10 mW) with a wavelength of 633 nm. Dust from the samples was removed by filtration through a Spartan filter with a pore size of 0.45 μm. 1 H NMR spectra at 500 MHz and 2D-COSY spectra at 600 MHz were recorded in D2O on a Bruker DRX500 SF 5 500.13 MHz instrument at T 5 299 K. IR spectra (KBr) were recorded on a Fourier-spectrometer Magna IR-720 (Nicolet, USA). Optical microscopy of crosslinked polymeric sorbents was performed on an Eclise H550S (Nikon, Japan) microscope equipped with a Kodak DC 120 Digital Camera (USA). Dot-blot assays were performed on Bio-Dot device (Bio-Rad, USA).

9.2.3 METHODS 9.2.3.1 Synthesis of N-glycyl-β-glycopyranosylamines Synthesis of N-glycyl-2-actamido-2-deoxy-β-D-glucopyranosylamine (N-Gly-GlcNAc) Firstly, 2-Acetamido-N-chloroacetyl-2-deoxy-β-D-glucopyranosylamine was prepared from 2-acetamido-2-deoxy-β-D-(3-O-glucopyranosylamine under the action of (ClCH2CO)2O in DMF according to the procedure reported in Likhosherstov et al. (1996). The yield was 69%, m.p. 223 225 C (from a EtOH-EtOAc mixture, with decomp.), [α]D20 129.1 degrees (c1, H2O). 1H NMR δ: 2.02 (s, 3H, CH3); 3.46 3.57 (m, 2H, H(4), H(5)); 3.65 (t, 1H, H(3), J2,3 5 J3,459 Hz); 3.74 3.96 (m, 3H, H(2), H(6a), H(6b)); 4.15 (br.s, 2H, CH2Cl); 5.11 (d, 1H, H(1), J 5 9 Hz); cf lit. data(Paul et al., 1980): m.p. 220 221 C, [α]D20 1 28.2 degrees (c1, H20). 2-Acetamido-N-chloroacetyl-2-deoxy-β-D-glucopyranosylamine (4 mmol) was dissolved in a 30% aqueous solution of NH3 (100 mL). The reaction solution was kept at 10 C for 40 hours and concentrated to 20 mL. Then Dowex 1 3 8 (OH2) anion-exchange resin (15 mL) was added and the reaction mixture was stirred for 30 minutes. The resin was filtered off and washed with H20 (150 mL). The filtrate and washings were concentrated to 50 mL. Then AcOH was added to pH 5 and the reaction mixture was kept at 20 C for 16 hours. KU-2 (H1) cationexchange resin (30 mL) was added and the suspension was stirred for 1 hours. The resin was filtered off and washed with water (300 mL) and 0.5 M pyridine in water (300 mL). The product was eluted with 2 M NH4OH (300 mL), the corresponding fractions (control by paper electrophoresis, pH4.5) were concentrated to dryness, and the residue was crystallized. The yield was 76%.

Synthesis of N-glycyl-4-O-β-D-galactopyranosyl-β-D-glucopyranosylamine (N-Gly-lactose) Lactose (0.06 mmol) and powdered ammonium carbamate (47 mg, 0.6 mmol) were dissolved in 25% aqueous solution of NH3 (0.6 mL), followed by the

9.2 Experimental

addition of MeOH (1.2 mL) and stirring and keeping for 24 hours at 37 C in a test tube with a screw stopper. The reaction mixtures were diluted with MeOH (5 mL) and concentrated to B0.5 mL at B40 Torr. This operation was repeated for another B5 times, monitoring the completeness of the ammonium carbamate removal using an indicator paper placed over the solution of the compound in MeOH, and the last concentration at B10 Torr was carried out to dryness. The residue was dried and the obtained lactosylamine was immediately subjected to N-acylation. Lactosylamine (B0.06 mmol) was dissolved in H2O (0.125 mL), cooled with ice, followed by the addition of N-hydroxysuccinimide ester of N-Boc-glycine (49 mg, 0.18 mmol) in DMF (0.5 mL), stirring and keeping for 2.5 hours at B20 C. The reaction mixtures were diluted with MeOH (6 mL), concentrated to B0.5 mL and this operation was repeated three times. Diethyl ether (10 mL) was added to the solutions obtained with stirring. The liquids after clearing were separated by decantation from oily precipitates. The precipitates were washed several times with Et2O (3 mL each) and a mixture of Et2O—acetone (1:1) and dried. The residues were dissolved in H2O (1 mL) and subjected to chromatography on columns with silica gel C18 in H2O loading 10 mg of the compound per 1 g (B3 mL) of silica gel. The columns were washed with H2O until the UV absorption decreased to a minimum value, and then washed with 25% aqueous MeOH. The aqueous methanol fractions containing the desired compounds were combined and concentrated to dryness. The residues were dissolved in 10% solution of Et3N in 50% aqueous MeOH (2 mL) and kept for 3 hours at B20 C. The reaction mixtures were diluted with MeOH (7 mL), concentrated to dryness, and this operation was repeated twice. The residues were dissolved in H2O and subjected to chromatography on columns with silica gel C18 under the same conditions. The aqueous methanol fractions containing the desired compound were combined and concentrated to dryness. The residues were dried to obtain N-(NBoc-glycyl)-lactosylamine (29.4 mg, 76%). Trifluoroacetic acid (1 mL) was added to N-(N-Boc-glycyl)-lactosylamine (B0.04 mmol each), the mixture was stirred for B5 minutes at 20 25 C until dissolution. The solution was concentrated nearly to dryness at B10 Torr. The residue was diluted with toluene (10 mL), stirred, and concentrated nearly to dryness. The residues were diluted with MeOH (5 mL) with stirring, followed by the addition of toluene (10 mL) and concentration almost to dryness. This operation was repeated four more times, monitoring the completeness of the acid removal using an indicator paper placed over the solution of compound in MeOH. The residue was dissolved in MeOH (10 mL) and concentrated to dryness to obtain trifluoroacetates of compound. A Dowex 21K (OH2) (50 mg) was added to the solution of salt in H2O (1 mL) and the suspension was stirred for 45 minutes. The resin was filtered off and washed with water (4 3 0.5 mL). The solution was concentrated to B1 mL, filtered through a membrane filter (0.45 μm), concentrated to B0.3 mL, and lyophilized. The residue was dried. Amorphous powder

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of N-Glycyl-lactosylamine was obtained in B95% yields, m.p. 246 248 C, [α]D 20 13.0 (c 1, H20). Found (%): C, 42.37: H, 6.46; N, 7.06. C14H26N2O11. Calculated (%): C, 42.21: H. 6.58: N, 7.03. 1H NMR, δ:3.42 (br.s, 2H, NCH2); 3.45 3.60 (m, 2H); 3.64 3.88 (m, 8H); 3.92 3.99 (m, 2H); 4.48 (d, 1H. H(1) Gal, J 5 8 Hz); 5.04 (d, 1H, H(1) GIc, J 5 10 Hz).

9.2.3.2 Synthesis of neoglycoconjugates and glyconanoparticles Carbohydrate-polymer ester bond formation: general procedure EM GlcNAc synthesis (GC-1, Scheme 9.3A). For activation of the polymerstabilizer EM, a powdered sample was heated over diphosphorus pentoxide (P2O5) (3 hours, 110 C) in vacuo. Suspension of D-GlcNAc (0.20 g in 3 mL DMF and 0.5 mL pyridine) was added with vigorous stirring to the activated polymer (0.20 g in 5 mL DMF). After stirring (2 hours, B100 C) solution was cooled, threefold diluted with 20 mL water, concentrated by ultrafiltration, and then freeze-dried. The degree of carbohydrate substitution (elemental analysis) was 25 % (weight).

Carbohydrate-polymer amide bond formation: general procedure EM Gly-GlcNAc (GC-2.2, Scheme 9.3A) and EM Gly-lactose (GC-3, Scheme 9.3A) synthesis. For the activation of the polymer-stabilizer EM, a powdered sample was heated over P2O5 (3 hours, 110 C) in vacuo. To the activated polymer (0.17 g in 3 mL DMF) solution of 0.15 g N-Gly-GlcNAc, or 0.16 g N-Gly-lactose, in 10 mL H2O, and 2 mL of 0.1 M NaHCO3 were added with vigorous stirring. After stirring (24 hours, B20 C), the solution was threefold diluted with 20 mL water, concentrated by ultrafiltration, and then freeze-dried. Degree of substitution (elemental analysis, or NMR) was 24% or 29% (weight), respectively. Neoglycoconjugate GC-2.1 was synthesized identical. Degree of substitution was 19% (weight).

Synthesis of crosslinked glycoconjugates (CLGC-1 and -2, Scheme 9.3B): general procedure A crosslinked maleic anhydride copolymer (0.8 g) was added with vigorous stirring to the solution of 0.18 g of N-Gly-GlcNAc, or 0.10 g N-Gly-lactose, in 5 mL H2O at pH 8 (titration with 0.1 M NaHCO3). After stirring (24 hours, B20 C), polymers were filtered off and washed, firstly, with sodium phosphate buffer (PBS, 0.1 M, pH 7.0) until absorbance became less 0.01 at 280 nm and, secondly, with H2O. Crosslinked glycoconjugates were used as prepared, or washed with acetone, and dried. Degree of substitution, determined by MorganElson reaction,(Morgan and Elson, 1934) was 0.60 mM g21 for N-Gly-GlcNAc ligand, and 0.25 mM g21 for N-Gly-lactose ligand.

9.2 Experimental

Synthesis of silver, or gold glyconanoparticles (Scheme 9.1) Colloidal solutions of nano-sized silver and gold GNPs were obtained by the borohydride reduction of the metal salts at B20 C in the presence of polymeric neoglycoconjugates (GC 1-3, Scheme 9.3A). Samoilova et al. (2014) A proper amount (for the required ratio of reagents) of freshly prepared solution of AgNO3, or HAuCl4•3H2O (0.1 M), was added to the freshly prepared solution of GC (0.01 M, here, the molar concentration of copolymer refers to the monomeric maleic acid units) in water at pH 6.3 (for the preparation of gold NPs; it is pH of the initial GC solution (for the preparation of gold NPs), or pH 7 (for the preparation of silver NPs, titration with 5% aqueous NaOH) with vigorous stirring. After 5 10 minutes, a freshly prepared aqueous solution of NaBH4 (0.1 M, twofold molar excess with respect to silver ions, or threefold molar excess with respect to gold ions) was added to the polymeric salt with vigorous stirring. The reaction mixture was allowed to stand (24 hours, B20 C). Dried samples of polymeric GNPs (GC 1-3/Ag (Au0)) were obtained after ultrafiltration (membrane filter 10 kDa, Hydrosart, “Sartorius Stedim Biotech”) and freeze-drying.

9.2.3.3 Lectins binding assays Dot-blotting Detection of lectins using gold or silver GNPs was performed after lectin blotting on polyvinylidene difluoride (PVDF) membranes in a Bio-Dot device; the membranes were preliminary activated for 20 minutes in MeOH, and then washed with water. Ligand-protein binding was performed at room temperature. Aqueous lectin solution (5 μL, 0.0002 4.0 mg mL21, Tris-buffered saline (TBS), 0.02 M, pH 7.4)) was spotted into each well, allowed to dry (1 hours), and then 0.03 mL of gold (silver) GNPs (1 mg mL21) was added. After 3 minutes of incubation, each well was washed with 1 mL of 20 mM TBS (pH7.4), and the membrane was air-dried. Yellow-brown (silver) or red-brown (gold) staining was indicative of lectin-GNPs interaction.

UV-visible absorbance measurements Optical spectra of gold GNPs were registered at concentration 1 mg mL21, (0.02М TBS, pH 7.4), and silver GNPs at 0.1 mg mL21 (0.01М PBS, pH 6.5). Spectra changes were registered after addition of 5 40 μL of lectin (2 mg mL21, 0.02М TBS, pH 7.4, 25 C).

Binding properties of crosslinked neoglycoconjugate sorbents Biospecific sorption of crosslinked neoglycoconjugates (CLGC, Scheme 9.3B) was estimated as a difference between optical densities of the lectin starting solution and lectin supernatant solution after its contact with the sorbent at λ 280 nm. To the solution of a lectin (WGA, AHA, PAA, RCA 120, or STA in 2 mL of 0.02 M TBS, pH 7.4) 0.4 mL (12 mg) of suspension of CLGC containing 2.7 mg of ligand—N-Gly-GlcNAc, or 14.5 mg of CLGC containing 1.8 mg of

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N-Gly-lactose—was added. Optical density of the supernatant solution was measured after 24 hours (equilibrium).

9.3 RESULTS AND DISCUSSION 9.3.1 SYNTHESIS OF NEOGLYCOCONJUGATES AND METAL-LABELED GLYCONANOPARTICLES Colloidal GNPs (Scheme 9.1) were prepared in two stages: (1) synthesis of polymeric neoglycoconjugates (GCs); and (2) introduction of nanometal into GCs. All neoglycoconjugates—polymer-sugar derivatives—were obtained in organic (GC-1, Scheme 9.3A) or water-organic media (GC-2 and GC-3, Scheme 9.3A). Maleic anhydride groups of the copolymers used easily reacted with one of hydroxy groups of GlcNAc (GC-1, Scheme 9.3A), or primary glycine amino group of ligands N-Gly-GlcNAc and N-Gly-lactose (GC-2 and GC-3, Scheme 9.3A) in one stage. Thus, two types of GCs were prepared: GC-1 contained D-GlcNAc conjugated to polymer via ester bond; GC-2 and GC-3 obtained via formation of amide bond between glycyl-spacered carbohydrate and copolymer of maleic anhydride, without using any condensing agents. Resulted GCs contained 10% 11% of carbohydrates (mol), and epitopic density of glycan was about 1 carbohydrate unit per 10 11 maleic anhydride residues of the polymer (k/m 5 10 11, Scheme 9.3A). Thus, the distance between carbohydrate ligands ˚ , as distance between maleic acid residues in the copolymers was 50.2 55.2 A ˚ was 5.02 A (this value was calculated for an “ideal” polymer chain, in which the ˚ and 109 degrees, distance between carbon atoms and the bond angle are 1.54 A respectively, similar to those in polyethylene). The influence of ligand presentation on recognition by the mannose-specific lectin BC2L was investigated earlier (Toyoshima and Miura, 2009; Godula and Bertozzi, 2012; Reynolds et al., 2013). Higher ligand density may result in the lowering of binding ability of the neoglycoconjugates, thus, glycan valency can set thresholds for linking by lectins. Toyoshima and Miura (2009), (Godula and Bertozzi, 2012) To obtain silver (GC/Ag0) or gold (GC/Au0) colloidal GNPs, NaBH4 was used as a reducing agent for the metal cations-precursors in presence of the stabilizing neoglycoconjugate. This method provided formation of compositionally uniform and small particles, due to the noticeable difference of the redox-potentials of the reducing agent (NaBH4) and metal with an excess of the former in the system. The preparation of nanosilver, using, for example, glucose and gelatin as the reducing and stabilizing agents resulted in not monodispersed and large size of silver nanoparticles (Pulit and Banach, 2013). The process of the reduction of silver ions in the presence of maleic acid copolymers was studied by Samoilova et al. (2009). The mechanism of silver NPs formation in the presence of dicarboxylic acid copolymers differs markedly from that in the case of polymers of monocarboxylic acids. The first step of the silver

9.3 Results and Discussion

hydrosol creation is the formation of polymer, silver salt. At the used conditions (pH 6 2 9) carboxyl groups of maleic acid units may donate protons so that an electrostatic attraction occurs between metal ions and acid residues. Binding capacity value of maleic acid copolymer with respect to silver ions and cooperativity of polymer-silver ions binding process were assessed on the ground of the connecting isotherm curve of maleic acid copolymer and Ag1 binding. Samoilova et al. (2009) The coefficient “n” (Hill coefficient) describing the degree of cooperativity was calculated from the Hill equation:
  lg f =1 2 f 5 n lg Ag1 f 1 lgKd ;

where f 5 [Ag1]bonded/[CMA], [Ag1]f is the molar concentration of unbound Ag.1 All parameters were calculated for mononuclear binding process without taking of activity coefficient of interacting molecules and Donnan effect into consideration. For poly(ethylene-alt-maleic acid) the Hill coefficient n 5 0.9 ( 6 0.1), under mild basic conditions. The above results suggest that the process of maleic acid copolymer binding of silver ions under used conditions is noncooperative. The dissociation constants of the polymer-silver salts were higher than those of initial polymer acids (e.g., pKd of copolymer/Ag1 5 2.4; cf. pK1 5 3.6 and pK2 5 6.2 for parent polymer acid). It could be proposed that under our reaction conditions (θ , 100%) one residue of maleic acid of copolymer binds no more than one ion of silver and metal-carboxylate complex is of a two-coordinate type. This mechanism of polymer-silver ion binding predetermines the mode of formation of intermediate clusters and NPs of silver in the course of reduction. Summing up, the early stage of silver nucleation apparently occurs with participation of polymer bonded silver ions within isolated centers of nucleation (nanoreactors). Then, the growth of silver particles within micelles of maleic acid copolymers leads to formation of compact, sterically stabilized polymeric globules. Here we used ratio 1:1:2 (mol) of the polymeric glycoconjugate/Ag1/NaBH4. Optical spectrum practically did not change, comparing to the one of silver NPs prepared at polymer/Ag1/NaBH4 ratio 1/1/10 (Samoilova et al., 2009); presence of the carbohydrates conjugated with the polymer chain also practically did not effect on the size of silver nanoparticles. An absorption band was observed due to the surface plasmon resonance of silver NPs and centered at around 400 401 nm. Depending of the nature of polymer shell, 50% 60% (weight) of silver was introduced into GCs, and thus silver GNPs (GC/Ag0) were obtained. In the process of gold nucleation, in presence of copolymers or GC used, optimal ratio of the polymer/chloroauric acid (5 7/1 mol/mol) (“sweet zone”) at 2 5-fold molar excess of NaBH4, with respect to the metal cation, was found. For preparation of gold GNPs (GC/Au0), copolymer/chloroauric acid ratio 5/1 (mol/mol) was

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FIGURE 9.1 Absorption spectra of EM-Gly-lactose/Au0 in water (pH7) at different incubation time: 0.5 h (1), 4 days (2), 8 days (3), 11 days (4), and 14 days (5).

used, which resulted in 14% 21% (weight) gold introduction. An absorption band of plasmon resonance of gold NPs was centered at around 520 521 nm. This band appeared after “ripening” of reaction solution for 10 11 days (Fig. 9.1). This type of “ripened” gold GNPs was used in all analytical experiments. Additional introduction of alkyl thiol groups (cysteamine residues) into the polymeric matrix practically had no influence on the properties of gold GNPs. Resulted silver and gold GNPs, after purification and freeze-drying, may be stored without any loss of properties for a long time, as was shown earlier for the same nonglycosylated metal NPs (Samoilova et al., 2009). In the preliminary tests with lectins, silver or gold GNPs were used without purification, as sodium, nitrate, and borate ions did not influence specific binding. It should be noted that another method of GNPs synthesis is possible: preliminary reduction of silver or gold cations in the presence of copolymers of maleic acid, and then, after thermal activation of polymeric shell of silver or gold NPs, conjugation of resulted nanocomposites containing maleic anhydride groups with carbohydrate ligands.

9.3.2 CHARACTERIZATION OF COLLOIDAL NEOGLYCOCONJUGATES AND GLYCONANOPARTICLES NMR spectra of the neoglycoconjugates are presented in Figs. 9.2 9.4. 1H NMR spectrum of EM-Gly-GlcNAc (Fig. 9.2), chemical shifts (δH, ppm): δH 5.09

9.3 Results and Discussion

FIGURE 9.2 1

H NMR spectrum of EM-Gly-GlcNAc.

FIGURE 9.3 1

H 2 1H COSY spectrum of EM-Gly-lactose.

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FIGURE 9.4 1

H NMR spectrum of EM-Gly-lactose.

(d, 1H, H-a, β-configuration); the signal of proton H-a cross-couples with H-b (3.76, dd, 1H, Jab 5 9.7 Hz); proton H-b cross-couples with H-a and H-c— Jab 5 9.7 Hz and Jbc 5 4.9 Hz. Triplet corresponding to H-c is present at 3.63 ppm with equal coupling constant Jbc 5 Jcd 5 4.9 Hz. Other signals correspond to:3.83 3.91 (m, 2H, H-f), 3.76 (dd, 1H, Jab 5 9.7 Hz and Jbc 5 4.9 Hz, H-b), 3.63 (t, 1H, H-c), 3.47 3.54 (m, 2H, H-e,d), 3.33 (s, 2H, H-h), 2.31 (m, w, H-i,j), 2.01 (s, 3H, H-g, COCH3), 1.39 (m, w, H-k,l), 0.83 (w, H from CH3 of polymer chain). 2D 1H-1H correlation spectrum of EMGly-lactose (Fig. 9.3): signal of proton of the methylene group H-m was identified at δH 3.92, and show a cross-coupling with saccharide part of glycopolymer (EM-Gly-lactose) at 3.9 3.6 ppm in COSY pattern (Fig. 9.3). A doublet at δH 4.46 corresponds to proton H-g of the β-Gal group, and crosscouples with triplet at δH 3.56 corresponded to H-h. Chemical shift corresponding to proton H-a of the β-Glc group of EM-Gly-lactose is clearly indicated downfield as doublet at δH 5.01, and its cross peak indicate couplings between protons H-a and H-b (t, δH 3.46). 1H NMR for EM-Gly-lactose (Fig. 9.4): 5.01 (d, 1H, Jab 5 9.2 Hz, H-a, β-configuration), 4.46 (d, 1H, Jgh 5 7.8 Hz, H-g, β-configuration), 3.92 (s, 2H, H-m), 3.56 (m, H-h), 3.46 (m, H-b), 2.3 1,0 (polymer chain), 2.30 (m, w, H-o,n), 1.39 (m, w, H-q,p). Chemical shifts in the range 3.9 3.6 ppm correspond to overlapping multiplet protons (10H, H-c,d,e,f, i,j,k,l) of the saccharide part.

9.3 Results and Discussion

Using two-dimensional 1H-1H COSY spectra, all 1H chemical shifts could be assigned and interpreted, affording the structure EM-Gly-GlcNAc and EM-Glylactose. According to DLS data, colloidal solutions of GNPs had a bimodal particle size distribution with two diffusion modes, and two hydrodynamic radii, which may be attributed to the motion of unimers with Rh(1), and aggregates with Rh (2). The presence of two types of particles may be interpreted as follows: small particles were unimer micelles in which one metal nanoparticle played a role of a core, while the corona consisted of the glycopolymer stabilizer molecule (core corona structure), and large particles represented clusters of the core corona micelles. For EM-Gly-GlcNAc/Au0 hydrodynamic radii of colloidal particles were: Rh(1) 10.0 6 0.5 nm, and Rh(2) 176.0 6 2.0 nm; for VM-GlcNAc/Ag0 Rh(1) 6.3 6 0.2 nm, and Rh(2) 69.0 6 2.0 nm; for VM-GlyGlcNAc/Au0 Rh(1) 6.7 6 0.5 nm and Rh(2) 54.0 6 3.0 nm. Figs. 9.5 9.8 demonstrate IR spectrum changes of the polymer-stabilizer when going from EM/Au0 (Fig. 9.5) to the thermally activated polymer shell in EM/Au0 (Fig. 9.6), and further to the polymer-stabilized gold GNP (GC-2.2/Au0) (Fig. 9.7). The IR-spectrum of the complex EM/Au0 (Fig. 9.5) was characterized by the presence of the stretching vibration of the carbonyl of ionized carboxyl groups EMνC5О 5 1559, 1395 cm,21 which provided coulomb stabilization of NPs. Thermal treatment of the stabilized NPs led to the appearance of vibration bands at 1779 and 1710 cm21 in the spectrum, corresponding to the cyclic structure of succinic anhydride (Fig. 9.6). Characteristic vibrations of the amide bonds at 1642 cm21(amide I), 1569 and 1403 cm21(amide II) were found in the spectrum of the polymeric shell in EM-Gly-GlcNAc/Au0 (Fig. 9.7).

FIGURE 9.5 IR spectrum of EM/Au0.

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FIGURE 9.6 IR spectrum of EM/Au0 after thermal treatment.

FIGURE 9.7 IR spectrum of EM-Gly-GlcNAc/Au0.

The IR-spectrum of the GC-1 complex EM-Gly-GlcNAc/Ag0 (Fig. 9.8) was characterized by the presence of the stretching vibration of the carbonyl of ester groups EMνC5О 5 1733 and 1059 cm21 (νC-О-C), amide bond at 1653 cm21(amide I) appeared because of the presence of the sugar acetamide group. Fig. 9.9 shows, as an example, micrographs of (A) silver nanoparticles of sample EM-Gly-GlcNAc/Ag0, and (B) gold nanoparticles of sample EM-Gly-GlcNAc/Au0

9.3 Results and Discussion

FIGURE 9.8 IR spectrum of EM-GlcNAc/Ag0.

FIGURE 9.9 TEM micrographs of silver particles in GC-2.2 (A), and gold particles in GC-2.2 (B)

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obtained by TEM. The forms of silver and gold NPs were close to spherical in all samples of sols under study. Diameter of silver and gold particles was 2.5 6 0.3 and 4.5 6 0.5 nm, respectively.

9.3.3 SILVER (OR GOLD)-LABELED NEOGLYCOCONJUGATE: LECTIN INTERACTIONS STUDY 9.3.3.1 Development of lectin sensors Use of multivalent neoglyconjugates based on multiple carbohydrate ligands bound to a polymeric matrix can drastically increase their interaction with lectins, thus, making possible even low-affinity ligands binding studies. Pieters (2009) To estimate the activity of the synthesized GNPs, their interaction with two groups of lectins was studied. The first group consisted of three lectins, that is, WGA, STA, and PAA. All three lectins, although attributed to be GlcNAc-specific, in fact have much higher affinity to (GlcNAc)n oligomers(Liener et al., 1986): WGA: GlcNAc3 . GlcNAc2 . GlcNAc STA: GlcNAc4 . GlcNAc3 . GlcNAc2 .. GlcNAc PAA: GlcNAc6 . GlcNAc4 . GlcNAc2 .. GlcNAc

The second group contained two β-D-Gal-specific lectins, AHA and RCA120, and have the following specificity: AHA: Galβ1 2 3GalNAc .. Gal . Galβ1 2 4Glc RCA 120: Galβ1 2 4GlcNAcBGalβ1 2 4Glc . Gal

Thus, GlcNAc is the weakest ligand for all three GlcNAc-specific lectins, and lactose (Gal-β1-4Glc) is the weakest ligand for AHA, and the strongest one for RCA 120. In recent years, a number of high-precision methods to study lectincarbohydrate interactions were developed. These include, for example, computational modeling (Mishra et al., 2014), X-ray crystallographic studies (Palmer and Niwa, 2003), and NMR (Ferna´ndez-Alonso et al., 2012). Other techniques like quartz crystal microbalance-dissipation assay (Dam and Brewer, 2002; Wang et al., 2014c), fluorescent spectroscopy including fluorescence polarization (Bader et al., 2016; Kakehi et al., 2001), piezoelectric sensing, electrochemical impedance spectroscopy, surface plasmon resonance (SPR) (Vegas et al., 2007; Mahon et al., 2013; Pihı´kova´ et al., 2015; Gemeiner et al., 2009; Dai and Dong, 2008; Gao et al., 2008; Tsvetkov et al., 2012; Duverger et al., 2003; Pattnaik, 2005), and others, (Renaudet and Spinelli, 2011) have also been used for this purpose. Most of the technologies require expensive and often sophisticated equipment (e.g., SPR). Less-sensitive, but simpler, semi-quantitative methods for lectincarbohydrate interactions studies are known. Among them, lectin blotting

9.3 Results and Discussion

Table 9.1 Lectin Detection Limits (Dot-Blot Tests) Lectin STA

a

M(kDa)

Detection Limit, g (nmol)

Polymeric GNPs

100

0.001 (0.010) -“0.1 (1.0) -“0.1 (23.0) -“-“-“0.1 (32.0-52.6) -“-“-“1.0 (10.5) 0.1 (0.83)

GC-2.1/Ag0 GC-2.2/Ag0 GC-2.1/Au0 GC-2.2/Au0 GC-2.1/Ag0 GC-2.2/Ag0 GC-2.1/Au0 GC-2.2/Au0 GC-2.1/Ag0 GC-2.2/Ag0 GC-2.1/Au0 GC-2.2/Au0 GC-3/Au0 GC-3/Au0

WGAb

36

PAAc

19 31

AHAd RCA120e

96 120

a

2 subunits. 2 subunits, mixture of 3 isolectins. c mixture of 5 isolectins. d 4 subunits, mixture of 5 isolectins. e 4 subunits. b

(western blot and dot-blot methods) is an effective, simple, fast, and inexpensive technique (Piskarev et al., 2003; Cao et al., 2013). Here, the dot-blot technique was employed for lectin binding studies using colloidal silver or gold GNPs bearing lectin-specific carbohydrate ligands. Blotting was performed using the Bio-Dot device on the PVDF membrane, because our preliminary studies showed more intense staining on PVDF than on a nitrocellulose membrane, and, moreover, nitrocellulose membrane manifested rather strong unspecific staining for silver GNPs. As demonstrated, lectins retained their activity upon physical sorption on hydrophobic surfaces (Mielezarski et al., 2008). Our study aimed to show the specific interaction of the GNPs with lectins and determine the detection limits of the lectins used (Table 9.1). The results of blotting are shown in Fig. 9.10. Upon binding with lectins, aggregation of silver GNPs produced yellow-brown, and gold GNPs (red-violet staining of the PVDF membrane). Thus, aggregation of colored particles of gold or silver GNPs can be detected visually by the naked eye. The spots did not lose intensity for at least 3 months. Interestingly, lectin-staining intensity did not correlate with the amount of lectin applied after achievement of threshold of sensitivity. Since the size of the lectin spot was strictly fixed with special device, at lectin concentration higher than the sensitivity threshold, interaction of GNPs took place only with the outer layer of “heavy” (multilayer) protein cover.

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FIGURE 9.10 (1). Specificity of the silver (gold) GNPs (0.03 mL of gold (silver) (1 mg mL21) in dot-blot tests with lectins (5 μL, 2 mg mL21, TBS, 0.02 M, pH 7.4). (2)Dot-blot tests using silver (gold) GNPs (0.03 mL of gold (silver) (1 mg mL21) for different lectin concentrations (5 μL, 0.0002-4.0000 mg mL21, TBS, 0.02 M, pH 7.4)

To minimize multilayer of protein, dotting of a certain amount of buffer onto the spot immediately afterward was proposed, resulting in much bigger spots(AlDubai et al., 2008) for higher protein concentrations, and these experiments were carried out without using the Bio-Dot device. In this express dot-blot assay lectin detection was possible even with weak binding carbohydrate ligands (Table 9.1). All the lectins used had two or more binding sites, thus, each molecule of the lectin could bind two or more GNPs or their associates. The most intensive binding (sensitivity) of GNPs was shown for STA, which may be attributed to more the favorable microenvironment in the binding center and higher molecular mass of this lectin. Comparison of Ag0(Au0) labeled GNPs based on GC-1, or GC-2 and GC-3, indicated pronounced advantages of the latter two composites, demonstrating only specific activity, and showing the lack of unspecific interaction with AHA and RCA 120, as well as with BSA devoid of carbohydrate binding sites (Fig. 9.10 (1)). For GC-1 based GNPs, unspecific interaction may be attributed to close spacing of the carbohydrate to the polymeric chain. From the other side, D-GlcNAc in GC-2, or D-Gal in GC-3 were β-linked to three-atomic glycine spacer, ensuring more comfortable orientation of the ligand in the binding site of the lectin. Both poly(N-vinylpyrrolidone-alt-maleic acid)- and poly(ethylene-altmaleic acid)-based GNPs had about the same lectin-binding properties. We did not find any noticeable difference when silver- or gold-labeled neoglycoconjugates were used, with exception of STA, which showed higher sensitivity to silver GNPs (Table 9.1, Fig. 9.10 (2)). It was established that mannose-stabilized silver NPs exhibited longer linear dynamic range and faster reaction kinetics for the target lectin, comparing to gold ones [43]. GC-2.1/Ag0(Au0) and GC-2.2/Ag0(Au0) (Table 9.1) had sensitivity to WGA and PAA as low as 0.1 μg. The highest

9.3 Results and Discussion

FIGURE 9.10 (Continued).

detection limit was found for AHA using GC-3/Au0 (Table 9.1) because lactose is a weak ligand for AHA. For example, for semiquantitative determination of antigen (gold colloid), human serum albumin interaction was used for the dot-blot test (Kakehi et al., 2001) and the limit of detection was found to be 0.1 μg cm23

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The detection limit and dynamic range of the resonance light scattering (RLS) assay in microarray format for system Con A/Man-α were 100 pg mL21; 0.5 500 ng mL21, respectively, and the detection limit and dynamic range of the fluorescence assay in microarray format for the same system were 1 and 1 50 μg mL21, respectively. Gao et al. (2008) Both fluorescence assay in microarray format and blotting techniques were found to have approximately the same level of sensitivity, and these detection limit values are similar to the values we collected. RLS assay was more sensitive.

9.3.3.2 UV-visible absorbance spectroscopy Interaction of the lectins with polymer-stabilized gold or silver GNPs was also monitored using UV-visible absorbance measurements because of the unique optical properties of silver and gold NPs and their sensitivity to polymeric shell environment. Appearance of the specific complex upon binding of lectin with gold or silver GNPs caused changes in optical properties of the sols, and can be illustrated by changes of the spectra of the initial metal-containing GNPs (Table 9.2; Figs. 9.11 9.13). In both cases, a red shift of the spectra maximum occurred. Under appropriate concentrations of gold GNPs, absorbance changes could be measured at lectin concentration as low as 5 μg mL21. Concentration of silver GNPs used was lower than gold ones because molar extinctions of silver and gold nanoparticles had approximately an order difference (an extinction coefficient of silver GNPs was 1.1 3 104 dm3 mol21 cm21 and gold one 3.3 3 103 dm3 21 21 mol cm ). As could be expected, based on the lectin dot-blot data, more strict

Table 9.2 Shift of Absorption Maximum of GNPs in the Presence of Lectinsa Polymeric GNPs

Protein (Lectin)

Shift of Absorption Maximum Δλmax, (nm)b

GC-2.1/Au0 -“GC-2.2/Au0 -“GC-2.1/Au0 GC-2.1/Ag0 GC-2.1/Ag0 GC-3/Au0 GC-3/Au0

BSA STA -“WGA PAA STA WGA AHA RCA120

0 11 5 5 3 9 5 5 11

25 C and exposure time 15 min in the systems: 30 μL lectin or BSA (0.1 mg mL21, 50 mM PBS, pH 7.4) and 1000 μL silver GNPs (0.1 mg mL21); 30 μL lectin or BSA(2.0 mg mL21, 50 mM PBS, pH 7.4), and 500 μL gold GNPs. b initial λmax for silver GNPs 401 6 1 nm, for gold GNPs 521 6 1 nm. a

FIGURE 9.11 Absorption spectra of GC-2.1/Au0 (0.5 mL, 1 mg mL21) (1); addition of 5 μL STA (2 mg mL21, 50 mM PBS pH 7.4, 25 C): incubation 1 min (2), 15 min (3), 40 min (4).

FIGURE 9.12 Absorption spectra of GC-2.1/Au0 (0.5 mL, 1 mg mL21) (1); addition of STA (2 mg mL21, 50 mM PBS pH 7.4, 25 C, incubation 15 min): 5 μL STA (1 nM) (2); 15 μL STA (3 nM) (3); 35 μL STA (7 nM) (4)

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CHAPTER 9 Maleic anhydride copolymers

FIGURE 9.13 Absorption spectra of GC-2.1/ Ag0 (1.0 mL, 0.1 mg mL21) (1); addition of STA (2 mg mL21, 50 mM PBS pH 7.4, 25 C, incubation 15 min): 5 μL STA (1 nM) (2), 15 μL STA (3 nM) (3), 30 μL STA (6 nM) (4), 40 μL STA (8 nM) (5).

specificity and maximal shift was revealed for pairs VM-Gly-GlcNAc/Ag0(Au0) STA and EM-Gly-lactose/Au0 RCA 120 (Table 9.2). When using poly(ethylene-alt-maleic acid) as a gold stabilizer in GNPs, for EM-Gly-GlcNAc/Au0 STA pair, the shift was less, compared to VM-GlyGlcNAc/Au0 STA. Most probably, in the case of poly(N-vinyl-pyrrolidone-altmaleic acid) as the glyconanoparticles shell, N-vinyl-pyrrolidone residues may enhance binding through hydrogen bonds formation of GNPs with the protein. For the other lectins, upon interaction with GNPs, shift of the spectra maximum was less pronounced (Table 9.2). A relatively insignificant shift of plasmon resonance maximum for gold GNPs was also observed earlier (Liu et al., 2009; Sanchez-Pomales et al., 2012; Schofield et al., 2006). Aggregative stability of lectin metal-labeled neoglycoconjugate complex and red shift in spectra are apparently under significant influence of durability of stabilizing cover of nanoparticles and the size of metal nanoparticles. In our case, stabilizing cover of the nanoparticles was carbon-chain polymer, in contrast with the most popular low molecular weight stabilizers. Nanoparticles, stabilized by low-molecular weight cover, seemed to be subjected to agglomeration caused by the external influence (presence of massive linker) apparently more effectively than the polymer-stabilized one. What is more, in our case, nanoparticles size was significantly less,

9.3 Results and Discussion

compared to those prepared using citric acid, and so metal NPs were more stable in solution (the smaller nanoparticles size, the larger the active contacting area). For proteins without carbohydrate-binding sites, there were no spectral changes in protein-GNPs system, and the red shift of spectra maximum did not appear (BSA, Table 9.2). Absorbance spectra (Table 9.2) were recorded at 5 minutes after mixing of metal-containing GNPs with the lectins. Longer incubation of the mixtures resulted in a more-pronounced red shift of plasmon resonance maximum (Fig. 9.11). The change of absorbance spectrum of lectin-glycan complex upon an aggregation in time has been recorded (Schofield et al., 2006). Titration of GNPs with the lectins resulted in a pronounced red shift of plasmon resonance maximum (Fig. 9.12) and was characterized by a threshold value of lectin concentration after which changes in spectra were not observed. Aging of the system, in time, may even cause an aggregation process. This threshold ratio lectin GNP was specific for each lectin, and most probably was connected both with lectin structure (number of binding sites, microenvironment of binding sites, molecular mass, etc.), with the nature of polymeric shells in GNP (degree of hydrophobicity of polymer), and association degree of polymeric GNP in solution. Thus, at saturation conditions, lectins having two or more binding sites may cause clustering, association, and, eventually, aggregation in the system. Based on molecular mass of macromolecules, the degree of polymer substitution with carbohydrates, and the specific content of metal nanoparticles in GNPs, ratio of lectin GNP for threshold (saturation) conditions were calculated. For example, each molecule of STA in these conditions bound 19.4 EM-Gly-GlcNAc/ Au0 GNPs, 13.9 VM-Gly-GlcNAc/Au0 GNPs, or 3.9 VM-Gly-GlcNAc/Ag0 GNPs. That is, protein molecule bound, most probably, with associated GNPs, corresponding to the particles size registered in DLS method as those with Rh(2) values corresponding to aggregates (slow mode). According to DLS method, these data correlated with higher aggregation of GNPs based on poly(ethylene-altmaleic acid), in contrast to poly(N-vinyl-pyrrolidone-alt-maleic acid). This conception also explained a weak optical effect on GNPs lectin binding, as the protein surface interacted with an insignificant part of the associated form of GNPs.

9.3.4 CROSSLINKED LECTIN SORBENTS Crosslinked neoglycoconjugates (CLGCs, Scheme 9.3B) were used as models for elucidation of lectin-carbohydrate binding. These CLGCs may be also used for lectins sorption and isolation. Crosslinking of a mixture of equimolar amounts of two copolymers of maleic anhydride (with different hydrophobic nature) allowed preparation of the sorbents suitable for lectin isolation, with optimal degree of swelling in water and good mechanical properties. Spherical granulated copolymers of maleic anhydride (maleic acid) were easy-to-use, stable under a wide pH range, and easily regenerated and dried. CLGCs bearing glycyl-spacered

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FIGURE 9.14 Optical microscopy image of crosslinked neoglycoconjugate CLGC-1.

carbohydrates, N-Gly-GlcNAc or N-Gly-lactose, were synthesized using the anhydride form of the crosslinked polymer upon reaction with primary amino groups of the carbohydrate synthons. Synthesis was carried out in one stage in aqueous medium without any condensing agents. Sorbents were hydrophilic and easily swelled in water, demonstrating a swelling capacity of 5 mL g21 (pH 7), and 7 mL g21 (pH 8). Fig. 9.14 shows, for example, an optical microscopy image of CLGC bearing glycyl-spacered carbohydrate N-Gly-GlcNAc where the mean diameter of swollen spherical particles was 70 6 20 μm (pH 7). As in the case of colloidal GNPs, crosslinked neoglycoconjugates bearing GlcNAc-containing ligands demonstrated high affinity to WGA, STA, and PAA (Table 9.3). AHA and RCA 120 were not bound to this sorbent. Kd were calculated for the equilibrium conditions in the system lectin CLGC based on the optical density of the supernatant (Table 9.3). Evidently, the lowest Kd corresponded to STA, which correlated to data on colloidal GNPs in dot-blot assay and UV-visible absorbance measurements. STA also bound to N-Gly-GlcNAccontaining sorbent with a high efficacy, where the specific agglutinin content (equilibrium adsorption capacity) was 54 mg g21 sorbent. Lactose-containing crosslinked sorbent did not bind WGA, STA, and PAA, effectively bound RCA 120, although to a lesser degree—AHA (Kd 5 30.0 μM, agglutinin content 7.7 mg g21 sorbent) (Table 9.3). This fact may be attributed to the low specificity of AHA to lactose. Nonspecific binding of the lectins to nonglycosylated spherical granulated copolymers of maleic acid was negligible—within the limits of experimental errors (about 0.5%). Dissociation constants of lectin-neoglycoconjugate complexes for both types of colloidal glycoparticles, using equilibrium data of shift of plasmon resonance

9.4 Conclusions

Table 9.3 Properties of Crosslinked Neoglycoconjgates Properties of Crosslinked Neoglycoconjugates Carbohydrate Ligand: N-Gly-GlcNAc (CLGC-1)

Carbohydrate Ligand: N-Gly-lactose (CLGC-2)

ε0.1%(280nm), Lectin

21 mL (mg
cm)

Kd (μM)/Lectin Sorption (mg g21)

Kd (μM)/Lectin Sorption (mg g21)

WGA AHA STA PAA RCA120

1.65 86.0 1.1 3.0 1.17

20.0/22.1 30.0/7.7 7.3/54.2 17.5/28.4 3.3 /56.8

maximum in UV-VIS absorbance spectra, or ion-exchange fractionation of lectins and lectin-neoglycoconjugates, were also determined. For the system EMGlcNAc/Ag0 STA (GlcNAc linked to polymer via ester bond) Kd 5 0.40 μM; for the system VM-Gly-GlcNAc/Ag0 STA (GlcNAc linked to polymer via glycine spacer) Kd 5 0.64 μM, or 0.41 μM (estimated also basing on equilibrium data of plasmon resonance maximum shift). These constants were lower than those for the system crosslinked neoglycoconjugate-lectin. Kd for neoglycoconjugate-lectin complexes may vary greatly even for the same lectin-carbohydrate system and may also depend on type of carbohydrate immobilization, carbohydrate density, and the detecting system used. For example, for ConA, the variously mannosylated glycoconjugate, Kd values estimated by different methods were 12.88 nM (the binding was monitored with SPR imaging) (Smith et al., 2003), 83 nM (SPR experiments) (Liang et al., 2007), and 200 μM (colorimetric assay) (Chuang et al., 2009). It should be noted that colloidal neoglycoconjugates, gold (or silver)-labeled colloidal GNPs, together with crosslinked lectin sorbents may be isolated in the dry state and had a shelf life of a few years.

9.4 CONCLUSIONS A simple, convenient, and inexpensive technique for the preparation of two types of lectin-specific multivalent neoglycoconjugates—nanometal-labeled colloidal sensors, and cross-linked sorbents—were proposed in this chapter. Maleic anhydride copolymers used allowed efficient anchoring of the appropriate carbohydrates, or carbohydrate derivatives, via their hydroxyl or amino groups. Most specific lectin-neoglyconjugate binding was observed when glycyl-spacered glycosynthons (N-Gly-GlcNAc or N-Gly-lactose) were conjugated to the polymeric matrix. Gold- or silver-labeled colloidal GNPs were generated by reduction of the

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cationic metal precursors in the presence of copolymer-carbohydrate neoglyconjugates. Gold- and silver-labeled GNPs containing glycyl-spacered carbohydrates exhibited specific lectin binding and were, thus, used for rapid detection of lectins in dot-blot assays. Both poly(N-vinylpyrrolidone-alt-maleic acid)- and poly(ethylene-alt-maleic acid)-based Ag0 (Au0) GNPs had about the same binding properties. Negligible distinctions were found when gold- or silver-labeled specific conjugates were used (with the exception of STA). Lectinoblotting is a qualitative method which may be used to detect the lectin-binding glycoproteins in mixtures, for example, to compare the lectin-binding patterns to glycoproteins of different cells, or cells at different stages of differentiation. GNPs thus obtained may be also used as bactericidal agents, or vaccines (Cartmell et al., 2015). Crosslinked spherical granulated neoglycoconjugates bearing glycyl-spacered carbohydrates were also synthesized in one stage in aqueous system without any condensing agents. These crosslinked neoglycoconjugates proved to be efficient lectin sorbents and efficient models for lectin-binding studies. These GNPs and crosslinked sorbents can be used as an effective tool for the elucidation of specific adhesion and recognition processes where lectincarbohydrate binding is involved, and can also be used for characterization and isolation of new lectins. The method proposed for neoglycoconjugate construction is universal and can be employed even without preliminary synthetic steps using commercially available materials such as maleic anhydride copolymers and amino glycosynthons (e.g., aminophenyl glycosides, aminopropyl glycosides, etc.). Thus, maleic acid (anhydride) copolymers can be used for metal NPs stabilization, preparation of GNPs, and, upon crosslinking, as sorbents for various applications. This technique also allows preparation of maleic anhydride copolymers functionalized with other ligands such as carbohydrates (Cai et al., 2017a,b), proteins (including lectins) (Neu and Kuhlicke, 2017; Tawakoli et al., 2017; Coelho et al., 2017), and glycoproteins, peptides and glycopeptides, and compounds of other structural classes.

ACKNOWLEDGMENT This work was financed by a grant from the Russian Foundation for basic research (No1503-09337). We thank Dr. O.V. Vyshivannaya and Dr. I.V. Blagodatskikh for help with DLS investigations and Dr. T.A. Babushkina and Dr T.P. Klimova for help with NMR interpretation.

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FURTHER READING Kaplan, A.M., Morahan, P.S., Regelson, W., 1974. J. Natl. Cancer I 52, 1919 1923.

CHAPTER

Particulate systems of PLA and its copolymers

10

Anjali Jain1, Wahid Khan1 and Agnieszka Kyzioł2 1

Department of Pharmaceutics, National Institute of Pharmaceutical Education & Research (NIPER), Hyderabad, India 2Faculty of Chemistry, Jagiellonian University, Krako´w, Poland

10.1 INTRODUCTION Polylactide (PLA), a polyester derived from lactic acid, has been used widely in biomedical research and applications for the past five decades (Santoro et al., 2016; Jamshidian et al., 2010). Since, its discovery by Carothers in 1932 (DuPont), PLA has been used in different areas of medicine, including tissue engineering (Yang et al., 2005; Santoro et al., 2016), resorbable sutures (Cutright and Hunsuck, 1971), dental materials (Robert and Frank, 1994; Narayanan et al., 2016), ophthalmic implants (Lee et al., 2010), fracture fixation (Narayanan et al., 2016; Pihlajamaki et al., 1992), treatment of bacterial infections (Xiong et al., 2014; Veerapandian and Yun, 2011), and drug delivery (Smith, 1986; Makadia and Siegel, 2011). In drug delivery applications, the controlled release of a drug is always a desirable trait to reduce the dosing frequency (Bechgaard and Nielsen, 1978). Controlled release systems are advantageous to reduce premature degradation, improve drug uptake, sustain drug concentrations within the therapeutic window, and reduce side effects. Various delivery systems, such as depot, micelle, implant, injectable suspension, and particles have been explored to achieve controlled release profile of drugs (Uhrich et al., 1999). From among these systems, particulate systems have witnessed significant technological progress over the years. These systems include particles ranging from nano to micron size and are termed as nanoparticles or microparticles based upon their size range. PLA and poly(lactic-co-glycolic acid) (PLGA) matrices are the most widely used material for the preparation of particulate systems due to their biocompatibility, low levels of immunogenicity, and safety (Mathiowitz et al., 1997). Further, the physiochemical properties of these materials can be easily tailored through the selection of polymer molecular weight, copolymerization, and functionalization (Anderson and Shive, 2012). These polymers have been approved by the US Food and Drug Administration (US FDA) for all food type applications as well as for human use as sutures, bone implants, and screws, and in formulations for sustained drug Materials for Biomedical Engineering: Thermoset and Thermoplastic Polymers. DOI: https://doi.org/10.1016/B978-0-12-816874-5.00010-4 © 2019 Elsevier Inc. All rights reserved.

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delivery as well as vaccine antigens (proteins, peptides, and DNA) (Tyler et al., 2016; Saini et al., 2016). Micro- and nanoparticles offer large volume to surface area ratios, what provides a greater number of reaction sites than macro-size particles with smaller surface areas (Coelho et al., 2010). They have also been found to improve the solubility of unmodified drug compounds (Chrastina et al., 2011). These systems also have modifiable external shells and targeting and stimulus-responsive capabilities (Davis and Shin, 2008). Currently, there are many particulate systems based on PLA and its copolymers in research and clinical use. This chapter will provide an overview on PLA and PLGA-based micro-particulate systems, preparation methods, products in clinical trials as well as clinical use, and advancements and future prospective of research in PLA-based particulate systems.

10.2 PROPERTIES OF POLY(LACTIC ACID) 10.2.1 PRODUCTION OF POLY(LACTIC ACID) PLA is a biopolymer derived from renewable sources, such as corn starch, tapioca roots, sugarcane, etc. In general, it is produced from nonfossil renewable natural resources through the fermentation of polysaccharides or sugar extracted from corn, potato, cane molasses, sugar-beet, etc. (Murariu and Dubois, 2016; CastroAguirre et al., 2016). This allows the biological cycle to come full circle, including PLA biodegradation and the photosynthesis process (Fig. 10.1). PLA is affordable and available for a variety of biomedical applications. Moreover, due to its easy moldability, numerous shapes, including scaffolds, sutures, rods, films, nanoparticles, and micelles, are possible. Since the discovery of PLA, its global mass production from renewable agricultural resources has made PLA one of the most popular green materials. Market demand for PLA, especially in the packaging industry, has grown dramatically over the past decade. According to European trade association for the bioplastics industry, the total production of durable bioplastics, including PLA materials, is forecasted to increase by 535% from 2014 to 2019 (Nagarajan et al., 2016). Certainly, the fast development of various PLA and its copolymer formulations is expected to be made. Even though the market turned to “durable” materials, biodegradable plastics from renewable sources have still remained a high actuality. This is mainly because of potential pharmaceutical and medical applications, such as drug delivery systems, healing products, and surgical implant devices, orthopedic devices, bioresorbable scaffolds for tissue engineering, and many others. Unfortunately, the production of such formulations is still in its early stages. It is obvious, that a high added value enables a particular application of a new product in commercial sectors. Many efforts are currently being made to produce and characterize new grades of biopolymers with improved characteristics, such as stability, long durability, high mechanical and thermal resistance,

10.2 Properties of Poly(Lactic Acid)

FIGURE 10.1 The biological cycle of PLA. Reproduced with permission from Murariu, M., Dubois, P., 2008. The “green” challenge: high-performance PLA (nano) composites. JEC Compos. Mag. 45, 66 69.

easy scalability and processability, flame retardancy, and other tailored physicochemical properties, etc. However, in the particular case of medical applications, more important are such properties as biocompatibility, biodegradation to nontoxic products, high bioactivity, processability to complicated shapes with appropriate porosity, ability to support cell growth and proliferation, and finally appropriate mechanical characteristics. Currently, PLA is industrially obtained mainly through the polymerization of lactic acid or by the ring opening polymerization (ROP) of lactide (the cyclic dimer of lactic acid) (Murariu and Dubois, 2016; Lim et al., 2008). The main methods for PLA synthesis are summarized in Fig. 10.2. In brief, the existence of both a hydroxyl and a carboxyl groups in lactic acid enables its direct conversion into polyester via a polycondensation reaction. Since, the conventional condensation polymerization of lactic acid does not increase its molecular weight sufficiently, high molecular weight poly(lactic acid) is obtained through the ring-opening polymerization of lactide. Also, PLA can be produced by first preparing lactide through a decompression method with the application of a catalyst, the subsequent opening of its ring, and then allowing polymerization. It is noteworthy that direct condensation polymerization has fewer manufacturing steps, lower cost, and is easier to manipulate and commercialize. Thus, this method is more favorable when compared to ring opening polymerization (Murariu and Dubois, 2016; Lee and Hong, 2014). Remarkably, enzymatic polymerization has emerged as one of the most sustainable alternatives as an environmentally benign method that can be carried out under mild

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FIGURE 10.2 Synthesis of poly(lactic acid). Auras, R., Harte, B., Selke, S., 2004. An overview of polylactides as packaging materials. Macromol. Biosci. 4, 835 864 with permission from Wiley.

conditions while providing adequate control of the polymerization process. For instance, Chanfreau et al. reported the enzymatic synthesis of PLA using the enzyme lipase B from Candida antarctica (Novozyme 435) (Chanfreau et al., 2010).

10.2.2 UNIQUE PROPERTIES OF POLY(LACTIC ACID) AND ITS COPOLYMERS PLA has received attention in biopolymer research due to its unique properties, such as biodegradability, biocompatibility, and sustainability (Jain et al., 2016; Basu et al., 2016). PLA is a biodegradable hydrophobic aliphatic polyester appearing in two optical forms: D-lactide and L-lactide. This allows for modulation of its biodegradability and many physicochemical properties by racemization of the D- and L-isomers or the introduction of other hydroxyl acid comonomers. For instance, a semicrystalline polymer of poly-L-lactide (PLLA) is obtained from L-lactide, while an amorphous polymer poly(D,L-lactide) (PDLLA) is obtained from both isomers. PLA is biodegradable by hydrolysis and enzymatic activity, however, the rate of its degradation is strongly dependent on the degree of crystallinity. This can be tailored by grafting with other polymers. In particular, polyethylene glycol (PEG) is the most popular hydrophilic polymer for surface

10.2 Properties of Poly(Lactic Acid)

modification and is commonly used to modify hydrophobic PLA to form the amphiphilic copolymer PLA PEG. Also, for use in surgical implants and tissue repair, a PLA copolymer with poly(glycolic acid) (PGA) poly(lactic-co-glycolic acid) was developed. Now, PLA and PLGA are widely used for various biomedical applications, such as sutures, bone plates, abdominal mesh, and controlled release drug delivery (Lee et al., 2016; Jain et al., 2011). One of the main advantages of using PLA is its flexibility in making microand nanoparticles in various shapes, including spheres, capsules, cubes, and many others. It is noteworthy that physical properties, such as size and shape, as well as chemical properties, including molecular weight and ratio of copolymers, can be easily controlled to obtain desirable pharmacokinetic and biodegradable properties. Unfortunately, the inherent weaknesses of PLA in its raw state, such as brittleness, low toughness, low heat distortion temperature, slow recrystallization rate, and its inadequate crystallization ability and degree after fast processing notably limit the applications of PLA in industry and biomedicine (Tyler et al., 2016; Nagarajan et al., 2016). For instance, a slow crystallization rate is a huge obstacle since a high level of crystallinity is essential in final products as it influences mechanical and thermal properties. However, thermal and mechanical improvement can be assured by stereocomplexation formation, since PLA possesses two diverse isomeric forms (Tan et al., 2016; Farah et al., 2016). Furthermore, the toughness and ductility of PLA can also be enhanced with a wide range of approaches, including plasticization, copolymerization, and melt blending with different tough polymers, rubbers, and thermoplastic elastomers (Nagarajan et al., 2016). Currently, PLA serves as a top alternative to replace petroleum-based conventional polymers (e.g., polypropylene) in many commercial applications. Given the key benefit of having a short life cycle due to its compostable nature, it is possible to use PLA in single use packaging. However, the application areas for PLA are currently widening with its usage in durable products (e.g., screws, boxes, interlocking parts, ball and socket joints, tapping blocks, etc.) (Nagarajan et al., 2016).

10.2.3 BIOCOMPATIBILITY AND SAFETY OF POLY(LACTIC ACID) Biocompatibility is an undoubtedly important feature of any new formulation for medical application. Biocompatibility is not an intrinsic property of a material, but depends on many variables, such as the complex biological environment, specific drug polymer tissue interactions, and degradability. Biodegradable materials, natural or synthetic in origin, can be degraded in vivo, either enzymatically, nonenzymatically, or both. This should result in the production of biocompatible, toxicologically safe byproducts which are further eliminated by the normal metabolic pathways (Makadia and Siegel, 2011; Ramot et al., 2016). The first therapeutic use of PLA was for the healing of mandibular fractures in a dog (Tyler et al., 2016). The first PLGA-based drug delivery system was approved by the US FDA in 1989. It was the Lupron Depot drug delivery system in the form of microparticles made of leuprolide acetate and PLGA (L:G ratio of

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75:25). The system was dedicated for the treatment of advanced prostate cancer and liberates the drug over a period of 4 months after a single injection (Lee et al., 2016). At present, many types of formulations containing PLA have been approved by the FDA for numerous applications making this biopolymer suitable for expedited clinical translatability. Biomaterials based on PLA and its copolymers can now be easily fashioned into sutures, scaffolds, cell carriers, drug delivery systems, and a myriad of other fabrications. Moreover, PLA has been the vehicle and goal of a multitude of preclinical and clinical trials. Current advanced drug delivery systems are mainly based on biodegradable polyesters, such as PLA and its copolymer PLGA. This is particularly due to their biocompatibility, low immunogenicity, lack of toxicity, and the fact that their physicochemical and mechanical properties can be easily regulated by the selection of polymer molecular weight, copolymerization, and functionalization. PLAbased micro- and nanoparticles have been proposed for biomedical applications, such as for controlled release drug delivery systems, mainly for improving oral bioavailability of poorly water-soluble drugs. Unfortunately, they also suffer from a variety of limitations, such as (1) high initial burst release; (2) off-target side effects; and (3) the wastage of drug; (4) and, in the case of nanoparticles, there is a possibility of nonspecific uptake by the reticuloendothelial system (Lee et al., 2016; Jain et al., 2011). Most nanoparticulate formulations based on PLA and PLGA have been focused on drug delivery to target tumor cells (Jain et al., 2011). In this particular case, loading such particles with a poorly soluble drug can significantly increase the drug dissolution rate. Furthermore, such formulations can also enable the administration of highly toxic drugs, which in the case of chemotherapy have to be administrated. Nanoparticles are thought to be absorbed from the gastrointestinal tract after oral administration. Then, they are internalized in cells partly through fluid phase pinocytosis and also through clathrin-mediated endocytosis. It is claimed that they rapidly escape the endolysosomes and enter the cytoplasm within 10 minutes of incubation (Anderson and Shive, 2012). The intestine has a special mechanism to absorb particles of certain sizes; 100 nm particles showed a significantly higher uptake than larger particles (Murakami et al., 2000). It is noteworthy that PLA-based micro- and nanoparticles have shown a certain promise in improving protection from degradation by plasma enzymes, in the case of alternative routes of administration (e.g., nasal, oral, pulmonary, and mucosal) and prolonged gene delivery systems (Saini et al., 2016; Munier et al., 2005; Chen et al., 2013; Ramot et al., 2016).

10.3 MICRO- AND NANOPARTICULATE SYSTEMS OF POLY(LACTIC ACID) Most particulate formulations based on PLA and PLGA have been focused on drug delivery and targeting systems (Danhier et al., 2012). PLA based systems

10.3 Micro- and Nanoparticulate Systems of Poly(Lactic Acid)

are widely used in drug delivery application due to their flexibility and tunable physiochemical properties. Drugs or biomolecules are usually dispersed homogeneously within a PLA matrix (Kamaly et al., 2012).

10.3.1 PREPARATION METHODS OF POLY(LACTIC ACID) MICROAND NANOPARTICLES The formulation of particulate systems with desirable particle size, drug loading, and release profile requires consideration of different factors, such as the nature of drugs to be encapsulated, type of method, selection of solvent, and other components. The selection of preparation methods is done based on the properties of the active substance to achieve high drug loading, encapsulation efficiency, and controlled release rate. Various methods for the preparation of PLA micro- and nanoparticles are described in Sections 10.3.1.1 10.3.1.8 (Makadia and Siegel, 2011; Lassalle and Ferreira, 2007; Leo et al., 2004). All techniques were classified and summarized in Fig. 10.3.

FIGURE 10.3 Schematic classification of PLA micro- and nanoparticle preparation techniques.

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10.3.1.1 Emulsion-based methods Emulsion-based methods involve two major steps, namely the initial formation of an emulsified system, followed by the formation of nanoparticles by the removal of solvents. The selection of emulsion type is largely dependent on the nature of the material to be encapsulated. Lipophilic materials can be encapsulated using an O/W system, while the encapsulation of a hydrophilic substance requires a double emulsification technique (Lee et al., 2016). An emulsified system can be formed by providing sufficient energy to form small droplets. The second step, involving the removal of solvent from the droplet, can be achieved by different methods, such as solvent evaporation, solvent diffusion after dilution, or salting out. In general, the principle of this second step defines the name of the method (Vauthier and Bouchemal, 2009).

Emulsification solvent evaporation This method involves the preparation of a polymer solution in volatile solvents, such as dichloromethane, chloroform, ethyl acetate, etc. Nanoparticles suspension is formed by the evaporation of the polymer solvent, which is allowed to diffuse through the continuous phase of the emulsion (Anton et al., 2008; Allemann et al., 1993). This is a slow process performed under vacuum. Solvent evaporation conditions can be modified by preparing the emulsion with a combination of partially soluble solvents and removing the volatile organic solvent contained in the oil droplets of the dispersed phase by distillation (Quintanar-Guerrero et al., 1999). Mainardes et al. prepared praziquantel loaded PLGA nanoparticles using an emulsion solvent evaporation method and studied the effects of some process variables on the size distribution of nanoparticles. They suggested that preparative variables, such as concentration of stabilizer and polymer, time of sonication, diffusion rate of organic solvent, and ratio of external to internal phases, are important factors for the formation of PLGA nanoparticles (Mainardes and Evangelista, 2005).

Emulsification solvent diffusion This method is also known as the emulsification solvent displacement method. In this method, the solvent selected to dissolve the polymer should be partly soluble in water (Leroux et al., 1995). Initially the solvent is mixed with water and the emulsion is prepared with water saturated with the polymer solvent composing the oil phase and with an oil phase saturated with water as continuous phase. When this primary emulsion is diluted with an excessive amount of water, the additional organic solvent contained in the dispersed droplets diffuses out of the droplets resulting in the precipitation of the polymer. The formation of nanoparticles occurs mainly due to a diffusion mechanism (Quintanar-Guerrero et al., 1997). Kwon et al. prepared estrogen containing PLGA nanoparticles using an emulsification solvent diffusion method. The effects of various preparative variables, such as the type and concentrations of stabilizer, homogenizer speed, and

10.3 Micro- and Nanoparticulate Systems of Poly(Lactic Acid)

polymer concentrations, were studied. The results indicated that the diffusion coefficient of the solvent was proportional to the kelvin temperature of the system and inversely proportional to the viscosity of the continuous phase. Further, the use of DMAB as a stabilizer resulted in a particle size below 100 nm (Kwon et al., 2001).

Emulsification reverse salting out This method differs from the emulsification solvent diffusion method in terms of the nature of the solvent used. The solvent used to dissolve the polymer is completely miscible in water, that is, acetone (Alle´mann et al., 1992). The polymer solution is emulsified in the aqueous phase containing high concentrations of sucrose or salts, such as magnesium chloride, calcium chloride, and magnesium acetate (Ibrahim et al., 1992). These components retain water molecules for their own solubilization; hence, they manipulate the miscibility properties of water with other solvents, such as acetone. When the emulsion is diluted in an excess of water, a sudden drop in the concentration of salt or sucrose in the continuous phase of the emulsion occurs, resulting in the precipitation of the polymer in the form of particles (Vauthier and Bouchemal, 2009). The salting-out process requires optimization of the process conditions, for example, the salt type and concentration, the type of polymer and solvent, and the ratios of these compounds, in order to obtain particles of the desired size range (Wischke and Schwendeman, 2008). This method does not involve heat production and thus may be useful when heat-sensitive drugs have to be encapsulated (Leroux et al., 1995; Quintanar-Guerrero et al., 1997; Kwon et al., 2001; Lambert et al., 2001).

10.3.1.2 Nanoprecipitation method The nanoprecipitation method or solvent displacement method involves the immediate precipitation of polymer in the form of nanoparticles upon the addition of the polymer solution to a nonsolvent (Fessi et al., 1989). The polymer solvent is chosen from among organic solvents that are miscible with a nonsolvent and easy to remove by evaporation. Typical solvents used for nanoprecipitation are acetone, acetonitrile, dimethylacetamide, dimethylformamide, dimethylsulfoxide (DMSO), 2-pyrrolidone, N-methyl-2-pyrrolidone (NMP), PEG, and tetrahydrofuran (Makadia and Siegel, 2011). Acetone is the most preferred solvent (Legrand et al., 2007). Sometimes, it consists of binary blends of solvents, acetone with a small amount of water, or blends of ethanol and acetone (Thioune et al., 1997). This polymer solution is added to the nonsolvent to produce nanoparticles that form instantly upon rapid diffusion of the polymer solution to the nonsolvent. The resulting colloidal suspension contains polymer particles with well-defined size (typically 200 nm in diameter) characterized by a narrow distribution. Nanoprecipitation offers optimum performance when the polymer is dissolved in a theta solvent and the concentration in the polymer dissolved in the solvent remains below the limit between the semidilute and dilute solubilization regime (Legrand et al., 2007). Furthermore, Balati et al. reported that the dielectric

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constant and interaction parameter between the solvent and antisolvent and between the solvent and polymer plays a major role in the determination of particle size of microparticles prepared using the nanoprecipitation method (Bilati et al., 2005). This method is simple, rapid, economic, and reproducible, but it is only limited to the encapsulation of hydrophobic molecules (Jain, 2000).

10.3.1.3 Dialysis The dialysis method is based on the principle of nanoprecipitation with some modification. The polymer is dissolved in an organic solvent and placed in a dialysis tube surrounded by a nonsolvent miscible with an organic solvent. The displacement of the organic solvent with nonsolvent induces precipitation of the polymer as particles with narrow size distribution. The formed particles can be collected after fully displacing the organic solvent with the nonsolvent (Chronopoulou et al., 2009; Hornig and Heinze, 2008).

10.3.1.4 Spray drying Spray drying is a simple one step process requiring the drug containing polymer solution to be spayed as ultra-fine droplets using a spray dryer. The organic phase is instantly evaporated and the dried particles are collected under low pressure with dry air flow (Liu et al., 2015; Bodmeier and Chen, 1988). The method offers the advantages of unit operation and scalability, but the efficient control of drug distribution within the particles is the main problem associated with its use (Makadia and Siegel, 2011; Sosnik and Seremeta, 2015).

10.3.1.5 In situ method for particle formation This method involves the formation of a drug polymer solution in a water miscible solvent and administration of this solution at a target site via injection. Watermiscible solvents dissipate in vivo with the formation of either microparticles or a depot-based system. Solvents such as NMP and DMSO are commonly used for this purpose. The toxicity of solvents must be examined before selection as microparticles formation involves the migration of the solvent in vivo. The solvent removal process may also be responsible for a high burst release (Jain et al., 2000). This method is simple and cost effective in comparison to other methods (Luan and Bodmeier, 2006). This method was used to prepare Lupron depot for the delivery of leuprolide acetate (Jain et al., 2016).

10.3.1.6 Supercritical fluids technique In supercritical fluid (SCF) technology, particles are formed using SCFs which above their critical point exhibit the unique properties of liquids as well as gases (Byrappa et al., 2008). Two principal processes have been developed for the production of nanoparticles using SCFs, namely rapid expansion of supercritical solution and rapid expansion of supercritical solution into a liquid solvent. Sacchetin et al. prepared PLA particles with/without 17α-methyl-testosterone incorporated using a supercritical antisolvent method. Dichloromethane was used

10.3 Micro- and Nanoparticulate Systems of Poly(Lactic Acid)

as a solvent for the polymer and the hormone and CO2 as the SCF. The influence of the operating pressure, polymer solution concentration, and flow rate on the size and morphology of particles was evaluated. The PLA particles were prepared with mean diameters in the range of 5.4 20.5 μm (Sacchetin et al., 2013). SCF and dense gas technology uses more environmentally friendly solvents and also results in particles with high purity and no residual solvents (Kim et al., 1996). However, the method involves the use of high concentrations of the lowmolecular weight PLAs and it is difficult to control particle size and morphology (Yeo et al., 2001).

10.3.1.7 Particle formation using template/mold This method utilizes a hydrogel-based template for the formation of particles of a defined size. The hydrogel template is prepared using a hard-master template by pouring a warm aqueous hydrogel solution and allowing it to solidify at low temperatures based on the property of the hydrogel. The solidified mold is peeled off and the drug/polymer solution is poured onto the hydrogel mold to allow for its even spreading into empty cavities. The solvent is dried to prepare nanoparticles that are collected by dissolving the hydrogel mold in water (Acharya et al., 2010). Gelatin and polyvinyl alcohol are the most widely used materials for mold formation. PVA molds have many advantages over gelatin gel molds, including stronger mechanical strength, ease of handling for preparation of mold, and storage in a dry chamber before use. The shape of microparticles depends on the PVA mold pattern. A template-based technique offers several advantages, such as monodispersed and predetermined microparticle dimensions, easy scale-up, and reproducibility. Moreover, this method can provide high drug loading for water soluble drugs, which is difficult with other methods. Limitations of this method include large particle size (in the range of microns) (Lu et al., 2014).

10.3.1.8 Microfluidic technique Microfluidic approaches are widely used for the preparation of PLA based microparticles. They offer uniform distribution of drug and precise control of particle size. The initial step involves the preparation of an emulsion system by dissolving the polymer and drug in an organic solvent followed by droplet solidification through solvent evaporation, diffusion, or extraction (Zhao, 2013). Microfluidic devices contain a T- or Y-junction where the polymer solution is filled and injected slowly into a large amount of stabilizing solution. Solvent used to dissolved the polymer starts diffusing into water phase resulting in the formation of particles (Liao and Su, 2010). Wantanabe et al. fabricated PLA microspheres by microfluidic emulsification and subsequent dilution in water. PLA concentration and flow rate of microfluidic device were adjusted to obtain particles of 6 50 μm (Watanabe et al., 2011). Xu et al. prepared bupivacaine loaded PLGA microparticles using microfluidic flow-focusing devices. The particles were nearly monodispersed with particle sizes ranging from 10 to 50 μm. These particles resulted in a lower initial burst and an overall slower release than with other conventional

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methods (Xu et al., 2009). Microfluidic techniques are also used in the preparation of multi core shell microparticles (Zhao, 2013).

10.3.2 CHALLENGES WITH PARTICULATE SYSTEM There are many challenges associated with particulate systems. Formulation development, efficient encapsulation of drug within the particle, removal of complete residual solvents, and control over initial burst release are some unsolved problems. Most of the preparation methods of particulate systems use a highenergy mechanical mixing process that may denature the biological drugs, such as proteins, peptides, and macromolecules (Wu and Jin, 2008). Moreover, the hydrophilic nature of proteins and peptides produces a significant challenge in achieving high entrapment of these molecules (Sah and Sah, 2015; van de Weert et al., 2000; Alimohammadi and Joo, 2014). Defining suitable conditions for in vitro release is also challenging in order to mimic in vivo conditions. Though manageable, the characterization of particulate systems still requires advancements to provide an interface between their physiochemical properties, structure, and in vivo behavior. Furthermore, the sterilization of particulate systems is also difficult and requires standardized protocol to be established (Mitragotri et al., 2014). At a large scale, scalability and batch to batch uniformity is very important. However, few methods can produce particles that meet these requirements. Further, the costly setup to produce a particulate system increases the cost of these products (Lee et al., 2016). In terms of biological behavior of particulate systems, little information is available. Their mechanism to reach the target site and mode of action is still not clear (Kamaly et al., 2012). Also there are many unknown facts about their in vivo behavior and the regulatory requirements for particulate systems are not completely established creating a problem in decision making during the approval of the products (Vasir et al., 2005; Nampoothiri et al., 2010).

10.4 PRODUCTS UNDER PRECLINICAL AND CLINICAL TRIAL Novel technologies for research and production of controlled delivery systems have revolutionized the field of drug discovery, providing with many benefits. First of all, controlled release formulations enable the achievement of a maximum therapeutic effect by increasing the efficiency and duration of drug activity with the simultaneous minimization of side effects. Secondly, such formulations increase patient compliance through decreased dosing frequency, convenient routes of administration, and reduction of the unwanted adverse effects by improving site specific delivery. Other important advantages of drug delivery systems are their high bioavailability, rapid kinetic of absorption, as well as avoidance of the hepatic first pass effect (Wong and Choi, 2016; Kamaly et al., 2016).

10.4 Products Under Preclinical and Clinical Trial

Various types of PLA-based formulations have been proposed for applications in many biomedical fields, such as orthopedics, cardiology, dentistry, general and plastic surgery, gynecology, radiology, oncology, transplantology, and many others. In particular, poly(lactic acid) and its copolymers have been extensively studied as controlled release drug delivery systems from more than two decades. PLA and PLGA have been recommended for the development of devices for controlled delivery of small therapeutics (e.g., anticancer, anti-HIV drugs), proteins, vaccines, genes, and other macromolecules. Unfortunately, PLA-based controlled release systems still have many goals to reach and challenges to overcome: (1) low reproducibility between batches, (2) particle heterogeneity in shape and size, (3) low drug loading capacity, (4) low encapsulation efficiency, (5) difficulty in terminal sterilization, (6) problems with controlled release kinetics (high initial burst release with incomplete release of drug), (7) poor scalability of manufacturing process from laboratory to large scale production (Lee et al., 2016; Tyler et al., 2016). Nevertheless, many PLA-based formulations are currently under preclinical or clinical trials and many PLA-based products are on the market (Anselmo and Mitragotri, 2016). Selected examples of PLA-based formulations that have been examined in vitro and in vivo are summarized in Table 10.1. Examples of preclinical trials are divided into groups according to synthesis method. An example of a remarkable proof-of-concept is a novel platform for drug delivery based on tunable PLGA micro- and nanoparticles. In this context, Yeredla et al. proposed biocompatible and biodegradable particles that are capable of incorporating hydrophobic, large biomolecules, and hydrophilic drugs under mild conditions. This is a feat that is problematic to achieve using synthetic protocols for conventional PLGA particle formation. It is noteworthy that the particles exhibit temperature-responsive drug release due to the use of a temperature responsive polymer as one of the aqueous phase forming polymers (a Pluronic F127/dextran aqueous two-phase system assisting self-assembly) (Yeredla et al., 2016). This platform opens new designs and concepts for forming polymeric/liposome nano- and microparticles with novel physicochemical and stimuli responsive properties that could be applied in drug delivery and other relevant biomedical applications, such as theranostics (Fig. 10.4). Another example that is worth mentioning here because of the significant and outstanding research idea is the work of Muntimadugu et al. The group prepared salinomycin (SLM) and paclitaxel (PTX) loaded PLGA nanoparticles using an emulsion solvent diffusion method. These nanoparticles (particle size below 150 nm) resulted in a biphasic sustained release pattern of SLM and PTX for more than a month. The combination of these two drugs was used for the complete eradication of both cancer cells and cancer stem cells, where the SLM targeted the cancer stem cells and PTX was used to kill the cancer cells. SLM nanoparticles were coated with hyaluronic acid (HA) for targeting CD44 receptors overexpressed on cancer stem cells. This combination therapy resulted in a

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Table 10.1 PLA-Based Formulations Under Preclinical Trials (In Vitro and In Vivo Studies) SN

Polymer

Drug

Indication

Outcomes

Reference

Nanoparticles exhibited slow and sustained release. Nanoparticles showed IC50 value twice higher than free drugs. Nanoparticles exhibited efficient co-delivery of both drugs and were effective in killing cancer stem cells and cancer cells. Novel active films based on hydroxypropylmethylcellulose containing nanoparticles loaded with green tea extract exhibited antioxidant capacity, extending the shelf-life of food products with high fat contents. Surface functionalized nanoparticles delivering nitric oxide exhibited no cytotoxicity toward mouse fibroblasts, while causing a 30% reduction of Escherichia coli culture growth. Moreover, the combined treatment with tetracycline resulted in a 90% increase in antibiotic effectiveness.

Dalmolin et al. (2016) Altmeyer et al. (2016) Muntimadugu et al. (2016)

Chitosan based films with incorporated nanoparticles exhibited a classic biphasic sustained release of macromolecules across the buccal mucosa. Nanoparticles are examples of site-specific delivery of drug whose site of pharmacological activity is cell nucleus. In vivo tests proved higher survival rate for drug loaded NPs compared to the free cisplatin group.

Giovino et al. (2012)

Single Emulsion Method 1.

PLA

Vanillin

Antioxidant activity

2.

PLA

Tamoxifen

Cancer treatment

3.

PLGA

Salinomycin and paclitaxel

Cancer treatment

4.

PLA

Green tea extract

Food packing

5.

PLGA PV

S-Nitrosocysteamine

Antibacterial treatment

Wrona et al. (2017)

Reger et al. (2017)

Double Emulsion Method 1.

PEG-b-PLA

Insulin

Buccal delivery of macromolecules

2.

PLGA-mPEG

Cisplatin

Cancer treatment

Gryparis et al. (2007), Mattheolabakis et al. (2009)

3.

PLA

Lamivudine

HIV treatment

Hybrid nanoparticles (PLA and chitosan) with anti-HIV drug were found to be nontoxic toward mouse fibroblast cells (L929). The use of both biopolymers enabled drug protection in the acidic environment in the stomach, while providing sustained release in the neutral pH of the intestinal tract. H2O2-responsive nanoparticles for dual controlled release of platinum anticancer drugs and O2 was developed. Synergistic cytotoxic effect toward cisplatin resistant cancer cells was achieved due to incorporation of nanocarrier with catalase. Such a novel hybrid system combines the advantages of chemotherapy and oxygen therapy.

Dev et al. (2010)

4.

PLGA

Cisplatin

Cancer treatment

Savoxepine

Cancer treatment

Nanoparticles exhibited controlled drug release for up to 1 week.

Leroux et al. (1996)

Nanoparticles exhibited a continuous release of the entrapped drug for 10 days and strong antitumor effect toward human carcinoma. Nanoparticles exhibited significant therapeutic effect with reduced side effects. Nanoparticles exhibited efficient activity toward cancer cells in terms of its sustained release kinetics revealing a novel vehicle for the treatment of cancer. Development of a novel method of nanoencapsulation of hydrophilic drugs, that are difficult to incorporate into a hydrophobic polymeric matrix. Antibacterial activity of antibiotic-loaded nanoparticles was proven against Gram-( ) intracellular microorganism Salmonella typhi and were significantly more effective than antibiotics alone. The formulation was proposed for oral administration in intracellular chemotherapy.

Jeevitha and Amarnath (2013)

Chen et al. (2014)

Salting Out 1.

PLA PEG PLA

Nanoprecipitation 1.

PLA

Anthraquinone

Cancer treatment

2.

PLA

Tamoxifen

Cancer treatment

3.

PLA

Quercetin

Cancer treatment

4.

PLGA

Protamine sulfate, diclofenac sodium, N6-cyclopentyladenosine

Treatment of various diseases

5.

PLGA

Azithromycin

Antimicrobial treatment

Sanjeev et al. (2015) Pandey et al. (2015)

Dalpiaz et al. (2016)

Mohammadi et al. (2010)

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CHAPTER 10 Particulate systems of PLA and its copolymers

FIGURE 10.4 Three-color fluorescent micrographs of microparticles composed of PLGA, Pluronic F127, and dextran. The fluorescence is from TRITC-dextran (red, A), FITC-Pluronic (green, B), and Cy5-PLGA (cyan, C). The particles exhibit a core shell structure: dextran can be seen in the outer shell, whereas the Pluronic F127 and PLGA polymers are located in the core (D, E). Scale bar 10 μm. Adapted with permission from Yeredla, N., et al., 2016. Aqueous two phase system assisted self-assembled PLGA microparticles. Sci. Rep. Nat. 6, 1 8 under CC License-Open access.

synergistic cytotoxic effect on MCF-7 cell lines. Cellular uptake was better in the case of fluorescein isothiocyanate (FITC)-loaded nanoparticles in comparison to FITC solution and further improved 1.5 times in the case of HA coated nanoparticles. Furthermore, the study, performed over MDA-MB-231 cell lines, resulted in highest cytotoxicity of combination of HA coated SLM nanoparticles and PTX nanoparticles against CD441 cells (Muntimadugu et al., 2016) (Fig. 10.5). The early development of PLA/PLGA delivery systems were concerned with small molecules. At the time, the delivery of luteinizing hormone-releasing hormone (LHRH) analogs were intensively investigated by pharmaceutical companies. The first successful product Decapeptyl was introduced to the European market in 1986 by the Debiopharm group. This first degradable injectable microparticle depot system that obtained regulatory approval was dedicated to the treatment of prostate cancer (it is still on the market) (Jain et al., 2016). Selected current clinical trials involving marketed products based on PLA and its copolymer particles are summarized in Table 10.2.

FIGURE 10.5 (A) Scheme of preparation method of SLM-HA NPs (DMAB introduces positive surface charge to the nanoparticles in the primary emulsion, while HA in external phase partially neutralizes the positive charge present on surface). (B) Release kinetics of PLGA NPs loaded with SLM and PTX (Intaxel, the marketed product of PTX, was used as a reference). (C) Fluorescent images of MCF-7 cells treated with FITC, FITC-NPs, and FITC-HA NPs (cell nuclei stained with DAPI, magnification 60 3 at 4 3 zoom). (D) Proposed mode of action of SLM and PTX loaded PLGA NPs designed for treatment of cancer cells and cancer stem cells: (1) passive targeting through EPR effect (NPs ,200 nm); (2) active targeting (HA coated SLM NPs specifically bind to CD44 receptors of cancer stem cells and kill cells by inhibiting Wnt pathway); (3) blockade of MDR-1 proteins overexpressed on tumor cells caused by SLM that block P-gp efflux pump. This results in its inhibition due to increased accumulation of PTX inside the cancer cells; (4) death of both cancer stem cells caused by SLM NPs and cancer cells by PTX NPs. Adapted from Muntimadugu, E., et al., 2016. CD44 targeted chemotherapy for co-eradication of breast cancer stem cells and cancer cells using polymeric nanoparticles of salinomycin and paclitaxel. Colloids Surf. B Biointerfaces 143, 532 546 with permission from Elsevier.

Table 10.2 Poly(Lactic Acid) and its Copolymer Particles Under Clinical Trials SN

Title

Condition

Remarks

NCT No.

Phase

Status

1.

A phase 1/2 dose escalation study of locally-administered OncoGel in subjects with recurrent glioma Efficacy and safety of OncoGel added to chemotherapy and radiation before surgery in subjects with esophageal cancer A study of BIND-014 in patients with urothelial carcinoma, cholangiocarcinoma, cervical cancer, and squamous cell carcinoma of the head and neck (iNSITE2)

Glioblastoma multiforme, brain neoplasms

PLGA-PEG-PLGA, paclitaxel, a bioerodible gel, OncoGel PLGA-PEG-PLGA, paclitaxel, a bioerodible gel, OncoGel

NCT00479765

I, II

Terminated (has results)

NCT00573131

II

Terminated (has results)

PEG-PLGA/PLA-PEG, docetaxel, nanoparticles for injectable suspension, BIND-014 PLA-b-mPEG, paclitaxel, micellar solution for injections, Paxceed PLA-b-mPEG, paclitaxel, micellar intravenous infusions, Paxceed PLA-b-PEG, paclitaxel, micelles, Genexol-PM PLA-b-PEG, paclitaxel, micelles, Genexol-PM

NCT02479178

II

Terminated

NCT00055133

II

Completed

NCT00006276

II

Completed

NCT00111904

II

Completed

NCT02064829

II

Completed

PLA-b-PEG, paclitaxel, micelles, Genexol-PM

NCT01770795

II

Completed

PLA-b-PEG, paclitaxel, micelles, Genexol-PM

NCT01426126

II

Completed

2.

3.

Esophageal cancer, adenocarcinoma of the esophagus, squamous cell carcinoma Urothelial carcinoma, cholangiocarcinoma, cervical cancer, squamous cell carcinoma of the head and neck Rheumatoid arthritis

4.

A study using intravenous Paxceed to treat patients with rheumatoid arthritis

5.

Micellar paclitaxel to treat severe psoriasis

Psoriasis

6.

Phase II clinical trial of Genexol-PM in patients with advanced pancreatic cancer Bioequivalence study of IG-001 versus nab-paclitaxel in metastatic or locally recurrent breast cancer (TRIBECA) A phase II trial of Genexol-PM and gemcitabine in patients with advanced nonsmall cell lung cancer Study of Genexol-PM in patients with advanced urothelial cancer previously treated with gemcitabine and platinum

Prostate cancer

7.

8.

9.

Metastatic breast cancer, locally recurrent breast cancer Nonsmall cell lung cancer

Bladder cancer, ureter cancer

10.

11.

12.

13.

14.

15.

16.

17.

A study of BIND-014 (docetaxel nanoparticles for injectable suspension) as second-line therapy for patients with KRAS positive or squamous cell nonsmall cell lung cancer A phase II study to determine the safety and efficacy of BIND-014 (docetaxel nanoparticles for injectable suspension) as second-line therapy to patients with nonsmall cell lung cancer A study of BIND-014 given to patients with advanced or metastatic cancer

A phase II study to determine the safety and efficacy of BIND-014 (docetaxel nanoparticles for injectable suspension), administered to patients with metastatic castration-resistant prostate cancer A phase II study of weekly Genexol-PM in patients with hepatocellular carcinoma after failure of sorafenib A phase II trial of doxorubicin and Genexol-PM in patients with advanced breast cancer Study to evaluate the efficacy and safety of Genexol-PM once a week for gynecologic cancer A clinical trial of paclitaxel loaded polymeric micelle in patients with taxanepretreated recurrent breast cancer

KRAS positive patients with nonsmall cell lung cancer, squamous cell nonsmall cell lung cancer Nonsmall cell lung cancer

Metastatic cancer, cancer solid tumors

CRPC, prostate cancer

Carcinoma, hepatocellular

Metastatic breast cancer

Gynecologic cancer

Recurrent breast cancer

PEG-PLGA/PLA-PEG, docetaxel, nanoparticles for injectable suspension, BIND-014 PEG-PLGA/PLA-PEG, docetaxel, nanoparticles for injectable suspension, BIND-014 PEG-PLGA/PLA-PEG, docetaxel, nanoparticles for injectable suspension, BIND-014 PLGA/PLA-PEG, docetaxel, nanoparticles for injectable suspension, BIND-014 PLA-b-PEG, paclitaxel, micelles, Genexol-PM PLA-b-PEG, paclitaxel, micelles, Genexol-PM, doxorubicin PLA-b-PEG, paclitaxel, micelles, Genexol-PM PLA-b-PEG, paclitaxel, micelles, Genexol-PM

NCT02283320

II

Completed

NCT01792479

II

Completed

NCT01300533

I

Completed

NCT01812746

II

Completed

NCT03008512

II

Recruiting

NCT01784120

II

Recruiting

NCT02739529

I

Recruiting

NCT00912639

IV

Enrolling by invitation

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CHAPTER 10 Particulate systems of PLA and its copolymers

10.5 PRODUCTS UNDER CLINICAL USE Enormous progress has been made with the application of PLA and its copolymers in almost every medical field. Impactful PLA-based formulations are still being investigated, fabricated, and tested clinically. In particular, such formulations are developed that can effectively deliver a variety of therapeutic compounds. From oncology to dentistry, orthopedics and tissue grafting, the mechanical, physicochemical, and biological properties of PLA have been engineered into numerous practical biomedical achievements. Clinically available products with a short description of the applied drug delivery system is given in Table 10.3.

10.6 ADVANCEMENTS 10.6.1 VACCINATION Oral route vaccination avoids the painful injection procedure, hence, there is a great attraction for the development of oral vaccines. However, inefficient antigen uptake caused by enzymolysis and hydrolysis in the gastrointestinal tract is the main challenge associated with this route. Tan et al. developed acid-resistant HP55/PLGA nanoparticles as an oral delivery system to protect Helicobacter pylori recombinant antigen CCF against the complex gastrointestinal environment. The developed HP55/PLGA-CCF NPs of 200 nm particle size, induced high levels of urease-specific antibodies and memory T cell responses in mice, and 43% of the mice were completely protected after a month of H. pylori exposure. This protection was found to be associated with a Th1/Th17-bias immune response. These NPs offered several advantages, such as antigen protection, slowrelease and targeting, and prevention of gastrointestinal infection (Tan et al., 2017). Dhakal et al. prepared a swine influenza virus (SwIV) vaccine for administration through the intranasal route. The inactivated SwIV H1N2 antigens (KAg) encapsulated in PLGA nanoparticles (PLGA-KAg) were prepared with particle sizes in the range of 200 300 nm. Pigs vaccinated twice with PLGA-KAg intranasally showed increased antigen specific lymphocyte proliferation and enhanced frequency of T-helper/memory and cytotoxic T cells (CTLs) in peripheral blood mononuclear cells. PLGA-KAg vaccination in heterologous SwIV H1N1 challenged pigs demonstrated no symptoms of clinical flu in comparison to control groups. Further, a reduction in viral antigenic mass in the lung sections with clearance of infectious challenge virus was observed in most of the PLGA-KAg vaccinated pigs. Immunologically, PLGA-KAg vaccine augmented the frequency of IFN-γ secreting total T cells, T-helper, and CTLs against both H1N2 and H1N1 SwIV. These vaccines reduced the clinical symptoms of the disease and induced cross-protective cell-mediated immune responses (Dhakal et al., 2017).

Table 10.3 Clinically Used Products of PLA and Its Copolymers

SN

Clinical Product (Approval Year)

Active Agent

1.

Sculptra (2004)

2.

Lupron depot (1989)

Leuprolide acetate

3.

Vivitrol (2006)

Naltrexone

4.

Risperdal Consta (1993)

Risperidone

5.

Decapeptyl

Triptorelin

6.

Profact/ Suprefact Trelstar (2000)

Buserelin acetate Triptorelin

7.

Delivery System

Indication

Company

The formulation of Sculptra contains PLA microparticles, nonpyrogenic mannitol, and sodium carboxymethylcellulose. This formulation is available as a lyophilate and can be injected by mixing with sterile water. Lupron Depot contains leuprolide acetate/leuprorelin encapsulated PLGA microspheres. A 1-month depot injection was developed using PLGA with MW of B14,000 kDa and a lactic acid/glycolic acid ratio of 75/25. Vivitrol are PLGA microspheres developed using Medisorb technology which involves encapsulation of naltrexone within a matrix of PLGA with 75:25 lactide to glycolide content. These microspheres of about 0.1 mm in size began to absorb water almost immediately after injection followed by swelling. Risperdal and Consta are PLGA microspheres based on Medisorb technology. PLGA copolymer of MW B90 kDa was used in Risperdal and Consta microspheres. The drug loading and particle size range for these microspheres were B38% w/w and 25 150 μm, respectively. Decapeptyl is a triptorelin containing PLGA (injectable suspension).

Facial lipoatrophy

Sanofi-Aventis

Advanced prostate cancer

Takeda Pharmaceuticals, Abbvie Endocrine Inc. Alkermes

Profact/Suprefact are buserelin acetate containing PLGA microspheres with 75:25 lactide to glycolide content. Trelstar are triptorelin containing PLGA microspheres.

Treatment of alcohol dependence and opioid dependence

Schizophrenia

Janssen Research Foundation

Locally advanced and metastatic prostate cancer Treatment of prostate cancer Palliative treatment of advanced prostate cancer

Ipsen Pharmaceutical Sanofi-Aventis DebioPharmaceutical

(Continued)

Table 10.3 Clinically Used Products of PLA and Its Copolymers Continued

SN

Clinical Product (Approval Year)

Active Agent

8.

Arestin (2001)

Minocycline

9.

Sandostatin LAR (1998)

Octreotide

10.

Somatuline (2007) Suprecur

Lanreotide

11.

Buserelin acetate

12.

Neutropin (1999)

Growth hormone

13.

Bydureon

Exenatide

Delivery System

Indication

Company

Arestin microspheres contain minocycline hydrochloride within a PLGA polymer matrix where the ratio of PLA and PGA within the copolymer played an important role in tailoring the duration of release. Dose of Arestin is variable and depends on the size, shape, and number of pockets being treated. Sandostatin LAR is a long acting dosage form containing octreotide acetate encapsulated within PLGA microspheres. Copolymer has an average MW B52 kDa and 55:45 M ratio of lactide to glycolide. Sandostatin LAR microspheres are of about 50 μm in size. Somatuline Depot consists of lanreotide acetate PLGA microspheres. Suprecur depot is an extended release PLGA microparticle formulation of buserelin acetate for the treatment of endometriosis in women. It is a long-acting formulation of micronized particles of recombinant human growth hormone (rhGH) embedded in biocompatible, biodegradable PLGA microspheres using ProLease technology. The Bydureon microspheres are prepared based on Medisorb microsphere technology (Alkermes Inc.). The polymer used in microsphere preparation is Medisorb 50:50 DL4AP, composed of lactide and glycolide monomers in an M ratio of 50:50. This polymer has a carboxylic end group and an inherent viscosity of approximately 0.4 dL g21. The microspheres contain 5 mg of encapsulated exenatide per 100 mg of microspheres.

Periodontal disease

Ora Pharma

Acromegaly

Novartis

Acromegaly

Ipsen Pharma Biotech Sanofi-Aventis

Endometriosis, infertility Growth hormone deficiency

Alkermes and Genentech

Glycemic control in type II diabetes mellitus

Byetta (Amylin Pharmaceuticals)

10.6 Advancements

10.6.2 SUPER PARAMAGNETIC IRON OXIDE NANOPARTICLES (SPIONS) Magnetic systems based on polymeric materials, such as PLGA, have been of a special interest for their biomedical applications. In recent years, magnetic nanoparticles have been increasingly studied and exploited for their potential applications as magnetic resonance imaging contrast agents, cancer treatment, targeted therapy, delivery vectors, and hyperthermia. Sivakumar et al. designed an aptamer-conjugated nanocomposite as a multimodal material capable of serving as a contrast agent for MR, photoacoustic, and optical imaging along with drug targeting. They encapsulated curcumin and super paramagnetic iron oxide nanoparticles (SPIONs) inside a PLGA nanocapsule that was conjugated with aptamer for cancer cell specific targeting. The hyperthermic ability of these nanocomposites was mediated by SPIONs upon NIR-laser irradiation. In vitro cytotoxicity as well as photothermal ablation were shown by curcumin-loaded nanoparticles of PANC-1 and MIA PaCa-2 cancer cell lines (Sivakumar et al., 2017). Mosafer et al. developed a method for the encapsulation of sized-controlled oleic acid-coated SPIO into PLGA nanospheres via a modified multiple emulsion solvent evaporation method for both anticancer therapy and magnetic resonance imaging. The SPIO NPs and SPIO-PLGA nanospheres showed spherical morphology with narrow size distribution of 10 and 130 nm, respectively. SPIO-PLGA nanospheres with high SPIO loading of 18.0% and magnetic properties of 5.9 emu g21 were obtained. These nanospheres also enhanced the contrast of relaxation time in tumor site (Mosafer et al., 2017b). The same group of researchers further developed SPION doxorubicin (Dox) coloaded PLGA nanoparticles attached with an AS1411 aptamer (Apt) for targeting murine C26 colon carcinoma cells. NPs were prepared using a modified multiple emulsion solvent evaporation method with a mean size of 130 nm. Dox and SPIO of 3.0% and 16% were loaded respectively. A release study on PBS with pH 7.4 indicated negligible burst release. The conjugation of Apt to NPs enhanced the cellular uptake of Dox in C26 cancer cells. Apt-NPs enhanced the cytotoxicity effect of Dox and resulted in significantly higher tumor inhibition with prolonged animal survival in mice bearing C26 colon carcinoma xenografts (Mosafer et al., 2017a).

10.6.3 CELLULAR INTERACTION The interaction of nanoparticles with cells and lipid bilayers is critical in drug/ gene delivery applications and require a firm control over surface properties (Verma and Stellacci, 2010). These surface properties can be modified by attaching some small molecules onto nanoparticles or by patterning the surface using lithographic techniques. Pillai et al. evaluated the effect of the surface stabilization of PLGA NPs on its cellular interactions, pharmacokinetics, and tumor accumulation. 2-Methoxyestradiol (2ME2) loaded PLGA NPs were prepared through a modified emulsion method and surface stabilized using casein or PEG. Both

371

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types of NPs resulted in similar particle cell interactions as well as antiangiogenesis activity. However, in vivo the pharmacokinetics and tumor accumulation of the drug were substantially improved for the PEGylated NPs. Futhermore, reduced protein binding was observed for PEG stabilized PLGA NPs (Pillai et al., 2017). Molecular imprinting is another potential strategy to promote the recognition and active transport of NPs into specific cells and tissues. Gagliardi et al. synthesized a novel molecularly imprinted nanocarrier based on PLGA and acrylic acid that combines biodegradability and molecular recognition properties. A radial polymerization method was used to synthesize narrowly-dispersed nanoparticles in the presence of biotin as a template molecule. The binding capacity of the imprinted nanoparticles toward biotin and biotinylated bovine serum albumin was 20-fold higher than nonimprinted NPs. These NPs also resulted in effective biotin-mediated cell internalization (Gagliardi et al., 2017).

10.6.4 GENE TRANSFECTION AND TISSUE ENGINEERING The transfection of a cocktail of genes into cells has received great attention in the area of stem cell differentiation. To regulate gene delivery into human mesenchymal stem cells (hMSCs), Park et al. employed multicistronic genes coupled with a nonviral gene carrier system. They fabricated three genes, namely SOX5, SOX6, and SOX9 in a single plasmid which was complexed with polyethylenimine. This complex was coated over PLGA NPs and these NPs showed higher uptake in vitro. Chondrogenesis of hMSCs treated with gene complex loaded PLGA nanoparticles was better than that of hMSCs treated with other carriers. This system provided a simple and effective strategy to deliver multiple genes within cells (Park et al., 2017). In tissue engineering applications, PLGA NPs have also been used widely. Lopes et al. developed doxorubicin (DOX)-loaded PLGA NPs conjugated with PEG and an oxalate variant of transferrin (TPDPs - doxorubicin-loaded poly(lactide-co-glycolide) nanoparticles conjugated with polyethylene glycol and transferrin). These NPs were incorporated into three-dimensional (3D) PLGA porous scaffolds. The PLGA scaffolds with TPDPs incorporated have been shown to release drugs for sustained delivery and provided a continuous release of DOX. An MTS assay suggested a threefold decrease in IC50 values for oxalate TPDPs than native TPDPs signifying greater potency (Lopes et al., 2017).

10.6.5 DENTAL ENGINEERING The innervation of teeth is mediated by axons originating from the trigeminal ganglia. It is essential for their function and protection. In transplanted tissues or bioengineered teeth, the immunosuppressive drug, cyclosporine A (CsA), is used to accelerate the innervations, but the side effects of CsA create problems with its use. To avoid the side effects of CsA, Kuchler-Bopp et al. reported the preparation of CsA loaded PLGA nanoparticles, their embedding on polycaprolactone

10.6 Advancements

(PCL)-based scaffolds and their possible use as templates for the innervation of bioengineered teeth. PCL scaffolds are capable of mimicking the extracellular matrix and are approved by the US FDA. These scaffolds were prepared by electrospinning and decorated with CsA-loaded PLGA nanoparticles to allow a local sustained action of CsA. In adult ICR mice, dental reassociations were coimplanted with a trigeminal ganglion on functionalized scaffolds containing PLGA and PLGA/cyclosporine. These designed scaffolds did not alter the teeth development after in vivo implantation as suggested by histological analysis. Results indicated that 88.4% of the regenerated teeth were innervated using the CsA-loaded PLGA scaffold. These active implants proved their potential for use in dental engineering (Kuchler-Bopp et al., 2017).

10.6.6 ACTIVE TARGETING Active targeting, also called ligand-based targeting involves attaching a ligand, antibody, or carrier protein specific to a receptor on a target site. Active targeting reduces undesired off target effects on other organs and reduces toxicity. ElHammadi et al. developed 5-Fluorouracil (5-FU) loaded folic acid (FOL) decorated PEGylated PLGA nanoparticles for targeted delivery of 5-FU to colon and breast cancers. NPs of PLGA, PEG-PLGA, and FOL-PEG-PLGA were prepared by nanoprecipitation and the release profile exhibited an initial burst drug release followed by a sustained 5-FU release. In normal (CCD-18 and MCF-10A) and tumor (HT-29 and MCF-7) human cell lines, NPs showed negligible cytotoxicity and they were found to be hemocompatible. Cytotoxicity studies in folateoverexpressed HT-29 colon cancer cells and MCF-7 breast cancer cells demonstrated a fourfold reduction in the IC50 of 5-FU-loaded FOL-PEG-PLGA NPs in comparison to 5-FU-loaded PLGA NPs (El-Hammadi et al., 2017).

10.6.7 PHEROID SYSTEM The combination of polymeric NPs as a core and lipid vesicles as a shell has emerged as a robust and promising drug delivery strategy. Chelopo et al. developed a novel combined delivery system, where PLGA NPs are entrapped within a Pheroid drug delivery system. The solid NPs were combined with the Pheroid vesicles using premix and postmix methods. The surface properties of the PLGA NPs were altered through the inclusion and exclusion of chitosan and polyethylene glycol, to evaluate their interaction with the Pheroid vesicles. The average particle size and zeta potential of the NP Pheroid system were found to be in the range of 1990 2450 nm and 218 to 230 mV, respectively. The results suggested that a maximum of 2.5% (w/v) NPs can be optimally added to the Pheroid vesicles without compromising the structure and the stability of the NP Pheroid combined system (Chelopo et al., 2017).

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10.7 CONCLUSIONS Versatile use of PLA and its copolymers in the biomedical field is the outcome of their remarkable properties, which allow their use in the design of different systems and incorporation into various modifications as per the need. Particulate systems, including micro- and nanoparticles, are among the most widely explored systems using these polymers. The remarkable properties of PLA and its copolymers made their way from laboratory to the clinic and introduced many products in market for treatment of various life-threatening diseases. Furthermore, ongoing clinical trials indicate the research possibilities and enormous scope of PLAbased particulate systems for biomedical applications.

10.8 FUTURE PERSPECTIVES PLA and its copolymers have made significant progress in the area of drug delivery. Specially, the particulate systems of PLA provided a mean for the controlled release of therapeutic agents in the treatment of a myriad of diseases. These systems are explored widely for active targeting in cancer treatment and currently the focus is centered toward the theranostic use of these biodegradable polymers by incorporating magnetic particles within these particulate systems. Surface modification of PLA based particles was found to improve cell particle interactions and opened the way for further research. Gene delivery, tissue engineering, and vaccination are also areas where PLA and PLGA based particulate systems are getting significant attention. Though, PLA and PLGA based particulate systems are available in clinical use, there are many issues which require a great deal of research. The conventional techniques have several disadvantages, including the relatively high cost of particle production and the potential toxicity of solvents and reagents, like stabilizers, emulsifiers, and other additives used for forming particles. Moreover, these techniques result in the low encapsulation efficiency of hydrophilic drugs. Modifications in the existing methods may provide a higher encapsulation efficiency. Similarly, the encapsulation of biological drugs, such as proteins, peptides, and macromolecules suffers with low drug encapsulation efficiency and lack of reproducibility. These molecules are more sophisticated and a high energy method may alter their structure, resulting in suboptimal activity. Furthermore, advancements in characterization methods of nanoparticles could further help in understanding the biological interaction of these systems in vivo.

REFERENCES Acharya, G., et al., 2010. The hydrogel template method for fabrication of homogeneous nano/microparticles. J. Control. Release 141 (3), 314 319.

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CHAPTER

Polylactide: the polymer revolutionizing the biomedical field

11

Muhammad Imran Asad, Naveed Ahmed, Asim Ur-Rehman and Gul Majid Khan Department of Pharmacy, Quaid-i-Azam University, Islamabad, Pakistan

11.1 INTRODUCTION Biodegradable polymers have been extensively used in the medical and biomedical engineering fields. Due to the presence of flexible ester bonds in aliphatic polyesters, they can degrade into nontoxic materials in different pH solutions which make them encouraging biodegradable candidates for medical applications. In the medical field, polylactones such as polylactic acid (PLA), polycaprolactone (PCL), polyglycolic acid (PGA), and their copolymers are commonly used synthetic polymers due to their excellent biodegradability (Cava et al., 2006). PLA is an absolutely diverse, aliphatic, biodegradable polymer which is fully obtained from renewable resources like sugar beets and corn. It has also been used as sutures in biomedical applications (Lipinsky and Sinclair, 1986). Carothers in 1932 for the first time prepared a low-molecular-weight PLA (Holten, 1971; Lunt, 1998). DuPont’s work was patented in 1954 resulting in the production of high-molecular-weight PLA (Lowe, 1954) which can be synthesized by ring-opening polymerization (ROP) and direct condensation of lactic acid. As direct condensation is an equilibrium process, it is hard to remove the minute quantities of water in the last stages of polymerization which limits this method to achieve the desired molecular weight and diverts the focus toward ROP (Drumright et al., 2000). There are three stereoisomers of lactide: L-lactide, D-lactide, and meso-lactide. The stereochemical makeup of the resulting polymer is determined by the stereochemical composition of the lactide monomer and, as a result, the melting point, rate, and ultimate polymer crystallization extent is affected. PLA synthesized from pure L-lactide, also known as poly(L-lactide), has a glass transition temperature of 60
C and equilibrium melting point of 207
C (Nijenhuis et al., 1991). It has good crimp and crease properties which give PLA excellent oil and grease resistance with low-temperature sealability and as a barrier to aromas/flavors (Kaplan, 1998). Standard PLA grade has a good strength and high modulus, but lacks toughness compared to polystyrene (PS). Toughness can be increased by copolymerization (Spinu et al., 1996), blending (Ljungberg et al., 2005), and orientation (Lee et al., 2007). Materials for Biomedical Engineering: Thermoset and Thermoplastic Polymers. DOI: https://doi.org/10.1016/B978-0-12-816874-5.00011-6 © 2019 Elsevier Inc. All rights reserved.

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PLA was approved by the US Food and Drug Administration (FDA) in 1970 and since then has been extensively used in food packaging, screws and plates, and clips and sutures (Ogaki et al., 2006; Lampe et al., 2009). The FDA has also recognized PLA resins which are commonly used in surgical and drug delivery applications (Doi and Steinbu¨chel, 2002). Due to its biocompatibility and bioresorsability in the human body, the biomedical products of PLA such as micro filtration plates, sutures, screws, fracture fixation devices, and drug delivery devices are being used (Ramakrishna et al., 2001). Applications of PLA in bone plates became possible in vivo for when a plate loses its rigidity to provide a shielding effect against stress and degrades in the body within a few weeks after implantation (So¨derga˚rd and Stolt, 2002). Micro- and nanoparticles (NPs) of PLA have great influence on drug delivery due to their hydrolytic degradation and low toxicity. The material properties and design greatly affect the release rate of the drug and polymer degradation in the case of microparticles and NPs (Serizawa et al., 2006). Hence PLA has become an alternative to traditional and nonbiodegradable drug delivery systems. In spite of these properties, desired ones are achieved by modifying the polymer by making microparticles and NPs, blending it with other polymers, and coating it with barrier materials. PLA and other filler materials are conjugated to form composite matter with unique properties such as biofibers (Huda et al., 2006), glass fibers (Mathew et al., 2005), cellulose, and nanoclays (Paul et al., 2005; Pluta, 2006; Ray and Bousmina, 2005; Sinha Ray et al., 2003; Ray and Okamoto, 2003; Sasatsu et al., 2006). Obviously PLA has some limitations. In specific requirements, degradation rates cannot be achieved with a sufficiently wide range and in tissue engineering it has no sites for cell recognition which is important for tissue compatibility (Garlotta, 2001). Modifications can be extensively made in PLA by the incorporation of organic and inorganic components which contribute to further biomedical applications of the polymer (Wang et al., 2013). Like other polyesters, PLA’s structure also lacks active cell motif sites, this being the main reason for its modification. PLA has been modified to increase its elongation strain, boost hydrophilic properties, developing biodegradation products with an acidic nature, enhance bioactivity, and increase motif sites for cells in the structure. (Manavitehrani et al., 2016).

11.2 POLYLACTIC ACID SYNTHESIS Polylactic acid is prepared by the polymerization of monomer lactide which is obtained from lactic acid (Hartmann, 1998; Auras et al., 2004; Mehta et al., 2005; So¨derga˚rd and Stolt, 2002; Doi and Steinbu¨chel, 2002). PLA synthesis can be done through three main pathways as shown in Fig. 11.1. Simple polymerization of lactic acid yields a low-molecular-weight and brittle polymer which is normally of no use and requires a coupling agent addition for the formation of

11.2 Polylactic Acid Synthesis

O

CH3

O

CH3 O

O

OH

O

OH O

O

CH3

CH3

Low molecular weight polymer Mw 100,000

Azeotropic dehydration condensation

–H2O H OH

C

OH

C

+

C

HO

CH3

H

CH3

C

HO

O L-Lactic

O

acid

D-Lactic

CH3

acid

CH3

O

O O

O OH

O O

Opoly

CH3

O

CH3

Low-molecular-weight polymer Mw = 1000-5000

O

O C H CH3

C

CH3 H

C

C O

Chain coupling agent

O

Lactide

Ring-opening polymerization

CH3

CH3

O

O

O

O

OH

O O

CH3

OH O

CH3

Low-molecular-weight polymer Mw > 100,000

FIGURE 11.1 Synthesis methods for obtaining low molecular weight.

suitable PLA. The second pathway for synthesis is azeotropic dehydrative condensation of lactic acid which can yield high-molecular-weight polymer without the use of chain extenders or special adjuvants (Auras et al., 2004). The third option which is mainly utilized to obtain PLA with high molecular weight is ROP

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of lactide which has been patented by Cargill (United States) in 1992 (Shin and Narayan, 2010; Hiltunen et al., 1996). Finally, lactic acid units can be part of a more complex macromolecular architecture as is the case in copolymers.

11.2.1 PRECURSORS 11.2.1.1 Lactic acid Either by chemical reaction or by fermentation of carbohydrate, lactic acid is obtained which is the precursor in PLA production. Different forms of lactic acid are presented in Fig. 11.2. Lactic acid has a hydroxyl group with an asymmetric carbon atom and two L and D configurations of optically active isomers which are produced by bacterial systems, whereas the L-isomer is only produced and assimilated by mammals in process of metabolism. Carbohydrates fermented by bacteria are the major source of lactic acid production in huge quantities (200 kT/year). The fermentation methods are classified on the basis of the type of bacteria consumed, namely: (1) the hetero-fermentative class of bacteria produce 1.8 mole lactic acid using one mole of hexose along with higher metabolites like CO2, glycerol, acetic acid, mannitol, and ethanol; and (2) the homo-fermentative technique produces huge amounts of lactic acid with minimum quantities of by-products (Auras et al., 2004). More than 90% of lactic acid is yielded by glucose conversion. The bacterial species of Lactobacilli contribute huge lactic acid yield via the process of fermentation. The D-isomer and mixture of L and D are produced by some strains of bacteria while L-isomer is produced by Lactobacilli casei, L. amylophilus, and L. bavaricus (Auras et al., 2004; Mehta et al., 2005). The sources for basic sugars are corn, sugar beet, sucrose, and potato which are used by homo-fermentative bacteria to produce lactic acid. Other products like amino acid, nucleotide, and B vitamins are produced along with carbohydrates. The conditions required for the process of production are oxygen in a low concentration, temperature around 40
C, and an acidic pH close to 6. To produce soluble solutions of lactate Mg (OH)2, Ca(OH)2, CaCO3, NH4OH, and NaOH are separated out and then the solution is filtered

FIGURE 11.2 Chemical structures of L-, meso-, and D-lactide.

11.2 Polylactic Acid Synthesis

from insoluble by-products and biomass. The crude lactic acid is obtained by evaporation, crystallization, and acidification with sulfuric acid. Lactic acid is purified further if it is intended to be used for the food industry. Separation methods such as ion-exchange, nano- and ultra-filtration, and electrodialysis are usedwhen polymerization is to be done.

11.2.1.2 Lactide The formation of L or LL lactide, D or DD lactide, and LD lactide (meso-lactide) are the results of combining two molecules of lactic acid cyclic diamer. A racemic mixture of L and D-lactide is called raclactide. When reduced, pressure is applied to low-molecular-weight PLA which produces L, D mixture, and meso-lactide by process of depolymerization. The temperature, lactic acid feedstock, and nature of the catalyst defines the percentage of different isomers of PLA (Auras et al., 2004; Mehta et al., 2005). A separation technique, for example, vacuum distillation, based on differences in boiling points of L, D, and meso-lactide plays a vital role among stereoisomers to handle the final structure of the PLA.

11.2.2 POLYLACTIC ACID POLYMERIZATION 11.2.2.1 Condensation and coupling of lactic acid Condensation is considered as the least expensive polymerization method, but high-molecular-weight polymers cannot be obtained by this method. The chains length can be increased by using coupling or the agents that promote esterification (Auras et al., 2004; Mehta et al., 2005), but high costs and multistep complex processes are involved. The coupling agents either react with carboxyl end groups or hydroxyl (OH) of the polymer PLA (Auras et al., 2004, Mehta et al., 2005), thus providing telechelic polymers (Hiltunen et al., 1997). It is important to control the end groups of the chains (Hartmann, 1998; Auras et al., 2004). For controlling end groups, chain extending agents are used because these are economical, used in small quantities, and there is no separation process in the melt reaction for these agents. The copolymers with tunable functional groups are vastly expanded. The disadvantage associated with this process is the residual of unreacted chain extending agents and impurities of metallic catalyst in the final polymer. Furthermore, some agents also lack biodegradability properties (Hartmann, 1998). Isocyanates, epoxies, and anhydrides are examples of chain extenders (Schwach, 2004). PLA blends can also be prepared by the same compatibilization products, but toxicity is associated with the use of isocyanates (Auras et al., 2004). Esterification has the advantage that the final polymer product is free from catalyst residue and is highly purified. By-products and high cost are some of disadvantages associated with these steps and by-products or additives must be removed or neutralized.

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11.3 POLYLACTIC ACID MODIFICATION 11.3.1 MODIFICATION BY HIGH ENERGY RADIATIONS AND PEROXIDES Peroxides and radiation carrying high energy and are used for the structural modification of PLA in radical reaction. Structural modification in poly(L-lactide) (PLLA) by branching with peroxides in weight percent 0.1%0.25% and 0.25 cross-linking concentration has been proposed to be the most dominant (Doi and Steinbu¨chel, 2002). PLA modifications in the original structural properties have been based on the melt reaction of peroxide with PLA. Irradiation causes scissions in PLA chains while cross-linking also modifies the main structure depending on the intensity of radiation and cross-linkers used (Stridsberg et al., 2002). Modifications in the bulk properties of PLA can be made with other materials by blending. PLA is naturally derived from dextran, so renewable material can be obtained. PLA-dextran blends improved the cell affinity and hydrophilicity compared to pure PLA significantly (Cai et al., 2002). The desirable properties of PLA can be achieved by fabrication followed by modification induced by methods such as plasma, physical, chemical, and radiation for biomedical purposes (Huang et al., 2007; Khorasani et al., 2008).

11.3.2 GRAFT COPOLYMERIZATION To improve the multiphase systems or blends having interfacial properties, compatibilizers are frequently used as graft copolymers. Irradiation with electrons and X-rays or UV modifies the huge quantity of polymer while chemical modification is done by plasma discharge and grafting induced into the trunk of the polymer. The surface properties of polymers are modified by the plasma-induced grafting technique of merging an organic vapor into the inorganic gases. The grafting of a substance can be performed either on the bulk or the surface of substance by adjusting the intensity and depth of penetration of radiation (Doi and Steinbu¨chel, 2002). Lactide-based polymers are chemically modified by graft copolymerization which is reported for L-lactide/CL and L-lactide homopolymer and copolymer, respectively, with different contents (Doi and Steinbu¨chel, 2002; Fang and Hanna, 1999). Lactic acid chains can be grafted onto the OH groups for the modification of polymers like carbohydrates (e.g., amylose). Gupta and Deshmukh (1983) found that the addition of coupling agents or peroxides to the melt blend of starch or PLA as the compatibilizer in the processing improves the properties of the substrate. PLA-amylose graft was obtained after purification of amylose from water and butanol residue by ROP using tin (II) bis (2-ethylhexanoate) at 100
C for 20 h in toluene.

11.4 Physicochemical Properties of Polylactic Acid

11.4 PHYSICOCHEMICAL PROPERTIES OF POLYLACTIC ACID 11.4.1 RHEOLOGICAL PROPERTIES Different methods of rheological characterization such as rational and capillary rheometers were extensively used to investigate the rheological properties of PLA and is systems (i.e., composites and blends) (Fang and Hanna, 1999; Hamad et al., 2010, 2011b, 2012; Huneault and Li, 2007; Shin and Narayan, 2010). PLA at low shear rates (,10 s1), showed Newtonian behavior whereas nonNewtonian behavior was observed at high shear rates ( . 10 s1) (shear thinning) as is the case for all thermoplastic polymers. The rheological behavior of PLA has been reported by many studies and expressed that PLA follows the power law (Eq. 11.1) in the same pattern as other polymers over a specific range of temperatures and shear rates (Wang et al., 2011; Hamad et al., 2014; Saeidlou et al., 2012; Dorgan et al., 2000; Lehermeier and Dorgan, 2001; Palade et al., 2001). τ 5 Kγn

(11.1)

where τ is the shear stress, K is the consistency index, γ is the shear rate, and n is the non-Newtonian index. The deviation from Newtonian fluid flow behavior is expressed by the n value, that is also known as the index of flow behavior. A higher value of n shows that the flow behavior is less affected by shear rate. In other words, shear rate is less noticeable with changes in viscosity. Additionally, melting of PLA at different shear stress and shear rate follows (Arrhenius equation) (Eq. 11.2) as given below: n 5 AeEA RT

(11.2)

where E represents the flow activation energy, consistency A is associated to the formulation and structure, and R represents the gas constant. The temperature sensitivity of the viscosity was reflected by the flow activation energy. Therefore greater E indicates the behavior of high-temperature sensitive materials. The effects of (L- and D-isomers) composition on the rheological properties of the PLA polymers were stated in the late 1990s (Fang and Hanna, 1999). Two types of PLA polymers were studied: (1) PLA in amorphous form having (L-lactide) 82% and (D-lactide) 18%; and (2) PLA of a semicrystalline nature containing (Llactide) 95% and (D-lactide) 5%. By increasing the L-isomer in the mixture of the L/D-isomer, the shear viscosity and crystallinity of PLA were also increased (Saeidlou et al., 2012). Dorgan et al. explored two different types of PLA, namely linear and branched, to examine the structural properties of PLA on its rheological characteristics. Their findings presented that the branched PLA had higher viscosity in the Newtonian range compared to linear PLA, whereas the branched PLA had lower viscosity in the non-Newtonian range. This was ascribed to the high shear rates lowering the viscosity and resulting in shear thinning behavior of the polymer (Lehermeier and Dorgan, 2001; Palade et al., 2001).

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11.4.2 MECHANICAL PROPERTIES The behavior of material under different effects like shear, pressure, tensile, impact, and loading modes are described by the mechanical properties of that material. PLA has excellent mechanical properties such as tensile Young’s modulus and flexural and tensile strength in comparison with other traditional polymers like PS, polyethylene (PE), and polypropylene (PP). However, polyamide (PA), polyethylene terephthalate (PET), PP, and PE were higher in breaking at elongation and impact strength compared to PLA. Young’s modulus and tensile strength of PET were compared to PLA and found that its applications were restricted at higher levels of stress due to poor toughness which motivated further interest toward PLA toughening over next few years (Liu and Zhang, 2011; Rasal et al., 2010; Liu et al., 2014a,b). Oyama et al. used PE glycidyl methacrylate (EGMA) for toughening PLA and noticed the annealing effects and strength impact of low and high molecular weights of PLA. The strength impact of PLA could be enhanced by the addition of 20% toughening agent (EGMA) and annealing for 2.5 h at 90
C. Fig. 11.3 represents the EGMA phase of distribution was finer when fabricated with PLAH, having high molecular weight PLA compared to that fabricated with PLA of lower molecular weight (PLA-L). PLA-H had higher melt viscosity than PLA-L which created high shear force at the PLA interface and fine distribution at the EGMA phase.

FIGURE 11.3 Impact strength of PLA and EGMA blends before and after annealing treatment at 90
C for 2.5 h. With permission from Oyama, H.T., 2009. Super-tough poly (lactic acid) materials: reactive blending with ethylene copolymer. Polymer, 50, 747751 (Oyama, 2009).

11.4 Physicochemical Properties of Polylactic Acid

Liu et al. (2014b) used polycaprolactone for PLA toughening using epoxy and polybutylene succinate-colactate as compatibilizers. The blends were prepared and scanning electron microscopy (SEM) identified a decent adhesion among the blend components, which improved the toughness and heat resistance of PLA materials. Recently, an unsaturated biobased aliphatic polyester elastomer (UPE) was implied to produce highly tough thermoplastic PLA blends. The results revealed the improved impact toughness and tensile strength of PLA and UPE.

11.4.3 THERMAL PROPERTIES PLA is an amorphous or semicrystalline polymer with a melting temperature (Tm) at 180
C and a glass transition temperature (Tg) of 55
C. The different structural parameters like composition and molecular weight can affect the thermal properties of PLA. The relationship between molecular weight and Tg was defined by Eq. (11.3) (FloryFox equation) as Tg is related to molecular weight of the polymer. Tg  Tg;N  K=M

(11.3)

where Tg is the glass transition temperature at the immeasurable molecular weight, K is a constant related to the free volume, and M is the molecular weight. This relation was proposed by Dorgan et al. on PLA polymers (Dorgan et al., 2005) where thermal properties of PLA polymers were investigated on the basis of composition (L/D ratio) and their molecular weights. PLA polymers with different composition and molecular weights are shown in Fig. 11.4 with their Tg values. By increasing the molecular weight, the Tg increased quickly but reached a constant value. By increasing L-stereo-isomer content, the Tg of the infinite polymer increased. Several studies also reported the crystalline behavior of PLA (Liang et al., 2013; Hughes et al., 2012; Day et al., 2006; Yuryev and Wood-Adams, 2012) which showed that depending on thermal history and stereochemistry, PLA could be either semicrystalline or amorphous. Its crystallinity was commonly determined by measuring the heat of fusion, Hm, differential scanning calorimetry (DSC), and Hc, the heat of crystallization. The crystallinity [C (%)] can be calculated using Eq. (11.4): Cð%Þ 5

ΔHm 2 ΔHc :100 93:1

(11.4)

where the constant, 93.1 J/g, is the Hm for 100% poly-levo-lactic acid (PLLA) or poly-dextro-lactic acid (PDLA) crystalline homopolymers. Pyda et al. (2004) studied how crystallization behavior of PLA is affected by stereochemistry. The DSC results in Fig. 11.4 show that with B8% and B16% of D-stereoisomer, PLA polymer was amorphous after heating isothermally even at 145
C for 15 h. PLA polymer with 1.5 % D-stereoisomer showed endothermally melting peak at 177
C and suggested that 1.5% D-stereoisomer was semicrystalline.

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FIGURE 11.4 PLA curves shown by differential scanning calorimetry (DSC) with different D-stereoisomer contents: B1.5%, B8%, and B16%. With permission from Pyda, M., Bopp, R., Wunderlich, B., 2004. Heat capacity of poly (lactic acid). J. Chem. Thermodyn., 36, 731742.

Pillin et al. (2008) inspected the effects of a thermomechanical action on PLA properties. Thermomechanical cycles were used to determine the mechanical and rheological properties of PLA. After the first cycle of treatment, the results showed that the viscosity of PLA was decreased by a factor of 0.82. The only constant in thermomechanical cycles was the tensile modulus whereas elongation at break and tensile strengths were decreased after treatment with thermomechani˙ cal cycles (Zenkiewicz et al., 2009). The results of the multiextrusion process showed that impact strength and tensile strength of PLA decreased somewhat, whereas the permeability of water vapor and oxygen and melt flow index were increased after the multiextrusion cycles. As extrusion temperature is increased, the viscosity of the extrudate decreased because molecular weights associated with the shear deformation decreased. During the extrusion process the chain extenders were introduced into the polymer to enhance the thermal stability of PLA which also showed an improved complex viscosity of the extruded polymer (Hamad et al., 2011a).

11.4.4 BIODEGRADATION PROPERTIES The Japanese Biodegradable Polymer Society (JBPS) described the biodegradation procedure in which the polymer is decomposed to carbon dioxide (CO2) and water (H2O) by the activity of microorganisms frequently present in the normal environment, and the JBPS named this type of polymer biodegradable polymers. The two well-known types for the biodegradation of products are aerobic and anaerobic.

11.5 Biomedical Applications of Polylactic Acid

When the original substrate is wholly transformed into gaseous products and no residue remains, it means that complete mineralization and complete biodegradation are established. The rate of biodegradation is affected by various factors (i.e., medium including humidity and temperature) and the chemical parameters of PLA such as composition and molecular weight. The degradation studies of PLA for medical applications like sutures, implants, and drug delivery were carried out in humans and animals. In these environments hydrolysis degrades PLA into soluble oligomers which are then metabolized by the cells. Furthermore, the degradation upon disposal is more challenging because PLA is attacked by microorganisms in ambient conditions. The molecular weight of the polymer must be reduced by hydrolysis at an elevated temperature (58
C) before degradation. The hydrolysis is a first-order kinetics reaction as demonstrated by Eq. (11.5) (Kale et al., 2006, 2007): MW 5 aebt

(11.5) 1

where the constants are a and b which are equal to 230 kDa and 0.18 s , respectively, while t is the time calculated in days. Ohkita and Lee (Ohkita and Lee, 2006) investigated the composites of PLA and corn starch (CS) for the application of rapid biodegradation. The composites were treated in soil under real conditions of composting (B30
C and B80% RH) and test results of biodegradation presented a clear figure that in starch-containing composites, the biodegradation rate was increased because of its higher biodegradability.

11.5 BIOMEDICAL APPLICATIONS OF POLYLACTIC ACID 11.5.1 TISSUE ENGINEERING A sketch of various biomedical applications is presented in Fig.11.5 and discussed in detail as under. Tissue engineering is a highly captivating and promising medical treatment technology in the modern era that gives hope to patients for recovery of lost or malfunctioning organs by using their own cells or tissues grown on polymer support for the regeneration of damaged cells. This polymer support is called the scaffold which serves as an adhesive substrate and physical guide for implanted cells to form new organs. PLA is identified as the bioresorbable alternative polymer in the field of biomedical applications. PLA is a fascinating biocompatible and biodegradable polymer (Lopes et al., 2012). The fabricated scaffolds with gradient pore sizes appropriate for tissue engineering applications were developed by using the sintering technique of solventfree microspheres. To enhance the bioactivity of scaffolds, the emulsification method was used to fabricate poly (D, L, lactide) where TiO2 NPs were used in the preparation procedure as both particulate emulsifier and surface modification agent. A bone-like structure formation was confirmed by FTIR, SEM, XRD, and EDX analysis in the simulated body fluid at day 14 after immersion and had the ability to bond the bone tissues. In vitro studies of osteoblasts also showed the

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FIGURE 11.5 Biomedical applications of PLA.

progressive activity of scaffolds and gave some hope that these bioactive scaffolds can be used for bone tissue engineering as suitable biomaterial (Rasoulianboroujeni et al., 2016). Periodontitis is a medical condition associated with periodontal tissue destruction which results in tooth loss and related problems. Atrisorb is used as the synthetic absorbable barrier containing doxycycline guided for tissue generation in periodontitis (Hatem, 2012). Atrisorb is a synthetic barrier accommodating 4% doxycycline and PLA polymer. The bacterial colonization of the membrane is minimized by this barrier and potential of the barrier is checked by the depth of the probe measurement, level of attachment, and regeneration of damaged tissues (Shue et al., 2012; Polson et al., 1995; Coonts et al., 1998; Jain et al., 2016). Frydrych et al. (2015) used the solvent-casting technique followed by freeze drying and curing to develop a biomimetic scaffold blend of PLLA/PGS (polyglycerol sebacate) for adipose tissue engineering. The PLA/PGA scaffolds possess optimal porosity, mechanical strength, and hydrophilicity for soft tissues and better cell penetration compared to neat scaffolds of PLA. Ease of processing, structural integrity, and mechanical strength make PLA a strong candidate for scaffolds in tissue engineering. PLA blends with other natural and synthetic polymers provide fast degradation, better wettability, and compatible biodegradability for the recovery of tissues and provide an opportunity for modifying the tissue growth (Saini et al., 2016). Scaffolds can also be used to strengthen the mechanical support of ligaments. Scaffolds of poly( L-co-D, L-lactic acid) (PLDLA) matching the mechanical strength and fibrous properties of collagen have been employed for anterior

11.5 Biomedical Applications of Polylactic Acid

cruciate ligament (Surrao et al., 2012). Mechanical and chemical properties of PLA have enabled tissue engineering to manufacture both soft and hard tissues. The fibrous scaffolds of PDLA/PLA have higher tensile and compression strengths compared to pure scaffolds of PLA or PDLA (Ren, 2010). Further research was focused on the design of copolymers with PLA to develop new load-bearing biomaterials so that the mechanical properties of PLA could be enhanced (Habraken et al., 2007). Peripheral nerve damage due to severe injury, pathology, or surgical procedures cause malfunctions and deficits. Analogue nerve grafting is currently a standard procedure to repair damaged nerves, but has some drawbacks including limited donor tissues, mismatch between donor and receptor site, and morbidity of the donor site. Scaffolds are used for nerve tissue repair that matches both the mechanical support to nerve cells and release of the growth factor (Gu et al., 2011). In various cardiovascular problems there is need to use tissue engineering strategies for replacement of damaged cardiovascular tissues (Butcher et al., 2011). There is also a need for PLA nanofibrous scaffolds for the blood vessels and heart regeneration (Hasan et al., 2014; Eschenhagen et al., 2012). PLA nanofibrous scaffolds and patches are most commonly used for cardiac muscle repair (Badrossamay et al., 2010). Chitosan was used to modify the surfaces of PLLA microspheres via graft coating and hydrolysis to enhance their biocompatibility to chondrocytes. The fabrication of PLLA microspheres was done using the solvent evaporation method and hydrolyzed with alkaline solution to produce carboxyl group on large number. Chitosan was covalently grafted onto the microspheres by using the water-soluble coupling agent carbodiimide. At neutral pH, the unbounded chitosan was insoluble which yielded a large amount of coated chitosan. Articular chondrocytes culture of a rabbit model was assessed to evaluate the biological performance of chitosan-coated and controlled PLLA microspheres. The chitosan-coated PLLA microspheres maintained the secretion of chondrocytes and showed a strong ability to promote cell proliferation and attachment. Therefore PLA microspheres coated with chitosan can possibly be employed as injectables for chondrogenesis in cartilage tissue engineering (Lao et al., 2008).

11.5.2 DRUG DELIVERY WITH POLYLACTIC ACID PARTICLES Drug delivery systems have moved from conventional to micro and nano fields. Drugs can be more specific to the target site through the development in nanotechnology. The efficacy and specificity of drugs can be enhanced by nano techniques. A research group developed, evaluated, and characterized PLA NPs for antioxidant activity of diphenyl diselenide (PhSe)2 and these NPs were also characterized in terms of physical stability, polydispersity index (PDI), encapsulation efficiency, mean size, polymerdrug interaction, in vitro drug release, and thermal properties. NPs had an encapsulation efficiency of over 90% with mean size of 210 nm, low PDI, and zeta potential of 224 mV. XRD and DSC results revealed that (PhSe)2 was dispersed as amorphous state in PLA matrix.

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Lyophilized NPs maintained physical stability over 3 months and nanoencapsulation did not change the antioxidant activity of (PhSe)2. All the results prove that PLA NPs can be evaluated for antioxidant activity of oxidative damaged animal models as these NPs are potential carriers of (PhSe)2 (Junior et al., 2017). For glaucoma management, effective intraocular pressure (IOP) lowering drugs exist, but the efficacy of the eye drop regimen to patients was limited. This study tested the hypothesis of reducing IOP for 1 month following a singletargeted injection to the supraciliary space of the eye using microneedles for the administration of brimonidine medication. PLA microspheres loaded with brimonidine were formulated which released the drug at a constant rate for 35 days. This experiment was conducted using the eye of white rabbit models from New Zealand and IOP was observed to be reduced for more than 1 month. Histological examination revealed some foreign body reactions with the microspheres but, overall, administrations were well tolerated. This was the first highly targeted delivery of brimonidine-loaded microspheres for glaucoma (Kim et al., 2014). The development and administration of sustain release microspheres to the ciliary body using microneedles was quite successful. Drug release was at zero order with minimal burst effect when using PLA microspheres (Chiang et al., 2016). Another advancement was done by a group where they developed PLA/PCL microspheres in different clodronate (CL) ratios for the treatment of osteoporosis in drug delivery systems by using a modified double emulsion method. PLA/PCL microspheres had a spongy, porous, inner structure with minute perforations on the surface having an encapsulation efficiency of 80% and ceaseless degradation and release of drug. The controlled release of CL from PLA/PCL microspheres was directly affected by the microspheres degradation. SEM results revealed that bigger holes in the microspheres induced channels to the burst release of CL. It was also revealed that the rapid degradation was attributed to burst CL release while slow degradation led to slow release. Therefore CL release could be adjusted via the PLA/PCL ratios and the best ratio for microspheres was PLA 25% and PCL 75% which gave ceaseless CL release within 20 days with slow release after 10 days. The cumulative in vitro release of CL was also enhanced with microspheres degradation, so PLA/PCL could be used for both the fast and slow release of CL by adjusting the ratios of PLA/PCL (Zhou et al., 2015). Drugs with poor water solubility are presented to the drug delivery system by versatile nanocarriers of amphiphilic polymeric materials. Solvent evaporation of the multiple emulsion technique was used in this work for the production of new amphiphilic copolymers-based nanovectors such as α, β-poly(N-2 hydroxyethyl)D-2-aspartamide-polylactic acid (PHEA-PLA) to synthesize and use controlled release of poorly water-soluble active molecules. For this purpose, the amphiphilic derivative of PHEA, a hydrophilic polymer, was prepared from the backbone with PLA hydrophobic grafts. This newly formed PHEA-PLA copolymer was used for the production of tocopherol (Vitamin E) loaded nanoparticles using multiple emulsion solvent evaporation technique aided with ultrasound energy. This PHEA-PLA system having 100% entrapment efficiency can be useful to

11.5 Biomedical Applications of Polylactic Acid

deliver both hydrophilic and hydrophobic drugs. The available data also confirmed that these NPs can be used for the production of semi-long-term drug release systems (Cavallaro et al., 2015). PH sensitive formulation of PLA/PEG conjugated through acid labile linkage of hydrazone with docetaxel (DTX) for controlled drug delivery. These nanoconjugates are useful carriers of anticancer drugs (Hami et al., 2014). To enhance the water solubility of curcumin and for improvement of absorption following oral administration, PLA/PEG 100 mixed micelles were synthesized and evaluated which showed improved bioavailability of curcumin (Duan et al., 2016). By using the multiple emulsions method, novel PEGylated nanocapsules of PLA (PEG-AcPLA) with high loading capability of water-soluble drugs were synthesized without using conventional stabilizers. Normally nanocapsules were obtained by using sodium deoxycholate with a mean diameter of 203 nm and PDI of 20.1. The model gemcitabine (GEM) hydrochloride was used as the hydrophilic candidate. The nanocapsules of GEM-loaded PEG-AcPLA followed the zero-order drug-release kinetics and showed high encapsulation efficiency. Confocal laser scanning microscopy evidenced the improved cell interaction after 6 h incubation and the 3-(4,5-dimethylthiazol-2-Yl)-2,5-diphenyltetrazolium bromide (MTT)-assay showed that GEM-loaded PEG-AcPLA nanocapsules have increased antitumor effect on cancer cell lines compared to the free drug. Xenograft murine models of solid human tumor showed in vivo antitumor activity of nanocapsules which supported the clinical applications of nanodrugs (Cosco et al., 2015). The biodegradable polymeric microspheres of PLA were synthesized as carriers for nimesulide nonsteroidal antiinflammatory drugs. The emulsion solvent evaporation method was used to prepare these particles. Particle sizes for blank and loaded nimesulide PLA microspheres were found to be 42.9 nm and 2.1 μm, respectively. SEM showed that particles had spherical shapes with PLA microspheres having a loading efficacy of 70%. In vitro dissolution studies revealed that PLA microspheres showed a sustained release system of nimesulide. The results concluded that the developed preparation could be used as a potential vehicle of nimesulide for intramuscle sustained release system (Freitas and Marchetti, 2005).

11.5.3 VACCINE DELIVERY Vaccine is difficult to deliver as it degrades at different body environments that’s why various nano formulations are designed to optimized its delivery i.e. PLGA/ PLA nanoparticles gained reputation as efficient vaccine delivery systems. Similar study was conducted with improved stability of PLA nanoparticles in GI (gastro intestinal) environment via copolymerization of PLA and PEG. By using dissimilar block copolymers AB, ABA, and BAB (where A represents PLA and B represents PEG) NPs were formulated that encapsulated HBsAg (hepatitis B surface antigen) and their effectiveness was evaluated via an oral vaccine delivery

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system. The efficiency of copolymeric NPs to retain encapsulated antigen and average particle size even after 2 h of incubation in simulated gastric and intestinal fluids were engraved in the findings of in vitro studies. As a result of fluorescence microscopic studies, the immunized mucosal gut of mice models efficiently uptake the copolymeric NPs. Finally, after oral administration, these mice models were evaluated for their axillary role in producing immune response. Because of stability issues, PLA NPs could not produce an effectual immune response. However, orally administered copolymeric NPs showed efficient levels of cellular immune response (TH1) and humoral immunity with mucosal (sIgA). It was also observed that the NPs containing BAB depicted improved uptake from mucosa, leading to enhanced immune response in comparison to other NP preparations. So, the effectiveness of NPs containing BAB as an optimistic carrier for oral immunization was authenticated. In BAB copolymers, the presence of PEG at both ends of PLA gave a productive cellular uptake and promisingly elicited efficient immune response. Moreover, without the need of a booster dose, copolymeric NPs exhibited cellular, humoral, and mucosal immunity after oral administration. It can be foreseen that this particular delivery of nanoparticulate schemes can be efficiently used in near future experiments on a larger scale to lengthen the period of immunity and as an approach to improve the immune response. However, additional studies are required for establishing clinical potential (Jain et al., 2010). For controlling the mortality rate caused by various infectious diseases worldwide, protection and safety can only be attained by recurrent immunizations with suitable vaccines. To achieve this purpose, biodegradable poly(lactic-co-glycolic acid) (PLGA) microspheres have been preclinically studied and proved to be most promising for various antigen subunits. The formulations of vaccines with final candidates were preclinically optimized in order to prepare a microspherebased formulation of the tetanus vaccine designed for clinical trials. In guinea pigs, the significance of specific materials and methods of processing for neutralization of antibodies were evaluated. Small-sized vaccines (B5 mm), coadjuvant with mingled alum and fabricated from polymers with fast degradation were found to be the most efficacious. Surprisingly, the form of antigen-stabilizing excipient, population of microsphere mixed together, or the method of microencapsulation used, that is, coacervation versus spray-drying, showed very little effect on the immunogenicity. The clinical samples for testing the immunogenicity and safety in humans were prepared. Admixed alum significantly increased the immunogenicity of microencapsulated tetanus toxoid, while the excipients of a proteinaceous nature were technologically important for synthesizing MS from W/O emulsions. On the basis of these investigations and those made earlier, the tetanus vaccines of MS is at an advanced phase in planning of clinical trials for its assessing immunogenicity and safety (Johansen et al., 2000a). For more than 10 years, PLA and PLGA MS have been evaluated for enhancement in immune response and the controlled delivery of antigens. Early developments of these well-established vaccines stem from the biocompatible polymers

11.5 Biomedical Applications of Polylactic Acid

in concert with their distinctive properties to modify the rates of release and bioerosion. Other features of these microspheres such as capability to produce cellular effector (cytotoxic T-cell responses) and additional antibody responses already observed in earlier studies have become more appealing. Parenteral microsphere-based vaccines, development of related issues, as well as the analysis of immunological fundamentals, and data related with Ag delivery by microsphere were also studied (Johansen et al., 2000b). Microspheres prepared from biodegradable polyester have proven to be effective for single-dose vaccines. In this study, the immunogenicity of diphtheria toxoid (Dtxd) microencapsulated in different types of PLA and PLGA MS synthesized by coacervation and spray-drying methods were examined. The effects of immunogenicity, in vitro release, polymer type and excipients (Bovine Serum Albumin/trehalose) co-encapsulated on Dtxd were investigated in guinea pigs. The Dtxd entrapment efficiency was improved when encapsulated alone with BSA as compared to trehalose in coacervation process. Specific and sustained responses from antibodies over 40 weeks were observed in Dtxd-MS with a relatively hydrophilic ratio (50:50) of PLGA. Whereas, MS made with hydrophobic polymers exhibited very low antibody responses following immunization. Interestingly, large (15 6 60 μm) and small (1 6 5 μm) microspheres presented similar primary antibody responses. So the obtainable data confirmed the viability of MS vaccines to induce very strong and long-lasting defensive antibody responses after a single immunization (Johansen et al., 1999).

11.5.4 TUMOR TREATMENT Over the past few years efforts have been made to improve tumor treatment which led to the survival chances of patients by optimized formulations and the development of carrier and targeted therapeutics that are more specific in their action. The surfaces of polymeric NPs are made functional with additional polymeric layers or ligands by simple polydopamine-based surface modification. In this effort, they synthesized of DTX loaded preparations with polydopamine modified NPs by using D-α-tocopherol PEG1000 succinate-polylactic acid (PD-TPGSPLA/NPs) to boost the delivery of DTX via ligand-mediated endocytosis and targeted cancer cells. They conjugated the prepared NPs with galactosamine (GalPD-TPGS-PLA/NPs). In vitro studies showed similar release profiles of TPGSPLA/NPs, PD-TPGS-PLA/NPs, and Gal-PD-TPGS-PLA/NPs; however, flow cytometry and results of confocal laser scanning microscopy displayed that coumarin 6-loaded Gal-PD-TPGS-PLA/NPs had better efficiency of cellular uptake in HepG2 liver cancer cell lines. DTX-loaded Gal-PD-TPGS-PLA/NPs also inhibited growth of HepG2 more than other two preparations. The in vivo results revealed that Gal-PD-TPGS-PLA/NPs precisely targeted the tumor cells and the tumor size of hepatoma-bearing nude mice was reduced. These ligand-mediated Gal-PD-TPGS-PLA/NPs interacted specifically with hepatocellular carcinoma cells and may be efficiently used as a liver cancer-targeting drug delivery system.

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Therefore Gal-PD-TPGS-PLA/NPs for targeting as drug delivery systems could potentially be used in liver cancer and other liver diseases (Zhu et al., 2016). Conventional chemotherapy has not improved the patient survival rate of the world’s most malignant cancers of small cell lung cancer (SCLC) significantly. An alternative treatment to enhance antitumor efficacy may be the targeting treatment. This study dealt with the SCLC therapy through targeted nanodrug delivery systems. The drug DTX and targeting peptide (AHSGMYP named as AP) separated from H446 cells by phage display technique were conjugated and encapsulated into polylactide NPs (Drug nanoparticles (DN)). In vitro and in vivo experiments were performed to investigate the cellular uptake, biodistribution, therapeutic efficacy, and cell cytotoxicity of AP-DN (AHSGMYP, targeting peptide drug nanoparticles). Encapsulation efficacy with sustained release was .94% and the mean size of AP-DN was found to be 260 nm. AP-DN inhibited the liver metastases effectively with enhanced tolerance and showed excellent antitumor efficiency. The results showed that increased cellular uptake and drug accumulation was attributed to the excellent antitumor efficacy of AP-DN. This was reported for the first time in 2016 and offered a novel system of targeted drug delivery as a therapeutic alternative for SCLC treatment via conjugation of peptide and DTX NPs. In future this approach may be utilized as an active targeting drug delivery system for the treatment of SCLC (Zhu et al., 2016). Human breast cancer cells were evaluated for antitumor efficacy of polylactide/quercetin (PLA/Qt) and NPs were synthesized by embedding Qt in PLA NPs. A novel emulsified nanoprecipitation method was used for the preparation of PLA-Qt with an entrapment efficacy of 62% 6 3% and varying diameters of 32 6 9 nm with different polymer concentrations, drugs, emulsifiers, and temperatures. The larger particles (152 6 9 nm) showed more sustained release of Qt from PLA-Qt NPs compared to smaller particles (32 6 8 nm). The sustained release of Qt from the polymer matrix was due to strong interaction and delayed diffusion in PLA-Qt, killing more than 50% breast cancer cells within 2 days at a drug concentration of 100 g/mL and 40% cell destruction within 5 days was indicated by in vitro cytotoxicity studies. The results showed that sustained release kinetics of PLA-Qt NPs has better antitumor efficacy and could be a novel way to treat cancer. The human breast cancer cell lines (MDA-MB231) were assessed for morphological cell density and percentage cell viability through in vitro activity based on the Qt release rate from PLA-Qt NPs. These in vitro antitumor results suggested that PLA-Qt NPs may be used as a novel way for cancer treatment as compared to Qt alone (Pandey et al., 2015). Triple negative breast cancer (TNBC) has no effective clinical therapy without the expression of estrogen and progesterone receptors and epithelial growth factor receptor-2 of humans. In this study, synthetic lethality-base siRNA and molecularly targeted therapy is used by cationic lipid-assisted PEG/PLA NPs loaded with siRNA. NPs with a carrier (NPsiCDK1) induced decreasing cell apoptosis and viability through expression inhibition of RNAi-mediated CDK1 by delivering siRNA targeting cyclin-dependent kinase 1 (CDK 1) only in the c-Myc overexpressed TNBC cells and not in the normal mammary cells of epithelium. Tumor

11.5 Biomedical Applications of Polylactic Acid

growth in mice models having SUM19 and BT549 was suppressed by systemic delivery of NPsiCDK1, while xenografts showed that there was no immune response or systemic toxicity suggested the therapeutic efficacy of NPs in the overexpressed TNBC. This approach of siCDK1 delivery through NPs provoked no immune response or systemic toxicity and suggested a novel way to treat TNBC (Liu et al., 2014d). To investigate the mimicking action of cytokine and its local antitumor efficacy, Zhao et al. (2013) prepared sustained release microspheres of dextran/ PLGA-PLA loaded with rIL-2. These microspheres were prepared in two steps. In the first step, the researchers loaded rIL-2 into the dextran via the aqueousaqueous emulsion stability method with subsequent encapsulation into the PLGA/ PLA microspheres. In vitro sustained release behavior was achieved for 25 days. A single and multiple dose of rIL-2 microsphere solution were injected into the subcutaneous colon of carcinoma BALB/c mice models to compare the longacting effects of formulation on tumors. The local effects of microspheres solution with a single dose were much effective than the multiple doses shown by the experimental results from the animal models. These experimental results concluded that dextran/PLGA-PLA microspheres loaded with rIL-2 promised the best approach to treat local cancer in animals. More studies on sustained release of rIL-2-loaded dextran/PLGA-PLA microspheres should be carried out for the potential applications of this approach to treat local cancers (Zhao et al., 2013). The copolymer of PLA, poly(ethylene glycol)-block-poly(propylene glycol)block-poly(ethylene glycol) was used to prepare camptothecin (CPT) loaded NPs and examined for in vitro release, characterization of particles and pharmacokinetic efficacy. The preparative ratio of PLA/PEG-PPG-PEG/CPT formulation was 35/35/4 w/w/w in an organic solvent dichloromethane (DCM) and evaporated at 18
C with an entrapment of drug content at 16% which was utilized for in vivo studies. The NPs showed high area under curve (AUC) and delivery of the drug to nearby tissues especially to the liver in normal rats followed by intravenous (I/ V) administration. The I/V administration of CPT-loaded NPs to sarcoma 180 (S180) solid tumor-bearing mice showed excellent tumor suppression without prominent weight loss and their effectiveness was better compared to CPT solution alone. The efficacy of CPT was potentially enhanced by the PLA/PEG-PPG-PEG NPs with high drug retention and gradual release of the drug. Different methods with variable conditions were used to evaluate NPs for drug content, encapsulation efficacy, particle size, and in vitro release. The NPs showed better drug content and excellent retention time at ratios 35/35/4 and organic solvent DCM. The antitumor activities of PLA/PEG-PPG-PEG with administration schedules at 2.5 mg CPT eq./kg3 stated that these NPs should be fruitful to improve the efficacy of CPT (Kunii et al., 2007).

11.5.5 IMMUNIZATION WITH POLYLACTIC ACID PARTICLES For controlled delivery of bioactive molecules, PLA is one with great importance. Ove the past two decades it has been an important contribution for the approval

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of various products by the European medicines agency (EMA) and FDA because of its biodegradability and biocompatibility. PLA has been extensively used in delivery systems for particle preparation, for example, PLA vaccine carriers which exhibit a slow and targeted release of bacterial, viral, and tumor antigens initiating an immune response. PLA-based carriers address not only diagnosis, but also prophylaxis and therapeutics in nanotechnology-based products (Sainz et al., 2015). However, despite knowing their active role in nano-based products in immune cell modulation there is no such marketable nano-based vaccine. Indeed, in the European market most nano-based products belong to the medical devices category. For medical devices, in fact, this has been fostered by the available solid regulatory framework, in comparison to guidelines supporting the evaluation of quality, efficacy, and safety of nano-based products. The foremost problem in nano-based systems is their classification according to current pharmaceutical regulations (Peres et al., 2016). Kojima et al. (2015) studied the release and absorption mechanisms of tacrolimus-loaded microspheres of PLGA or PLA synthesized by the o/w emulsion solvent evaporation method and found there was no decrease in the entrapment efficiency and Tg of PLGA microspheres with the tacrolimus addition. This is because entrapment of tacrolimus is affected by the polymer and drug intermolecular interaction. This data indicated that the release mechanism followed erosion not diffusion with high weight loss of the microsphere and polymer. The pharmacokinetics (PK) profile after subcutaneous administration was similar to intramuscular administration, showing that the release and dissolution are the rate-limiting steps. Also in a heart transplantation rat model graft-survival time was prolonged. The release period from the microspheres in vitro could be controlled by altering PLGA/ PLA ratio. The in vitro release and in vivo absorption results were also similar to these findings and showed that erosion of the polymer and dissolution of tacrolimus were prominent factors of the release mechanism. The findings of this work indicated that formulation facilitated PK optimization and the medication adherence of tacrolimus (Kojima et al., 2015). There is an urgent need to study NP-based vaccine carriers for advanced vaccine research to highlight their antigen delivery systems and immunostimulatory action. Nanoparticle based vaccine delivery system also highlights the delivery of antigen to determine immunogenicity. For this purpose PLGA/PLA nanoparticles encapsulated with Omp antigen of Aeromonas hydrophila were injected in fish by intra-peritoneal injection The preparation of antigen-loaded PLA-Omp and PLGA-Omp NPs was done using the double emulsion method. NPs of ,500 nm sizes were successfully endocyted in the body. Although there was low antigen loading in PLA-Omp, in vitro antigen release was slower in PLGA-Omp NPs. There was more significant hemolytic activity and agglutination of bacterial titer in PLA-Omp and PLGA-Omp immunized groups than all other groups. There was also a persistent increased antibody response of FIA-Omp, PLA-Omp, and PLGA-Omp postimmunization. PLA-Omp antibody response of NPs with treated groups of FIA-Omp indicated that both polymers could be replaced for Freund’s

11.5 Biomedical Applications of Polylactic Acid

adjuvant. Thus NP (PLA/PLGA) based delivery systems could be a novel application for fish parenteral immunization as antigen carriers (Rauta and Nayak, 2015). Recombinant viral subunit-based vaccines are safer compared to classic liveattenuated vaccines. Subunit antigens are less immunogenic, but need an adjuvant. New adjuvants are being developed. Various studies have been performed to check the properties of particulate vaccines, but less studies have evaluated the administration routes of adjuvantantigen combinations. To check the immune response, surface antigen of Hepatitis B was united with cationic microsphere or aluminum-based adjuvants. Microparticles (MP)-based vaccine modifies CD40 and CD80 expression and increase dendritic cell activation. If they are given via the I/M route, then IgG2 isotype of high amounts are generated. The secretions of cytokine IL-4, Th1-type cytokines (IL-2, IL-12, and IFN-γ), Granzyme B, and Th2 are increased by MP-based vaccine so these vaccines can give high humoral and cellular response. If given by I/M or S/C routes, the MP-based vaccine stimulated the activation of CD8 T cells and cytokine expression of Th1-type. MPbased vaccines are more operative and effective (Chen et al., 2014). NPs containing chitosan/polylactic acid were formed by nanoprecipitation in the size range of 300 nm. Chitosan was coated to increase the retention capability at the procorneal area. A rabbit model was used to check the immunosuppression of NPs in corneal transplants and treated corneal allografts had a median survival time of 23.7 6 3.20 days compared to 10.6 6 1.26 days for untreated groups, hence immunosuppression was increased by NPs. RAPA-loaded chitosan/PLA NPs can be prepared via the nanoprecipitation method, were labeled with 99mTc, and showed excellent retention ability at the procorneal area. These NPs are used to treat corneal allografts (Yuan et al., 2008). Th1 lymphocytes are activated to produce stronger immune responses and are induced by mixing a HER-2/neu synthetic cytotoxic T lymphocytes (CTL) peptide with PLA microspheres. PLA microspheres with peptide antigens showed potent in vivo immune response with profile T lymphocytes and cytokine while PLA microspheres alone produced a biased response. The mixture of HER-2/neu peptide with PLA microspheres indicated strong Th1 immune response on S/C administration and the incubation time did not affect the immune response (Nikou et al., 2005). Pa´linko´-Biro´ et al. (2001) conducted a study to observe the immunogenicity of inactivated microencapsulated parvovirus in goose and Muscovy duck (Cairina moschata). Inactivated suspension of duck parvovirus was microencapsulated into 1417 kDa PLA and PLGA by coacervation technique. The in vitro discharge of antigen from separate and mixed PLA and PLGA (50:50 H) microspheres was observed as biphasic for approximately 10 days with an initial lag phase monitored by further 12 days’ constant release. The antigen-loaded MS were introduced S/C into the ducks. The uncertain immune response after single dose administration of microsphere was observed below 200 over a period of 6 weeks, except for pre-immunized animals i.e. 3 weeks before inoculation of microsphere. Lower dose of microsphere and unsuitable release kinetics of antigen attributed to

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CHAPTER 11 Polylactide

weak immune reaction. The lower concentration of antigen of the virus suspension attributed to moderate antigen loading. Results indicated that, further advancement is vital to optimize the dosage form as it would be appropriate for the induction of a faster, long-lasting, and more elevated immune response (Pa´linko´-Biro´ et al., 2001).

11.5.6 DNA AND GENE DELIVERY Another promising approach for the treatment of cancer is combinatorial therapy through the codelivery of small anticancer drugs and minicircle DNA (mcDNA) via stimuli-sensitive nanocarriers. However, it is remarkably challenging to simultaneously load DNA and drugs into nanosized delivery systems. Micelles were synthesized based on triple-block copolymers of poly (2-ethyl-2-oxazoline)-poly (L-lactide) for codelivery of doxorubicin (Dox) and supercoiled mcDNA vectors when grafted with bioreducible polyethylenimine (PEOz-PLA-g-PEI-SS). The nonfouling oxazoline offered biological stability; PLA provided the hydrophobic core for encapsulation of the drug, while complexation and stimuli-responsive release of mcDNA was provided by bioreducible PEI-SS. In vitro results showed higher penetration of mcDNA-loaded micelleplexes into tumor models with higher gene expression and specific kinetics in comparison to nonbioreducible nanocarriers. Bioluminescence imaging of in vivo studies also showed the detection of gene expression for about 8 days after intratumoral administration of mcDNA micelles. The effective encapsulation of both mcDNA and Dox with high efficiency was also verified for the codelivery of drugs and genes in PEOz-PLA-gPEI-SS nanocarriers. Micelleplexes with dual loading showed remarkable uptake and cytotoxic effects represented in 2D cultures of cancer cells. The cancer cells viability and the reduction in volume of tumors was noticed in tumor-bearing mice after codelivery of mcDNA-Dox to B16F10-luciferase. Overall, these results indicated that bioreducible triple-block micelles have an effective delivery in vivo and represent possible combinatorial DNA therapy applications in the near future. PPP-SS bioreducible nanocarriers were utilized for the delivery of single genes as well as druggene codelivery. Furthermore, modifications in pH-responsive moieties of micelles and hydrophobic blocks could be beneficial for controlled drug release in the cancer cells compartment (Gaspar et al., 2015). The differences in physiochemical properties had remained a challenge to design the nanocarriers delivery system for both the nucleic acids and drugs. In this study, they manufactured triple amphiphilic micelles block of poly(2-ethyl-2oxazoline)-PLA-g-PEI (PEOz-PLA-g-PEI) for delivery of minicircle (mcDNA) vectors. The PEI moieties in the polymer backbone were replaced by the bioreducible PEOz. The formed micelles were hemo-compatible, stable upon incubation and showed low critical micelle concentration. The uptake of nanocarriers in MF7 cells was higher. Furthermore, mcDNA-loaded micelleplexes promoted higher gene expression and excellent 3D cell penetration. Additionally, this gave concept of codelivery of Dox and mcDNA encapsulated simultaneously in PEOz-PLA-g-

11.5 Biomedical Applications of Polylactic Acid

PEI nanocarriers with higher efficacy. The therapeutic applications of this system can be used for cancer therapy and, with modifications in chemical structures, can be utilized for cell targeting, imaging, and nano diagnostics/theranostics (Gaspar et al., 2014).

11.5.7 ANTIGEN LOADING NPs are the formulations used to deliver various therapeutic agents including drugs, DNA, proteins, antigens, and different molecules. These agents are formulated into NPs for targeted and controlled delivery. Microparticles (MPs) as adjuvants have attracted interest for the delivery of vaccines. The immune response can be regulated by many physiochemical properties of MPs like surface properties, morphology, and hydrodynamic size. As surface charge affects the adjuvanticity of MPs, Liu et al. (2014c) coated polylactide MPs (PLA-MPs) with different polymers having a positive charge and investigated how surface charge affected antigen loading, macrophage phagocytosis, and in vitro activation. Antigen loading and internalization into macrophages, promoted expression of MHC II and secretion level of TNF greatly enhanced by higher surface charge. They coated chitosan (CS), chitosan chloride (CSC), or polyethylenimine (PEI) on the surface of PLA-MPs with different surfaces charges. Then these MPs were loaded with HBsAg and evaluated by in vitro potential adjuvanticity using RAW 264.7 macrophages. The results showed greater adsorption of the antigen with increased surface charge of MPs which, in turn, increased the internalization of the antigen into the macrophages and their elevated activation (Liu et al., 2014c). Streptococcus equi. cause strangles which affects the upper respiratory tract of equidae. Control of this infectious disease and protection of animals has not yet been accomplished. Convalescent horses have a defensive immune response against opsonogenesic S. equi and SeM, an antiphagocytic protein. Liu et al. (2006) investigated the protein extract of purified recombinant S. equi and SeM entrapped in nanospheres of PLA and their potency was calculated and studied via I/M route. Spherical protein/antigen NPs ,500 nm were prepared by the solvent evaporation technique. Vaccination with S. equi and SeM entrapped PLA nanospheres elevated the antibodies and responses from the cellular immune system. The enzymatic extract of S. equi and nanospheres loaded with SeM-PLAGCS caused the induction of high levels of IgG and IgG2a compared to PLAPVA nanospheres. This shows that adjuvanticity of PLA particles elevate the immune response. The results suggested that PLA carriers are suitable candidates for a safe and effective vaccine against strangles as they provide targeted phagocytic cells such as dendritic cells or macrophages. As they are inexpensive and easy to obtain, they can be used for balanced immune response in veterinary vaccine (Liu et al., 2006).

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11.5.8 PROTEIN DELIVERY Oral protein delivery and increased bioavailability are major challenges related to proteins due to high molecular weight, enzymatic degradability, and low membrane permeability. NPs have solved these hurdles by formulating particles having small sizes and higher surface areas with enhanced bioavailability and selectivity. The different polymeric nanocarriers are very appealing for in vitro and in vivo biomedical applications. Hydrophobic cyanines (IR-780 or ZnPc) or curcumin (CUR) loaded poly(D,L-lactide) nanospheres stabilized with cremophor EL were formed via the nanoprecipitation method for cancer tissues. The cremophor nanospheres of EL/PLA/water showed the even shape, low PDI (,0.3), and high entrapment competence of the selected drug (over 90%). Lower levels of in vitro drug release were observed from these nanospheres which suggested their good stability upon prolonged storage. The affinity of obtained nanospheres for the plasma protein albumin was inquired in vitro because the half-life and biodistribution are related to the adsorption of the nanocarriers to the plasma proteins. The binding of cyanine IR-780 nanospheres with albumin showed lower efficiency than cyanines ZnPc or CUR in PS plate wells. This is an important phenomenon to understand the behavior of the polymeric nanocarriers in systemic circulation when these are administered through parenteral route (Pietkiewicz et al., 2016). The protein/peptide drug delivery system (DDS) was found to be an important application of biocompatible and biodegradable PLGA-based and chitosan-based microspheres. Higher constant value for the serum drug concentration can be maintained for an extended period of time with encapsulation of protein/peptide drugs in the microspheres. However, in the preparation of protein/peptide-loaded microspheres, various problems were encountered including large size distribution, protein deactivation, storage, and release. Microspheres with uniform size and controlled diameter from submicron to 100 μm can be prepared by using the membrane emulsification methods including rapid membrane and direct membrane emulsification processes. The deactivation of proteins can be prevented by using mild membrane emulsification conditions. The bioactivity of protein drug was maintained by adding various additives in protein solution and use of powdered solid drug instead of protein solutions and using encapsulation techniques. For examples in PLGA/PLA microsphere/microcapsules, hydrophilic poly (lactide)-poly (ethylene glycol) (PELA) was used as wall material for encapsulation. Similarly, protein drug was absorbed on a hollow porous structure of chitosan which was already crosslinked. Animal test results showed that higher drug concentration in the blood can be maintained up to 2 months with recombinant human growth hormone (rhGH) loaded PELA microspheres linked with microcapsules of PLA and PLGA. In another study use of poly(lactide)-poly(ethylene glycol) PELA alone improved bioactivity more than 90 % by maintaining encapsulated rhGH during 45 days release which is much higher than wall material of PLGA or PLA. To prevent the chemical cross-linking of proteins which were encapsulated in the chitosan microsphere, the stepwise cross-linking technique

11.5 Biomedical Applications of Polylactic Acid

and self-solidification arrangement was developed. Blank chitosan microspheres were prepared with hollow, hollow-porous, and macroporous morphologies as an alternative of conventional solid microspheres. Then the protein/peptide drugs were adsorbed to these blank chitosan microspheres. The highest loading efficiency of protein, lowermost burst effect, and comparatively constant release pattern were showed by the hollow-porous microsphere. The level of blood glucose was reduced greatly when hollow-porous microspheres were used in oral administration as an insulin carrier compared to the solid microspheres (Ma, 2014). The Shirasu emulsification membrane technique for porous glass Shirasu Porous Glass (SPG) and the double emulsion-evaporation method were combined to prepare recombinant human insulin (rhI) loaded PLA/PLGA microcapsules. The inner water phase (w1) was made of an aqueous phase containing rhI and the oily phase was made by dissolving PLA/PLGA and Arlacel 83 in a mixture DCM and toluene. A w1/o primary emulsion was formed by emulsifying the two solutions. The uniform w1/o/w2 droplets were formed by permeating the primary emulsion into an outer water phase through the uniform pores of a SPG membrane in the presence of nitrogen gas. The solvent was evaporated to acquire solid polymer microcapsules. The encapsulation efficiency of the drug was influenced by various factors of the preparation process and the cumulative drug release was explored systemically. The encapsulation efficiency of the drug and its cumulative release were influenced by the ratio of PLA/PLGA, concentration NaCl in outer phase of water, the volume of inner phase of water, loading amount of rhI, value of pH in outer phase of water, and the microcapsules’ size. The preparation process was optimized to increase the encapsulation efficiency of the drug up to 91.82%. The benefit of formulating the membrane emulsification method is that the size can be controlled accurately in drug-loaded microcapsules and microcapsules, and, therefore, high encapsulation efficiency can be obtained (Liu et al., 2006).

11.5.9 IMAGING AND DIAGNOSIS NPs, due to their unique features, are not only used to deliver therapeutic agents and drugs, but are also valuable candidates for imaging and diagnostics purposes. Imaging through NPs to specific areas has led to precise and fast diagnoses of diseases. The SM5-1 conjugates with PLA for specific targeting of bioconjugate to the HCC-LM3-fLuc cells were studied by Ma et al. (2014). SM5-1 is a humanized antibody which has very high affinity toward the overexpressed membrane protein of breast cancer, melanoma, and hepatocellular carcinoma. The PLA conjugated with SM5-1 were prepared and checked for tumor growth inhibition. Tumor inhibition by PLA-5FU-SM5 was 45.07 %, PLA-5FU 23.56 % and that of 5-FU alone was 19.05 % after using bioluminescent intensity for 31 days. Then the researchers injected the same formulation into the abdomen of tumor-bearing mice. The results showed that the conjugated product had more tumor growth inhibition activity. PLA NPs conjugated with SM5-1 and loaded with 5-FU

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Table 11.1 Application, Administration, Release, and Delivery System of Polylactic Acid Conjugated With Other Polymers PLA/Other Polymers

Drug/Active Ingredient

Delivery System

Particle Size

Application

Administration

Release

Reference

PLA/chitosan PLA

Microspheres Nanoparticles

 210 nm

Chondrogenesis Antioxidant activity

Injectable 

 

Lao et al. (2008) Junior et al. (2017)

PLA

 Diphenyl diselenide (PhSe)2 Brimonidine

Microspheres



Glaucoma

Injection

Kim et al. (2014)

PLA/PCL

Clodronate

Microspheres



Osteoporosis



Sustain release Controlled

PLA/PHEA

Vitamin E

Nanoparticles





PEG-AcPLA

Nanocapsules

203 nm



Semi-longterm release 

PLA

Gemcitabine hydrochloride Nimesulide

High entrapment efficiency Antitumor activity

Microspheres

2.1 μm

NSAIDS delivery

I/M injection

Sustain

PLA/PLGA

HBsAg

Nanoparticles



Immunization

Oral



PLA/PLGA

Tetanus vaccine Diphtheria toxoid Docetaxel Docetaxel

Microspheres



Immunization





Microspheres

15 6 60 μm

Immunization





Nanoparticles Nanoparticles

 260 nm

Liver cancer Small cell lung cancer

 

 

Quercetin

Nanoparticles

152 6 9 nm

Breast cancer



Sustain

PLA/PEG

siRNA

Nanoparticles







PLA/dextran/ PLGA

rIL-2

Microspheres



Triple negative breast cancer Colon carcinoma

S/C injection

Sustain

PLA/PLGA PLA/PEG1000 PLA/peptide (AHSGMYP) PLA

Zhou et al. (2015) Cavallaro et al. (2015) Cosco et al. (2015) Freitas and Marchetti (2005) Jain et al. (2010) Johansen et al. (2000a) Johansen et al. (1999) Zhu et al. (2016) Zhu et al. (2016) Pandey et al. (2015) Liu et al. (2014d) Zhao et al. (2013)

PLA/PEGPPG-PEG PLA/PLGA

Camptothecin

Nanoparticles



Tumor treatment

I/V



Tacrolimus

Microspheres





I/V, S/C



PLA/PLGA

Omp antigen

Nanoparticles

,500 nm

Immunization

Intraperitoneal



PLA/chitosan

RAPA

Nanoparticles

300 nm





PLA/PEI

mcDNA, Doxorubicin HBsAg

Micelleplexes



Immunosuppression corneal allografts Cancer treatment



Controlled

Microparticles



Vaccines delivery





Nanospheres

,500 nm

Strangles treatment





Nanospheres



Cancer tissues

Parenteral route



Microspheres

100 μm

Drug delivery systems

Oral



Pietkiewicz et al. (2016) Ma (2014)

Microspheres



Drug delivery systems

Injection



Liu et al. (2006)

Nanoparticles Microspheres

 

Antitumor activity Theranostic, imaging

Intraperitoneal 

 

Ma et al. (2014) Jin et al. (2013)

CS, CSC, or PEI PLA/GCS, PVA PLA/ cremophor EL PLA/PLGA

PLA/PLGA PLA/SM5-1 PLA fabrication

Streptococcus equi, SeM Albumin Growth hormone (rhGH) human insulin (rhI) 5-FU Graphene oxide

Kunii et al. (2007) Kojima et al. (2015) Rauta and Nayak (2015) Yuan et al. (2008) Gaspar et al. (2015) Liu et al. (2014c) Liu et al. (2006)

408

CHAPTER 11 Polylactide

actively targeted the tumor cells with sustain drug release. Both in vitro and in vivo NPs showed good activity against tumor cells. Bioluminescent results of conjugated NPs also revealed antitumor activity. The researchers concluded that PLA NPs and SM5-1 with 5-FU could be effectively used to improve the therapeutic index of antitumor drugs like 5-FU and to provide an achievable method for antitumor evaluation in mice models in the near future (Ma et al., 2014). The effectively fabricated theranostic MPs into PLA microcapsules by presenting gold (Au) NPs into PLA microcapsules with the double microemulsion technique by deposition of graphene oxide layer wise onto surface of microcapsules by self-assembly technique was studied by Jin et al. (2013). The proofs obtained from these prepared microcapsules showed that these contrast agents served to enhance X-ray CT imaging and US imaging both in vitro and in vivo. These microcapsules showed higher efficacy against photothermal cancer therapy when therapeutic examination was carried out. The inhibition rate of tumor growth was 83.8% in the presence of these microcapsules when ablated in near IR light. The real-time ultrasound combined with 3D CT imaging helped in interpreting results, identifying location and size of tumor, as well as monitoring and guiding photothermic therapy. The microcapsules agents can enhance the effectiveness of CT imaging and US imaging. This microcapsule technique could bring multimodal imaging and successful loading of Au NPs into PLA microcapsules. These microcapsules not only serve as theranostic agents, but also enhance US and CT imaging. Such a versatile system can diagnose and monitor the therapy and multimodal imaging guide for treatment of cancer (Jin et al., 2013). Table 11.1 represents the various biomedical applications of PLA with different modifications.

11.6 CONCLUSION The researchers are attracted to the biomedical applications of PLA having the capability of bioadsorbability and biodegradability, prepared from renewable resources. The physical and chemical modifications in the structure of PLA have proven that a single polymer may have diverse and useful applications. Various drug delivery systems like microparticles, scaffolds, NPs, micelles, and nanospheres have been prepared from PLA because of its biodegradability and bioresorsability. PLA has been used in bone-fixation devices due to its durability. Recently, PLA has been used for the codelivery of drugs and genes or DNA materials for controlled and targeted delivery. It is currently the most promising biodegradable polymer extensively used as a biomaterial for multidimensional medical applications. PLA is synthesized from the raw materials like starches and agriculture-based products. Thus besides its biodegradable nature, it is also cost-effective, making it a promising candidate for biomedical applications. With certain modifications, it can be used for drug and DNA delivery to tissue engineering, imaging, and diagnoses. In future, researchers will rely on PLA polymers because of their low production cost and versatile biodegradable nature.

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Lowe, C.E., 1954. Preparation of high molecular weight polyhydroxyacetic ester. Google Patents. Lunt, J., 1998. Large-scale production, properties and commercial applications of polylactic acid polymers. Polym. Degradation Stability 59, 145152. Ma, G., 2014. Microencapsulation of protein drugs for drug delivery: strategy, preparation, and applications. J. Controlled Release 193, 324340. Ma, X., Cheng, Z., Jin, Y., Liang, X., Yang, X., Dai, Z., et al., 2014. SM5-1-conjugated PLA nanoparticles loaded with 5-fluorouracil for targeted hepatocellular carcinoma imaging and therapy. Biomaterials 35, 28782889. Manavitehrani, I., Fathi, A., Badr, H., Daly, S., Negahi Shirazi, A., Dehghani, F., 2016. Biomedical applications of biodegradable polyesters. Polymers 8, 20. Mathew, A.P., Oksman, K., Sain, M., 2005. Mechanical properties of biodegradable composites from poly lactic acid (PLA) and microcrystalline cellulose (MCC). J. Appl. Polym. Sci. 97, 20142025. Mehta, R., Kumar, V., Bhunia, H., Upadhyay, S., 2005. Synthesis of poly (lactic acid): a review. J. Macromol. Sci. Part C: Polym. Rev. 45, 325349. Nijenhuis, A., Grijpma, D., Pennings, A., 1991. Highly crystalline as-polymerized poly (Llactide). Polym. Bullet. 26, 7177. Nikou, K.N., Stivaktakis, N., Avgoustakis, K., Sotiropoulou, P.A., Perez, S.A., Baxevanis, C.N., et al., 2005. A HER-2/neu peptide admixed with PLA microspheres induces a Th1-biased immune response in mice. Biochim. Biophys. Acta (BBA)-General Subjects 1725, 182189. Ogaki, R., Green, F., Li, S., Vert, M., Alexander, M., Gilmore, I., et al., 2006. G-SIMS of biodegradable homo-polyesters. Appl. Surface Sci. 252, 67976800. Ohkita, T., Lee, S.H., 2006. Thermal degradation and biodegradability of poly (lactic acid)/ corn starch biocomposites. J. Appl. Polym. Sci. 100, 30093017. Oyama, H.T., 2009. Super-tough poly (lactic acid) materials: reactive blending with ethylene copolymer. Polymer 50, 747751. Palade, L.-I., Lehermeier, H.J., Dorgan, J.R., 2001. Melt rheology of high L-content poly (lactic acid). Macromolecules 34, 13841390. Pa´linko´-Biro´, E., Ro´nasze`ki, G., Merkle, H.P., Gander, B., 2001. Release kinetics and immunogenicity of parvovirus microencapsulated in PLA/PLGA microspheres. Int. J. Pharm. 221, 153157. Pandey, S.K., Patel, D.K., Thakur, R., Mishra, D.P., Maiti, P., Haldar, C., 2015. Anticancer evaluation of quercetin embedded PLA nanoparticles synthesized by emulsified nanoprecipitation. Int. J. Biol. Macromol. 75, 521529. Paul, M.-A., Delcourt, C., Alexandre, M., Dege´e, P., Monteverde, F., Dubois, P., 2005. Polylactide/montmorillonite nanocomposites: study of the hydrolytic degradation. Polym. Degradation Stability 87, 535542. Peres, C., Matos, A.I., Conniot, J., Sainz, V., Zupanˇciˇc, E., Silva, J.M., et al., 2016. Poly (lactic acid)-based particulate systems are promising tools for immune modulation. Acta Biomater 48, 4157. Pietkiewicz, J., Wilk, K.A., Bazyli´nska, U., 2016. In vitro studies of serum albumin interaction with poly (D, L-lactide) nanospheres loaded by hydrophobic cargo. J. Pharm. Biomed. Anal. 117, 426435. Pillin, I., Montrelay, N., Bourmaud, A., Grohens, Y., 2008. Effect of thermo-mechanical cycles on the physico-chemical properties of poly (lactic acid). Polym. Degradation Stability 93, 321328.

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Pluta, M., 2006. Melt compounding of polylactide/organoclay: structure and properties of nanocomposites. J. Polym. Sci. Part B: Polym. Phys. 44, 33923405. Polson, A.M., Southard, G.L., Dunn, R.L., Polson, A.P., Yewey, G.L., Swanbom, D.D., et al., 1995. Periodontal healing after guided tissue regeneration with Atrisorb barriers in beagle dogs. Int. J. Periodontics Restorative Dentistry 15, 575589. Pyda, M., Bopp, R., Wunderlich, B., 2004. Heat capacity of poly (lactic acid). J. Chem. Thermodyn. 36, 731742. Ramakrishna, S., Mayer, J., Wintermantel, E., Leong, K.W., 2001. Biomedical applications of polymer-composite materials: a review. Composites Sci. Technol. 61, 11891224. Rasal, R.M., Janorkar, A.V., Hirt, D.E., 2010. Poly (lactic acid) modifications. Progress Polym. Sci. 35, 338356. Rasoulianboroujeni, M., Yazdimamaghani, M., Khoshkenar, P., Pothineni, V.R., Kim, K. M., Murray, T.A., et al., 2016. From solvent-free microspheres to bioactive gradient scaffolds, Nanomed, 13. pp. 11571169. Rauta, P.R., Nayak, B., 2015. Parenteral immunization of PLA/PLGA nanoparticle encapsulating outer membrane protein (Omp) from Aeromonas hydrophila: evaluation of immunostimulatory action in Labeo rohita (rohu). Fish Shellfish Immunol. 44, 287294. Ray, S.S., Okamoto, M., 2003. New polylactide/layered silicate nanocomposites, 6 melt rheology and foam processing. Macromol. Mater. Eng. 936944. Ray, S.S., Bousmina, M., 2005. Biodegradable polymers and their layered silicate nanocomposites: in greening the 21st century materials world. Progress Mater. Sci. 50, 9621079. Ren, J., 2010. Application in the field of biomedical materials. Biodegradable Poly (Lactic Acid): Synthesis, Modification, Processing and Applications. Springer. Saeidlou, S., Huneault, M.A., Li, H., Park, C.B., 2012. Poly (lactic acid) crystallization. Progress Polym. Sci. 37, 16571677. Saini, P., Arora, M., Kumar, M.R., 2016. Poly (lactic acid) blends in biomedical applications. Adv. Drug Delivery Rev. 107, 4759. Sainz, V., Conniot, J., Matos, A.I., Peres, C., Zupanoˇioˇ, E., Moura, L., et al., 2015. Regulatory aspects on nanomedicines. Biochem. Biophys. Res. Communicat. 468, 504510. Sasatsu, M., Onishi, H., Machida, Y., 2006. In vitro and in vivo characterization of nanoparticles made of MeO-PEG amine/PLA block copolymer and PLA. Int. J. Pharm. 317, 167174. Schwach, E., 2004. Etude de syste`mes multiphases biode´gradables a` base d’amidon de ble´ plastifie´: relations structure-proprie´te´s, approches de la compatibilisation: the`se pour le doctorat en sciences, spe´cialite´ Chimie des Mate´riaux. Reims. Serizawa, S., Inoue, K., Iji, M., 2006. Kenaf-fiber-reinforced poly (lactic acid) used for electronic products. J. Appl. Polym. Sci. 100, 618624. Shin, B.Y., Narayan, R., 2010. Rheological and thermal properties of the PLA modified by electron beam irradiation in the presence of functional monomer. J. Polym. Environ. 18, 558566. Shue, L., Yufeng, Z., Mony, U., 2012. Biomaterials for periodontal regeneration: a review of ceramics and polymers. Biomatter 2, 271277. Sinha Ray, S., Yamada, K., Okamoto, M., Ogami, A., Ueda, K., 2003. New polylactide/ layered silicate nanocomposites. 3. High-performance biodegradable materials. Chem. Mater. 15, 14561465.

Further Reading

So¨derga˚rd, A., Stolt, M., 2002. Properties of lactic acid based polymers and their correlation with composition. Progress Polym. Sci. 27, 11231163. Spinu, M., Jackson, C., Keating, M., Gardner, K., 1996. Material design in poly (lactic acid) systems: block copolymers, star homo-and copolymers, and stereocomplexes. J. Macromol. Sci., Part A: Pure Appl. Chem. 33, 14971530. Stridsberg, K.M., Ryner, M., Albertsson, A.-C., 2002. Controlled ring-opening polymerization: polymers with designed macromolecular architecture. Degradable Aliphatic Polyesters. Springer. Surrao, D.C., Waldman, S.D., Amsden, B.G., 2012. Biomimetic poly (lactide) based fibrous scaffolds for ligament tissue engineering. Acta Biomaterialia 8, 39974006. Wang, Y., Yang, L., Niu, Y., Wang, Z., Zhang, J., Yu, F., et al., 2011. Rheological and topological characterizations of electron beam irradiation prepared long-chain branched polylactic acid. J. Appl. Polym. Sci. 122, 18571865. Wang, Q., Bao, Y., Ahire, J., Chao, Y., 2013. Co-encapsulation of biodegradable nanoparticles with silicon quantum dots and quercetin for monitored delivery. Adv. Healthcare Mater. 2, 459466. Yuan, X.-B., Yuan, Y.-B., Jiang, W., Liu, J., Tian, E.-J., Shun, H.-M., et al., 2008. Preparation of rapamycin-loaded chitosan/PLA nanoparticles for immunosuppression in corneal transplantation. Int. J. Pharm. 349, 241248. Yuryev, Y., Wood-Adams, P.M., 2012. Crystallization of poly (L-/D-lactide) in the presence of electric fields. Macromol. Chem. Phys. 213, 635642. ˙ Zenkiewicz, M., Richert, J., Rytlewski, P., Moraczewski, K., Stepczy´nska, M., Karasiewicz, T., 2009. Characterisation of multi-extruded poly (lactic acid). Polym. Testing 28, 412418. Zhao, H., Wu, F., Cai, Y., Chen, Y., Wei, L., Liu, Z., et al., 2013. Local antitumor effects of intratumoral delivery of rlL-2 loaded sustained-release dextran/PLGAPLA core/ shell microspheres. Int. J. Pharm. 450, 235240. Zhou, W., Qian, H., Yan, L., Luo, D., Xu, N., Wu, J., 2015. Controlled release of clodronate from PLA/PCL complex microsphere. Mater. Lett. 152, 293297. Zhu, D., Tao, W., Zhang, H., Liu, G., Wang, T., Zhang, L., et al., 2016. Docetaxel (DTX)loaded polydopamine-modified TPGS-PLA nanoparticles as a targeted drug delivery system for the treatment of liver cancer. Acta Biomater. 30, 144154.

FURTHER READING Mathew, A.P., Oksman, K., Sain, M., 2006. The effect of morphology and chemical characteristics of cellulose reinforcements on the crystallinity of polylactic acid. J. Appl. Polym. Sci. 101, 300310. Ogata, N., Jimenez, G., Kawai, H., Ogihara, T., 1997. Structure and thermal/mechanical properties of poly (L-lactide)-clay blend. J. Polym. Sci. Part B: Polym. Phys. 35, 389396.

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Poly(propylene fumarate)based biocomposites for tissue engineering applications

12 Ana M. Dı´ez-Pascual

Analytical Chemistry, Physical Chemistry and Chemical Engineering Department, Faculty of Biology, Environmental Sciences and Chemistry, Alcala´ University, Madrid, Spain

12.1 INTRODUCTION Currently, researchers from chemistry, materials, medicine, biology, and engineering are working together in the field of tissue engineering in an effort to overcome some of the limitations associated with current methods of bone repair. Both polymeric materials and ceramics have been investigated as effective pathways to repair bone defects. Amongst these materials, biodegradable synthetic polymers present numerous advantages for the development of scaffolds for tissue engineering, including the capacity to tailor their mechanical properties and degradation kinetics to fit various applications. Also, they can be designed with various shapes and desirable pore size and morphology conducive to tissue in-growth as well as permit the facile integration of chemical groups, which promotes tissue growth (Okamoto and John, 2013). Synthetic biopolymers of the polyester family, such as poly(L-lactic acid) (PLLA), poly(ε-caprolactone) (PCL), and poly(glycolic acid) (PGA), are amongst the most widely used for tissue engineering applications. Even so, their low strength, limited barrier properties, and lack of variety of functional groups in the polymer backbone have hindered their uses. Lately, fumaric acid-based polyesters have drawn great attention for use in the biomedical field owing to their excellent biocompatibility and biodegradability (Kasper et al., 2009). One of the most broadly explored is poly(propylene fumarate) (PPF), which is comprised of two ester bonds and one unsaturated carbon-carbon double bond (Fig. 12.1) that enables crosslinking both via free radical polymerization of monomers or via photoinitiation in the presence of different photoinitiators (Fisher et al., 2001). Crosslinked PPF fulfills the main requisites of bone substitutes since it is thermally nonconductive, sterilizable, osteoconductive, and biocompatible, shows sterilizability, biocompatibility, and handling characteristics. In addition, it degrades through the hydrolysis of the ester bonds into fumaric acid and propylene glycol, which are nontoxic and the degradation can be tailored Materials for Biomedical Engineering: Thermoset and Thermoplastic Polymers. DOI: https://doi.org/10.1016/B978-0-12-816874-5.00012-8 © 2019 Elsevier Inc. All rights reserved.

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FIGURE 12.1 Chemical structure of poly(propylene fumarate) (PPF).

by controlling parameters like molecular weight, crosslinker type, and crosslinking density. In order to enhance the mechanical performance of PPF and extend its applications, new strategies are pursued like the inclusion of fillers (Mistry et al., 2010; Lalwani et al., 2013; Diez-Pascual and Diez-Vicente, 2016a,b) or polymer blending (Wang et al., 2010a,b). PPF is frequently mixed with ceramic particles, including hydroxyapatite (HA), calcium carbonate, and calcium phosphate (Cai et al., 2009; Kamel et al., 2013) for orthopedic applications. The resulting composites show good compressive strengths (2
30 MPa), favorable osteoconductivity, and the capability to support cellular functions of bioceramics and are, therefore, appropriate for substitution of cancellous bone. Besides, preceding works have explored the effectiveness of fullerenes and carbon nanotubes (CNTs) as reinforcing agents for crosslinked PPF (Lalwani et al., 2013; Shi et al., 2005). Fullerenes slightly improved the mechanical properties of the polymer, whereas CNTs provoked modest enhancements. Conclusions drawn from those studies resulted in with two further methods to additionally improve the mechanical properties of nanofiller-reinforced PPF composites: covalent and noncovalent modification of nanofillers to circumvent the development of agglomerates and diminution of the aspect ratio of the filler. To attain good distribution of the nanofiller and tailor the nanocomposite microstructure, noncovalent or covalent modification with polymers may be necessary. The noncovalent modification involves the physical adsorption of polymers onto the surface of the through different interactions, such as Van der Waals forces, electrostatic, hydrogen bonding, or π 2 π stacking. Its major benefit is the preservation of the filler integrity and characteristics. The covalent route consists of the chemical anchoring (grafting) of polymer chains to functional groups of the nanofiller surface. It is versatile owing to the high density of surface groups of organic fillers. Nevertheless, it commonly causes defects to the nanofiller surface that have negative effects on its properties, particularly mechanical and electrical (Diez-Pascual et al., 2015).

12.2 Poly(Propylene Fumarate): Synthesis, Properties, and Applications

Polyethylene glycol (PEG), also known as polyethylene oxide (PEO) or polyoxyethylene (POE), is a polyether that has been widely used for the development of hydrogels for tissue engineering (Okamoto and John, 2013) owing to its biocompatibility, biodegradability, water and organic solvent solubility, low protein adhesion, and nonimmunogenicity (Beamish et al., 2010). In addition, the hydroxyl groups of this polyether can be simply modified with functional groups, like carboxyl, thiol, and acrylate, or grafted to other molecules or bioactive agents in order to prepare nanofillers for biomedical purposes. Numerous works have been published on the covalent functionalization of both organic [i.e., graphene (G) and its derivative graphene oxide (GO)] (Yang et al., 2011; Jin et al., 2012) and inorganic nanofillers (i.e., MoS2) with PEG (Presolski and Pumera, 2016). However, scarce works have dealt with PEG-functionalized graphene (or GO) via noncovalent chemistry (Park et al., 2011). The current chapter focuses on the preparation and characterization of novel bionanocomposites based on PPF reinforced with different concentrations of PEG-functionalized graphene oxide (PEG-GO) or PEG-grafted boron nitride nanotubes (PEG-g-BNNTs) with the aim of applying them in the field of tissue engineering. The composites were prepared through a combination of sonication and heat curing, and their surface morphology, thermal stability, hydrophilicity, water absorption, biodegradation, cytotoxicity, mechanical, viscoelastic, tribological, and antibacterial properties have been carefully analyzed through a variety of techniques. The next sections will explain in detail the effect of nanofiller type and concentration on the properties of the final nanocomposites. At the end, conclusions and future perspectives will be drawn.

12.2 POLY(PROPYLENE FUMARATE): SYNTHESIS, PROPERTIES, AND APPLICATIONS 12.2.1 SYNTHESIS Several methods have been reported to synthesize PPF through step-growth copolymerization, which can be divided into two groups according to the number of steps in the synthesis process. Using a one-step method, propylene glycol (PG), fumaric acid, and an acid catalyst were used by Frazier et al. to prepare PPF (Frazier et al., 1995). Nevertheless, to remove impurities and excess PG as well as to augment the chain length of the polymer, elevated temperatures were necessary. A straightforward esterification of fumaric acid and PG catalyzed by ptoluenesulfonic acid was also used by Gresser et al. to synthesize this copolyester (Gresser et al., 1995). Kharas et al. synthesized PPF via a two-step reaction of diethyl fumarate (DEF) and PG using ZnCl2 as a catalyst (Kharas et al., 1997). Afterwards, Kasper et al. described in detail a similar method with improved experimental conditions to synthesize 500
4000 Da PPF (Kasper et al., 2009). This route has been

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FIGURE 12.2 Schematic representation of the two-step process used for the preparation of PPF matrix. Taken from Diez-Pascual, A.M., Diez-Vicente, A.L., 2016b. PEGylated boron nitride nanotube-reinforced poly (propylene fumarate) nanocomposite biomaterials. RSC Adv. 6, 79507
79519, with permission from the Royal Society of Chemistry.

modified for the preparation of the PPF/PEG-GO and PPF/PEG-g-BNNTs composites (Fig. 12.2). Briefly, both reagents in a molar ratio of 1:3 were mixed in a three-neck round-bottomed flask using hydroquinone as a crosslinking inhibitor and ZnCl2 as a catalyst, which were added in a 0.003:0.01:1 molar ratio to DEF, respectively. The solution was maintained under nitrogen and mechanically mixed at 130oC for 10 hours. This step produced ethanol, which was collected as a distillate, %and bis(hydroxypropyl) fumarate (BHPF) intermediate. In the second step, the intermediate was heated to 150 C and the reaction was run for 8 hours after the application of a vacuum. The polymer was cooled down, washed repeatedly with dichloromethane followed by several acid washes to get rid of the catalyst and two washes with distilled water, and lastly dried over sodium sulfate.

12.2.2 PROPERTIES PPF is amorphous and displays a molecular weight-dependent glass transition temperature (Tg) that ranges between
30 C and 32 C, a thermal decomposition temperature of 345 C, a density of 0.998 g cm23, and a melt viscosity in the range of 102
104 Pa•s. It is an injectable, biocompatible, and biodegradable copolyester that leads to extractable degradation products, primarily fumaric acid and PG, upon hydrolysis of its ester linkages. It can be crosslinked both by chemical reaction or UV light, and its mechanical characteristics are highly influenced by the degree of crosslinking as well as the polymer molecular weight. Thus, compressive strengths in the range of 2
30 MPa, flexural strength in the range of 1.8
16.1 MPa, and flexural modulus from 1.1 to 1.4 GPa have been reported (Frazier et al., 1995; Kharas et al., 1997). Thus, its mechanical properties can be controlled, making it appropriate for application in bone scaffolds. Nonetheless, it is hydrophobic in nature, which negatively effects cell adhesion. To increase its hydrophilicity and extend its variety of medical applications, copolymerization

12.3 Graphene Oxide: Structure, Synthesis, and Properties

with hydrophilic polymers, like PEG, or modification with peptides can be used (Shin et al., 2011).

12.2.3 APPLICATIONS Due to their biocompatibility and tailorable mechanical performance, biomaterials based on PPF are ideal for use in orthopedic tissue engineering. For instance, for the substitution of cancellous bone, PPF is frequently mixed with ceramic particles, including calcium carbonate or calcium phosphate (Cai et al., 2009; Kamel et al., 2013). Moreover, for cardiovascular applications, reduced platelet adhesion has been achieved via the incorporation of PEG into PPF (Suggs et al., 1999). For bone defect repair, like recomposing the load-bearing capability of vertebral bodies, injectable scaffolds of polycaprolactone (PCL)/PPF copolymers have been manufactured (Fang et al., 2014). For craniofacial bone repair, bone cements made of unsaturated PPF and crosslinked PPF microparticles have been synthesized (Henslee et al., 2012). Another interesting application is drug delivery; in particular, thermoreversible methoxy poly(ethylene glycol) (mPEG)/ PPF block copolymers have been used for such purposes (Behravesh and Mikos, 2003). Drug delivery applications (i.e., drug carriers for tumor treatment, drug dispensers, etc.) and magnetic resonance imaging, to mention but a few, are amongst other uses of porous PPF scaffolds. Besides this, ultrafine polymer fibers prepared via electrospinning can be used as biosensors, bioactuators, and neural interfaces.

12.3 GRAPHENE OXIDE: STRUCTURE, SYNTHESIS, AND PROPERTIES 12.3.1 STRUCTURE Graphene oxide (GO), monolayer or few-layer oxygen-functionalized graphene, has gained a lot of interest as a nanostructured material (Dreyer et al., 2010). Due to the presence of oxygen functionalities on the GO surface, such as epoxides and hydroxyls on the basal planes and carboxyls on the edges (Fig. 12.3), it behaves strongly hydrophilic and can be easily dispersed in organic solvents, water, and different matrixes. This is the main advantage when combining this nanomaterial with polymer or ceramic matrixes to enhance their properties. It has been proposed that GO sheets comprise of flat, graphene-like aromatic domains connected via a network of cyclohexane-like units in chair configuration, which are decorated by oxygen groups (Haubner et al., 2010). Each GO layer is composed of a dense 2D carbonaceous skeleton with a great number of sp3 hybridized carbon atoms and a small number of sp2 carbons.

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FIGURE 12.3 Structure of graphene oxide (GO).

12.3.2 SYNTHESIS A pioneering work on the synthesis of GO was reported by Brodie in 1859 (Chen et al., 2012). In this method, graphite was mixed with KClO3 (1:3 w/w) and reacted in HNO3 at 60 C for 4 days. Staudenmaier improved the former approach by replacing HNO3 with concentrated H2SO4 and adding KClO3 in higher portions, which enables a single-step reaction, albeit it still requires a long period of time (about 4 days). The most broadly applied method for the fabrication of GO was developed by Hummers and Offeman in 1958 (Hummers method). In this case, the oxidation of graphite was attained via the treatment of graphite powder with KMnO4 and NaNO3 (1:3:0.5 w/w) in a concentrated H2SO4 solution. It presents three important advantages over previous techniques: Firstly, the reaction can be completed within a few hours. Secondly, KMnO4 is employed instead of KClO3, which improves the reaction safety by preventing the release of explosive ClO2. Thirdly, the replacement of HNO3 by NaNO3 eliminates the formation of acid fog. The Hummers method has been extensively applied owed to its high efficiency and reaction safety. However, it still has two weaknesses: the oxidation process releases toxic gasses, such as NO2 and N2O4, and the residual Na1 and NO3 ions are difficult to eliminate from the waste water formed during the synthesis and purification stages. Tour et al. (Marcano et al., 2010) improved the Hummers method by eliminating the NaNO3, raising the amount of KMnO4, and performing the reaction in a mixture of H2SO4/H3PO4 (9:1 v/v). This variant increases the reaction yield, reduces toxic gas evolution, and uses twice as much KMnO4 and 5.2 times as much H2SO4 as those employed by the Hummers method. Baek et al. carried out the etching of the basal planes of highly ordered pyrolytic graphite with a mixture of H2SO4 and HNO3 (Shin et al., 2013), and attained an effective shortening and cutting of the graphene layers after a long-term treatment. This reveals that the H2SO4/HNO3 mixture used in the Hummers method acts as a chemical scissor and drill to facilitate the penetration of the oxidation

12.3 Graphene Oxide: Structure, Synthesis, and Properties

solution between graphene planes. On the other hand, KMnO4 is one of the strongest oxidants in the acid medium, hence, with its aid, a total intercalation of graphite with concentrated H2SO4 can be attained, in which each graphene sheet is sandwiched by bisulfate ions. This complete intercalation ensures the full oxidation of graphite, hence, NaNO3 is no longer necessary for the synthesis of GO. Numerous variants of these methods have been developed, with improvements continually being investigated to attain superior results at a lower cost. For the preparation of GO-reinforced PPF composites, GO was synthesized using a modified Hummer’s method (Diez-Pascual and Diez-Vicente, 2016a). Briefly, graphite powder, H2SO4, K2S2O8, and P2O5 were heated at 80 C for 6 hours. After cooling, deionized water was added and subsequently the product was filtered, dried, and oxidized via the addition of H2SO4, KMnO4, and water in an ice-water bath. Upon dilution with water, excess KMnO4 was decomposed through the addition of a H2O2 aqueous solution. The product was filtered once more, washed with deionized water, and finally freeze-dried at reduced pressure.

12.3.3 PROPERTIES GO has excellent electronic, thermal, optical, and mechanical properties, with some values that go beyond those attained in any other material (Dreyer et al., 2010). In particular, it presents outstanding thermal conductivity (4000
6000 W m21 K21), large surface area (B2000 m2 g21), high electron mobility (105 cm2 V21 s21), a Young’s modulus of 207 GPa, and strength of about 100 GPa (Suk et al., 2010). Further, it is an electrical insulator because of the disruption of the sp2 bonding. These unique properties make GO a perfect candidate for a wide range of applications, such as sensors, electrical devices, batteries, supercapacitors, and solar cells, amongst others (Zhu et al., 2010). GO can be mixed with different polymers to enhance the performance of the resulting composite materials. The utmost improvements can be attained when GO is regularly dispersed within the matrix and the external load is effectively transferred via strong GO
polymer interfacial adhesion (Diez-Pascual et al., 2015). Nonetheless, the large GO surface area and strong van der Waals forces between the sheets can provoke important aggregation in the composite matrix. To achieve good GO distribution and tailor the nanocomposite microstructure, noncovalent or covalent functionalization with polymers is typically required (Salavagione et al., 2014). The noncovalent functionalization, which is based on van der Waals forces, H-bonding, electrostatic, or π 2 π stacking interactions, offers effective ways to control the properties and solubility of the GO flakes without modifying their chemical structure. Prior to synthesizing the PPF nanocomposites, PEG was used to noncovalently functionalize the matrix, and this process resulted in the exfoliation of the carbon nanomaterial into smaller sheets (about 50 nm thick) since the hydrogen bonding between the oxygenated moieties of GO and the OH groups of PEG prevail over the π 2 π stacking interactions that held together the GO layers (Diez-Pascual and Diez-Vicente, 2016a).

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Furthermore, the PPF chains intercalated between the GO flakes, although complete exfoliation was not accomplished given that no individual sheets were observed. Moreover, the comparison of the thickness and width distribution histograms of GO and PEG-GO demonstrated a unimodal distribution of GO thicknesses, in the range of 20
100 nm with a mean of 65 nm, and a broad trimodal distribution of GO width, ranging between 0.2 and 1.2 μm, with maxima at 0.25, 0.55, and 0.95 μm, and an average value of 0.60 μm. Conversely, both the sheet thickness and width distribution of PEG-GO were unimodal, with average values of 55 nm and 0.54 μm, respectively, and the polydispersity was reduced. This confirms that the sonication stage in the presence of PEG polymer induces the exfoliation of GO flakes in thinner sheets. The covalent functionalization relies on the reaction between the functional groups on the GO surface and certain functional groups on the polymer, and gives versatile possibilities owing to the rich surface chemistry of GO. On the other hand, the anchoring of GO to polymeric segments usually causes defects on the sheets that can have disadvantageous effects on certain properties, especially thermal and mechanical (Georgakilas et al., 2012).

12.4 BORON NITRIDE NANOTUBES: STRUCTURE, SYNTHESIS, AND PROPERTIES 12.4.1 STRUCTURE The metastable allotropic form of hexagonal boron nitride nanotubes was first proposed by Rubio et al. based on the analogies between the lattice structure of graphite and hexagonal boron nitride, Fig. 12.4 (Rubio et al., 1994). They are structural analogs of CNTs but with boron (B) and nitrogen (N) atoms instead of carbon atoms. BNNTs can be visualized as a rolled up hexagonal boron nitride layer and they have different chiralities (Wang et al., 2010a,b). According to this, three types of single-walled BNNTs have been described (Fig. 12.4): (A) arm-chair; (B) zig-zag; and (C) chiral. The B
N bonds have a somewhat ionic character since the electron density of B is attracted to the N atoms owing to their higher electronegativity. This provokes a large band gap between the valence and conduction bands of B5.5 eV; hence they are electrically insulating.

12.4.2 SYNTHESIS BNNTs can be synthesized by a variety of techniques. The first works on BNNTs employed techniques similar to those used for the synthesis of CNTs, such as arcdischarge, laser heating and vaporization, chemical vapor deposition (CVD), and high temperature ball milling. Complete reviews on the synthesis approaches of BNNTs can be found elsewhere (Zhi et al., 2010; Kalay et al., 2015). Among the different paths employing the CVD approach, those reported by Tang et al. should

12.4 Boron Nitride Nanotubes: Structure, Synthesis, and Properties

FIGURE 12.4 Left: Lattice structure of hexagonal boron nitride; Right: Atomic models of BNNTs.

be highlighted, who employed metal oxides, like MgO, FeO, and Li2O, as precursors to synthesize reactive boron oxide (BxOy) vapor that reacted under an NH3 atmosphere in an induction heating chamber at 1300 C to produce the BNNTs (Tang et al., 2002). This approach has been further improved by Zhi et al. to produce high-purity BNNTs (Zhi et al., 2005). This procedure was successful for synthesizing BNNTs in a lab, although the commercialization and practical use of BNNTs in industry is hampered by their low production rate and the necessity for a tailored chamber design. Using an analogous method, Lee et al. reported the preparation of high-purity and high-quality BNNTs at 1200 C in a traditional tube furnace (Lee et al., 2008). The fundamental characteristic of this technique is the use of a quartz tube closed at the end to entrap the vapors for the growth of the BNNTs. This method is denominated by the growth vapor trapping (GVT) approach, and can be tailored using catalytic nanoparticles coated on Si substrates. The raw BNNTs produced by this CVD/GVT method are verticallyaligned and display super-hydrophobicity due to their high surface roughness and reduced surface energy. High temperature ball milling is another widely used approach for the synthesis of BNNTs. Typically, boron powder is ball milled in a N2/NH3 atmosphere for about 150 hours, and subsequently mixed with ferric nitrate and cobalt nitrate in ethanol to create B ink. Afterward, the solution is treated with N2/NH3 for a few hours at 1300 C to yield the BNNTs. Nanotubes produced using the CVD method display a tubular structure with typical outer diameters in the range of 50
80 nm and lengths between 100 and 200 μm, while those synthesized via ball milling usually have a bamboo-like structure owing to the metallic particle impurities.

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The laser ablation technique has been employed for the synthesis of BNNTs since 1996. The nanotubes were grown on substrates at 600oC using Fe catalysts % via pulse laser deposition of a BN target. It results in thin nanotubes, with typical diameters in the range of 10
20 nm. Using this laser-ablation approach, singlewalled BNNTs can be synthesized via laser heating at 1 atmosphere under N2. A modification of this method was developed in 2009 using a laser evaporation technique (Smith et al., 2009). A CO2 laser was employed to vaporize a boron target located in a room filled with high pressure nitrogen which was condensed into liquid boron droplets that acted as nucleation sites. This approach results in multiwalled BNNTs with a cotton-ball like structure. A thermal plasma reactor was employed for large-scale synthesis of BNNTs with a mean diameter of 5 nm (Kim et al., 2014). Boron nitride powder and N2 and H2 gases were introduced in a high temperature induction plasma at atmospheric pressure. The precursor materials were decomposed into their elements and nanosized boron droplets were condensed in the reactor and further acted as nucleation sites to grow BNNTs. These nanotubes displayed a cloth-like morphology and appeared as an entangled network of fibrils with nontubular impurities, like amorphous boron and boron nitride. These plasma techniques have several drawbacks: they require high synthesis temperatures in the range of 4000K
8000K as well as purification of the nanotubes and a chamber for high pressure and fast cooling conditions in order to obtain elevated mass production. A number of approaches have been developed to purify raw BNNTs (Lee et al., 2011), including acid treatment, thermal oxidation in air atmosphere, and the addition of a surfactant or polymer wrapping separation by means of functionalization. Despite BNNTs being chemically inert and resistant to oxidation up to 1000 C, ultrasonication processes and treatments with strong acids result in the damaging and shortening of the BNNTs, and thus modify their intrinsic properties. Boron nitride impurities are difficult to eliminate given that they are also chemically stable and oxidation resistant. For the preparation of PPF/PEG-g-BNNT composites, BNNTs were first synthesized by CVD following a variant of the method reported earlier (Ferreira et al., 2011). In short, amorphous boron, NH4NO3, and Fe2O3 powders were mixed in an alumina boat in a 15:15:1 mass ratio and heated to 600 C. The temperature was steadily increased to 1200 C under a nitrogen flow and maintained at a constant level for 2 hours. Upon cooling, BNNTs were purified by washing with a 4 M HCl solution at 90 C. Finally, they were washed with deionized water and dried overnight. The resulting nanotubes were uniform, with a bamboo-like arrangement, and showed outer diameters between 30 and 80 nm, and lengths higher than 5 μm (Diez-Pascual and Diez-Vicente, 2016b). Afterward, the BNNTs were functionalized via the application of a two-step process, as illustrated in Fig. 12.5. Initially, they were sonicated in an acid solution (HNO3 65% w/w) for 4 hours at a frequency of 80 W and washed with ethanol by ultracentrifugation. This process leads to hydroxylated nanotubes (BNNTs-OH) with a functionalization degree (FD) of about 13.3%, as roughly

12.4 Boron Nitride Nanotubes: Structure, Synthesis, and Properties

FIGURE 12.5 Representation of the covalent functionalization of BNNTs with PEG. Reprinted from Diez-Pascual, A.M., Diez-Vicente, A.L., 2016b. PEGylated boron nitride nanotube-reinforced poly(propylene fumarate) nanocomposite biomaterials. RSC Adv. 6, 79507
79519, with permission from the Royal Society of Chemistry.

estimated from thermogravimetric curves (Diez-Pascual and Diez-Vicente, 2016b). The BNNTs-OH were then added to a silane PEG (PEG-Si) waterethanol solution, and the mixture was then ultrasonicated for 8 hours at a frequency of 40 W. The resulting grafted product (PEG-g-BNNTs) was afterward washed repeatedly with deionized water by ultracentrifugation and dried under vacuum. The extent of the grafting reaction was determined as 29% based on the difference between the FD of the BNNTs prior to and after the grafting reaction. According to SEM images (Diez-Pascual and Diez-Vicente, 2016b), PEG-gBNNT is a nonhomogeneous mixture comprising of free PEG chains that physically interact with the hydroxylated nanotubes and other polymeric segments chemically bonded to the surface of the BNNTs. Subsequent to the polymer grafting, the bamboo-like structure changed into a cylindrical shape, the surface was coarser, and the nanotube diameter raised due to the polymer wrapping.

12.4.3 PROPERTIES BNNTs display outstanding properties, such as excellent chemical stability, superior thermal conductivity, high elastic modulus (up to 1.3 TPa), piezoelectricity, resistance to oxidation and heat, hydrophobicity, hydrogen storage capacity, radiation absorption, lubricant behavior, and good biocompatibility (Kalay et al., 2015; Zhou et al., 2012). Due to these properties, they can be used to prepare composite materials with enhanced properties and in a number of applications ranging from protective shields, self-cleaning materials or nanoscale electrical devices working under extreme conditions to drug delivery, biosensors, and neutron capture therapy. Nonetheless, their hydrophobicity and limited dispersion in conventional solvents restricts their applications, and a lot of effort has been made to improve

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their dispersibility via surface modification, including noncovalent (Ciofani et al., 2012) and covalent approaches (Wang et al., 2008). Noncovalent methods were attained by wrapping the BNNTs with suitable surfactants/polymers. For instance, Zhi et al. dispersed the BNNTs in water via a single-strand DNA wrapping (Zhi et al., 2007); Chen et al. employed an amphiphilic dendritic structure that interacted with the BNNT surface by π 2 π stacking and hydrophobic interactions (Chen et al., 2009); and Wang et al. used an anionic perylene derivative with carboxylic groups to increase their aqueous dispersibility (Wang et al., 2008). However, large amounts of polymer can be detrimental for the material properties and the poor scalability is another drawback of the noncovalent process. On the other hand, covalent modifications of BNNTs have scarcely been explored owing to their chemical inertness. Only a few pathways have been reported, namely the anchoring of functional groups to the defect sites on the BNNT sidewalls (Ciofani et al., 2012), the formation of chemical bonds, like F
B, based on boron chemistry (Tang et al., 2005), or the reduction of the nanotubes to get them negatively charged (Shin et al., 2015).

12.5 PREPARATION OF PPF-BASED BIOCOMPOSITES PPF nanocomposites incorporating different amounts of PEG-g-BNNT or PEGmodified GO as nanofillers have been prepared via sonication and thermal curing (Diez-Pascual and Diez-Vicente, 2016a; Diez-Pascual and Diez-Vicente, 2016b). Firstly, the desired amount of PEG-g-BNNTs and PEG/GO mixtures were suspended in water and chloroform respectively, by bath sonication for 15 minutes. The resulting dispersions were found to be uniform and stable for more than 3 months. On the other hand, PPF and the crosslinker, N-vinyl-pyrrolidone (NVP), were mixed (1:1 w/w) and afterward the corresponding amount of PEG-g-BNNT or PEG/GO dispersion was added. The mixture was then ultrasonicated for an additional 15 minutes, and later the free radical initiator, benzoyl peroxide (BP), was dissolved in DEF and added to the mixture to begin the polymerization. The mixture was finally casted onto Teflon molds and cured under reduced pressure at 80 C for 24 hours. For each system, five nanocomposites were prepared with nanofiller weight percentages in the range of 0.1
4.0 wt.%.

12.6 CHARACTERIZATION OF PPF-BASED BIONANOCOMPOSITES 12.6.1 MORPHOLOGY AND STRUCTURE The surface morphologies of the nanocomposites were examined by scanning electron microscopy (SEM) with a SU8000 Hitachi microscope applying an

12.6 Characterization of PPF-Based Bionanocomposites

FIGURE 12.6 Representative SEM images at different magnifications of a PPF/PEG-modified GO nanocomposite with 3.0 wt.% GO loading. Adapted from Diez-Pascual, A.M., Diez-Vicente, A.L., 2016b. PEGylated boron nitride nanotube-reinforced poly(propylene fumarate) nanocomposite biomaterials. RSC Adv. 6, 79507
79519, with permission from the American Chemical Society.

acceleration voltage of 1.0 kV. Representative SEM micrographs of a PPF/PEGGO nanocomposite (3 wt.% GO loading) are displayed in Fig. 12.6. Pristine GO is a powder containing stacked sheets held by π 2 π, hydrogen bonding, and van der Waals interactions which exfoliate after noncovalent functionalization with PEG (Diez-Pascual and Diez-Vicente, 2016a). In contrast, the PEG-GO sheets are regularly distributed inside the PPF matrix. The micrographs shown in Fig. 12.6A
C show slim crumpled GO layers with thicknesses between 10 and 50 nm and an average value of B20 nm. The layers are well exfoliated but not individually dispersed and no GO monolayers can be observed. Furthermore, the flake edges have a tendency to roll and wrinkle, as detected at higher magnifications (Fig. 12.6C). A similar morphology was found for the rest of the composites, without the formation of aggregates. Fig. 12.7 shows the morphology of pure BNNTs, PEG-g-BNNTs, and a PPF/ PEG-g-BNNT nanocomposite with 4.0 wt.% loading. Raw BNNTs are collected in small bundles (i.e., 5 μm length and 30
80 nm diameter, Fig. 12.7A). Upon functionalization with the PEG chains, the diameter increases and the bamboolike appearance vanishes (Fig. 12.7B). The PEG-g-BNNTs have a cylindrical shape and show a rougher surface owing to the polymeric wrapping; they are homogenously distributed throughout the polymer matrix (Fig. 12.7C), revealing a twisted, entangled, and bundled arrangement and the absence of holes indicates good PPF-PEG-g-BNNTs compatibility, since the hydrogen bonding and polar interactions between the ester groups of the biopolymer and the ether and hydroxyl groups of the PEG-g-BNNTs improve the interfacial strength and prevent nanotube aggregation. The mean surface roughness of PPF and the biocomposites was measured, since it is an important factor that controls cell adhesion and proliferation: the higher the roughness, the better the adhesion strength (Chang and Wang, 2011). The roughness was found to rise with increasing GO loading (Diez-Pascual and

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FIGURE 12.7 Typical SEM images of (A) BNNTs, (B) PEG-g-BNNTs, and (C) PPF/PEG-g-BNNTs (4.0 wt.%). Adapted from Diez-Pascual, A.M., Diez-Vicente, A.L., 2016b. PEGylated boron nitride nanotube-reinforced poly(propylene fumarate) nanocomposite biomaterials. RSC Adv. 6, 79507
79519, with permission from the Royal Society of Chemistry.

Diez-Vicente, 2016a) from a value of about 1.3 μm for neat PPF to about 2.4 μm for the PPF/PEG-GO (3.0 wt.%). Thus, the improved surface roughness of the GO-reinforced nanocomposites compared to PPF would be beneficial for cell proliferation and differentiation. To obtain information about the filler
PPF interactions, the infrared spectra of BNNTs, PEG-g-BNNTs, PPF, and the nanocomposites were recorded and these are shown in Fig. 12.8. The spectrum of pristine BNNTs presents two features at 1368 and 790 cm21, attributed to the in-plane B
N stretching and the out-ofplane B
N
B bending vibrations, respectively (Sainsbury et al., 2012). The spectrum of PEG-g-BNNTs shows peaks at 3470 and 2890 cm21 related to the O
H and C
H stretching at about 1462 and 1350 cm21 ascribed to C
H bending vibrations, as well as bands in the range of 1280
1085 cm21 corresponding to C
O
C stretching vibrations. Moreover, Si
C and Si
O stretching vibrations are found in the range of 774
825 and 900
1050 cm21 respectively, and the Si
OCH3 rocking mode appears at 1192 cm21 (Diez-Pascual and Diez-Vicente, 2016b). Furthermore, broad strong peaks can be observed in the range of 1660
1280 cm21 due to the in-plane B
N stretching of the nanotubes, and the O
H stretching band shows higher intensity and is shifted toward a higher wavenumber compared to that of PEG-Si (Diez-Pascual and Diez-Vicente, 2016b); a phenomenon that has been ascribed to a change from intramolecular to intermolecular hydroxyl
hydroxyl interactions and that points toward H-bond formation between the BNNTs and PEG-Si. Most notably, the occurrence of a band at B1000 cm21 due to B
O
Si stretching vibrations confirms the grafting reaction. The spectrum of neat PPF shows a very intense band at around 1730 cm21 (inset of Fig. 12.8), attributed to the CQO stretching of the ester group. The C
O
C stretching vibrations appear between 1280 and 1067 cm21, the C
H stretching modes in the range of 2900
3000 cm21, the CQC stretching at 1652 cm21, and the H
C
O in-plane bending at 1385 cm21. The spectra of the

12.6 Characterization of PPF-Based Bionanocomposites

FIGURE 12.8 FTIR spectra of BNNTs, PEG-g-BNNTs, PPF, and the nanocomposites with 0.5 and 4.0 wt.% PEG-g-BNNTs content. The inset is a magnification showing the carbonyl band. Adapted from Diez-Pascual, A.M., Diez-Vicente, A.L., 2016b. PEGylated boron nitride nanotube-reinforced poly(propylene fumarate) nanocomposite biomaterials. RSC Adv. 6, 79507
79519, with permission from the Royal Society of Chemistry.

nanocomposites are similar to that of PPF, with some distinctive peaks at 3500, 2900, 1192, 1000, and 774 cm21 arising from the PEG-g-BNNTs; further, the intensity of these peaks increases upon raising BNNT content. The O
H stretching vibration appears wider in the nanocomposites compared to that of PEG-gBNNTs and the carbonyl peak is also broader and shifted to a lower wavenumber by B7 and 16 cm21 for the nanocomposites with 0.5 and 4.0 wt.% (inset of Fig. 12.8), suggesting H-bond formation with the BNNTs.

12.6.2 HYDROPHILICITY, BIODEGRADABILITY, AND PROTEIN ADSORPTION Prior to use the PPF-based nanocomposites for medical applications it is essential to assess their level of hydrophilicity, which can be evaluated by water contact angle (θ) measurements. The lower the θ, the more hydrophilic the surface and the better the cell attachment is. The best θ values for protein adsorption are between 40 and 70 degrees (Diez-Pascual and Diez-Vicente, 2015). Table 12.1 summarizes the contact angle data for the different PPF-based nanocomposites. PPF can be considered a hydrophilic material because it has a θ value ,90

431

Table 12.1 Contact Angle (θ), Water Uptake, Weight Loss in PBS, Protein Adsorption, Initial Degradation Temperature (Ti), Temperature of 10% Weight Loss (T10), and Temperature of Maximum Rate of Weight Loss (Tmax), for PPF and Composites With Different Nanofiller Weight Percentages (Values in Parenthesis) Nanofiller (wt.%)

θ ( C)

Protein concentration (µg mL21)

Water uptake (%)

Weight loss (%)

Ti ( C)

T10 ( C)

Tmax ( C)


PEG-GO (0.1) PEG-GO (0.5) PEG-GO (1.0) PEG-GO (2.0) PEG-GO (3.0) PEG-g-BNNTs PEG-g-BNNTs PEG-g-BNNTs PEG-g-BNNTs PEG-g-BNNTs

66.7 61.2 57.5 53.1 45.9 43.4 68.2 64.1 59.5 55.4 50.9

1.35 2.24 2.48 2.86 3.83 4.47 1.28 1.49 2.40 2.74 2.81

2.9 4.3 5.1 7.2 8.4 8.8 2.6 2.8 3.4 3.9 4.8

3.5 3.7 4.0 4.4 4.9 5.6 2.7 3.6 3.8 4.5 5.4

285.2 283.0 284.8 297.5 303.7 311.6 279.3 261.4 254.9 247.2 244.6

309.0 311.3 317.5 329.2 333.1 346.6 311.7 319.2 328.9 334.3 341.6

352.2 364.8 370.3 379.1 388.5 397.8 368.2 375.0 381.9 385.2 390.8

(0.1) (0.5) (1.0) (2.0) (4.0)

12.6 Characterization of PPF-Based Bionanocomposites

degrees. A fall in θ is found for both types of nanocomposites with increasing nanofiller loading, indicating superior wettability related to the great number of surface oxygen-containing groups of GO and BNNTs along with the OH groups of PEG that can easily form H-bonds with the water molecules. All the biocomposites exhibit θ values in the range of 43
68 degrees and are, therefore, likely appropriate for cell adhesion and proliferation. Composites with PEG-GO as fillers display higher hydrophilicity than their PEG-g-BNNT counterparts, likely because the oxygen functionalities on the GO surface (i.e., epoxides, hydroxyls, and carboxyls) make it highly hydrophilic, whereas BNNTs have been reported to be hydrophobic (Kalay et al., 2015). Thus, the strongest drop in θ (about 35%) is found for the nanocomposite with 3.0 wt.% PEG-GO (Table 12.1). The water uptake of the nanocomposites (Table 12.1) was calculated by immersing them in a simulated body fluid (SBF) at 37 C for 14 days and weighing them before and after immersion in SFB. Neat PPF has a low value (around 3%), which increases steadily upon increasing nanofiller content, which is consistent with the higher level of hydrophilicity of the nanocomposites. Accordingly, composites with PEG-GO as a nanofiller present higher water absorption than those filled with PEG-g-BNNTs, and the composite with the highest PEG-GO content has a value close to 9%. Nevertheless, the nanocomposites show fairly good water resistance, which is important in maintaining their dimensional stability upon getting in contact with fluids. An in vitro degradation study on phosphate buffered saline (PBS) at pH 7.4 and 37 C for 7 weeks was carried out to evaluate the biodegradation of nanocomposites. A high biodegradation speed is usually preferred for biomedical applications to ensure rapid elimination from the body. The percentage weight loss in PBS for the different nanocomposites is summarized in Table 12.1. PPF exhibits a weight loss of B3.5% owing to the hydrolytic degradation of the ester linkages, which produces fumaric acid and PG as the two main degradation products (Kalay et al., 2015). An augment in weight loss is detected for both types of nanocomposites, following their analogous tendency to uptake water, and the largest increment (around 60%) is again found for the sample with 3.0 wt.% PEG-GO. A direct correlation between protein adsorption and cell-growth has been reported (Chang and Wang, 2011). Therefore, the protein absorption ability of PPF-based nanocomposites was measured via immersion of the samples in a medium containing 10% fetal bovine serum for 4 hours at 37 C. Afterward, they were placed into well plates and washed three times with PBS followed by the addition of 400 μL of a 1% sodium dodecyl sulfate (SDS) solution. The concentration was assessed using a Micro BCA protein assay kit, and the results are presented in Table 12.1. An increase in protein concentration was detected for both types of nanocomposites, leading to an approximately 2 and 3.5-fold increment for the highest PEG-g-BNNT and PEG-GO concentration, respectively. Again, higher protein adsorption is found for composites with PEG-GO, related to their

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high wettability, hence, they are expected to be better for promoting cell attachment and proliferation.

12.6.3 THERMAL PROPERTIES Thermal stability was analyzed by thermogravimetric analysis (TGA) under an inert atmosphere and the results for the different composites in terms of the initial degradation temperature (Ti), temperature of 10% weight loss (T10), and temperature of maximum rate of weight loss (Tmax) are gathered in Table 12.1. PPF displays only one decomposition stage that begins at approximately 285 C and the Tmax at about 352 C (Table 12.1). A one-step decomposition process was also observed for PPF/PEG-GO nanocomposites (Diez-Pascual and Diez-Vicente, 2016a), while PPF/PEG-g-BNNTs exhibited two stages; the first related to the elimination of OH groups from the BNNT surface and the second to the decomposition of the polymeric backbone. As shown in Table 12.1, the Ti of composites reinforced with PEG-GO increases with increasing GO concentration, while that of composites with PEG-g-BNNTs decreases, since PEG-silane has lower thermal stability than PPF. Conversely, the T10 and Tmax of both types of nanocomposites rise upon increasing nanofiller content, the maximum increments being around 37 C and 45 C respectively, for the nanocomposite with 3.0 wt.% PEG-GO. Thus, composites with PEG-GO exhibit improved thermal stability compared to those with PEG-g-BNNTs, likely due to the large surface area of GO that acts as a barrier to protect the polymeric chains from the flame and delays the diffusion of decomposition products from the bulk of the polymer to the gas phase by means of the formation of a tortuous path. In addition, the strong PPF-PEG and PPF-GO interactions via H-bonding and dipole
dipole intermolecular forces would confine chain rotational movements, therefore, diminishing the amplitude of the molecules moving under the temperature influence, which leads to better thermal stability.

12.6.4 MECHANICAL PROPERTIES Certain mechanical properties of biomaterials are essential for tissue engineering applications; in general, an adequate balance between flexibility and strength is required. The tensile properties of PPF-based nanocomposites, namely the Young’s modulus (E), tensile strength (σy), elongation at break (εb), and toughness (T), measured under both dry (23 C and 50% RH) and wet (37 C in SBF for 6 weeks) conditions, according to the UNE-EN ISO 527-1 standard, are summarized in Table 12.2. The E of PPF is B1 GPa and rises steadily as the nanofiller concentration raises, the highest augment being around 200% with the addition of 3.0 wt.% PEG-GO. Interestingly, the increment is methodically larger for nanocomposites with PEG-GO than for nanocomposites with PEG-g-BNNTs, notwithstanding the moduli of the BNNTs (750
1200 GPa) (Santosh et al., 2009) is superior than the reported GO modulus (about 200 GPa) (Suk et al., 2010). The

12.6 Characterization of PPF-Based Bionanocomposites

Table 12.2 Mechanical Properties of PPF-Based Nanocomposites Nanofiller (wt.%)

E (GPa)

σ y (MPa)

εb (%)

T (MJ m23)


PEG-GO (0.1) PEG-GO (0.5) PEG-GO (1.0) PEG-GO (2.0) PEG-GO (3.0) PEG-g-BNNTs PEG-g-BNNTs PEG-g-BNNTs PEG-g-BNNTs PEG-g-BNNTs

0.99/0.66 1.36/0.72 1.72/0.98 2.33/1.08 2.51/1.20 2.91/1.26 1.14/0.95 1.29/1.15 1.72/1.36 1.98/1.39 2.34/1.62

39.1/30.1 47.2/34.9 59.4/42.6 70.3/48.8 74.9/47.2 90.2/54.5 37.3/35.1 45.2/40.2 56.3/48.7 59.5/49.4 67.2/54.8

4.31/3.71 4.42/3.84 3.78/3.66 3.33/3.35 2.67/3.24 2.38/3.37 4.42/3.73 3.87/3.44 3.91/3.35 3.52/3.62 3.06/3.26

6.9/4.6 7.2/5.2 9.6/5.7 9.5/6.3 7.9/6.2 8.4/6.9 6.3/5.9 6.6/6.2 8.9/7.0 9.2/7.4 9.4/7.8

(0.1) (0.5) (1.0) (2.0) (4.0)

E0 25 (GPa)

Tg ( C)

0.85 1.32 1.62 2.12 2.28 2.76 0.82 1.43 1.68 1.98 2.26

23.1 23.0 26.6 31.4 33.7 36.2 22.5 27.3 34.9 35.2 37.3

Young’s modulus (E), Tensile strength (σy), Elongation at break (εb), Toughness (T), Storage modulus (E0 ) at 25 C, and Glass transition temperature (Tg). The first and second value in each column correspond to the values measured under dry (23 C and 50% RH) and wet (37 C in SBF) environments respectively.

superior enhancement for the composites with PEG-GO is probably connected to better nanofiller distribution and strengthened PPF
nanofiller interfacial adhesion by polar and hydrogen bonding interactions. Moreover, the reinforcement effect observed for these PPF/PEG-GO composites is larger than that found upon the addition of the equivalent contents of other nanofillers, like CNTs (Shi et al., 2005) or fullerenes (Lalwani et al., 2013). A comparable trend is detected for tensile strength (Table 12.2), corroborating that noncovalent functionalization is more advantageous for improving the mechanical performance of PPF than the covalent approach, maybe because the oxidative process with HNO3 used during the grafting of PEG segments to the BNNTs generated defects on the walls of the nanotubes that are disadvantageous for the mechanical performance. Regarding the ductility of the nanocomposites (Table 12.2), a strong drop is found for those with PEG-GO as a nanofiller, the decrease being nearly 50% at 3.0 wt.% loading, whilst for the nanocomposite with 4.0 wt.% PEG-g-BNNTs it is only around 30%. This could be explained considering the strong PEG-GO/PPF H-bonding and polar interactions that considerably limit the ductile flow of the polymer segments, thus, resulting in lower εb values. With regard to the toughness of the nanocomposites measured as the area under the tensile curve (Table 12.2), different trends are observed for the two types of composites: those comprising of GO show a maximum in the range of 0.5
1.0 wt.% and then fall slightly, whereas for those filled with PEG-g-BNNTs, it increases steadily with increasing loading. For both cases, the maximum improvement compared to the value of the neat polymer is about 40%. This improvement is noteworthy, given that PPF is a brittle material and a high toughness is crucial for bone tissue engineering applications.

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Tissue engineering biomaterials often stay within the body for a long time, consequently their mechanical properties have to be assessed under physiological conditions. In particular, PPF-based samples were exposed to an SBF at 37 C and the data drawn from tensile tests under such an environment are also summarized in Table 12.2. Thoroughly, stiffness and strength data in air are higher than those obtained in an SBF medium, most likely related to the increase in hydrophilicity and biodegradation rate induced as the nanofiller content rises (Table 12.1); the higher the filler concentration, the more intense the plasticization effect of the absorbed water, which reduces the crosslinking density between the PPF chains and this is reflected in lower rigidity and strength. The differences between both conditions, for these two parameters, augment with increasing nanofiller content, as the plasticization effect (increase in the water uptake) becomes stronger (Table 12.1). The largest differences (around 56% and 38% in stiffness and strength, respectively) are found for the nanocomposite with 3.0 wt.% PEG-GO, which incorporates a large number of surface oxygen-containing groups. Nonetheless, the loss in properties found for these composites is lower than that reported for poly(propylene fumarate-co-caprolactone diol) reinforced with HA, which is likely caused by the higher hydrophilicity of HA compared to GO. Surprisingly, the elongation at break of PPF and the nanocomposites with small amounts of nanofillers is inferior in SBF compared to under air, whereas that of composites with filler concentrations equal to or higher than 2 wt.% are superior. This behavior could be attributed to the opposition of two facts: the shortening of the polymeric segments due to rupture of the ester bonds via hydrolytic degradation and the plasticizing effect due to the remaining humidity that increases the ductility. For small filler contents, the former factor predominates, whilst at loadings $ 2.0 wt.% the plasticizing effect prevails and the outcome is an augment in ductility. Even so, the toughness of both types of nanocomposites is thoroughly lower under SBF conditions, demonstrating that the loss in strength surpasses the raise in ductility for nanofiller loadings higher than 2.0 wt.%. The largest drops in ductility are found for the nanocomposites with 0.5 and 1.0 wt.% PEG-GO (around 40% and 33% respectively), whereas the smallest (about 17%) for the nanocomposites with 3.0 wt.% PEG-GO and 4.0 wt.% PEG-g-BNNTs. On the whole, for nanocomposites with either PEG-GO or PEG-g-BNNTs. the deterioration in mechanical performance owing to dipping in SBF is straightforwardly connected to the hydrophilicity and biodegradation characteristics of the nanocomposites, being the drop in properties higher as the level of hydrophilicity and weight loss due to hydrolytic degradation raise. More outstandingly, experimental data corroborate that biocomposites based on PPF maintain sufficient mechanical strength in a physiological environment to hold up new tissue formation. The temperature dependence of the storage modulus (E0 ) and tan δ (ratio of loss to storage modulus) for PPF and nanocomposites with different nanofiller contents has been investigated by dynamic mechanical analysis (DMA), a technique broadly used to characterize the mechanical properties of biomaterials as a function of temperature that offers information about their stiffness and damping,

12.6 Characterization of PPF-Based Bionanocomposites

a measure of how they can dissipate energy under a cyclic load. The results of E0 at 25 C and the glass transition temperature (Tg) obtained from DMA spectra of the nanocomposites are collected in Table 12.2 (Diez-Pascual and Diez-Vicente, 2016a; Diez-Pascual and Diez-Vicente, 2016b). E0 is indicative of the capability of a material to store mechanical energy without dissipation; the higher the E0 , the stiffer the material is. For both types of composites, E0 increases with increasing nanofiller loading, following a similar trend to the Youngs modulus. Thus, the largest increment (about 225%) is found for the nanocomposite with 3.0 wt.% PEG-GO. Interestingly, for nanocomposites with PEG-GO, the enhancement in modulus is more pronounced at temperatures above Tg (i.e., more than fourfold increase at 100 C for the nanocomposite with 3.0 wt.% loading, in agreement with the results reported for other polyester-based composites) (Diez-Pascual and Diez-Vicente, 2014), whereas for nanocomposites with PEG-g-BNNTs, the increase in E0 is larger below Tg, which points toward different stiffening mechanisms, however, the different behavior is not fully understood yet. With regard to Tg data for the different nanocomposites (Table 12.2), a regular augment is found on increasing nanofiller loading, demonstrating confined motion of the polymer chains in the presence of nanofillers. For instance, Tg increases by B14 C for the nanocomposites with the highest nanofiller contents in comparison to PPF. In general, the increments are slightly higher for composites with PEG-gBNNTs compared to those incorporating PEG-GO. Further, it was found that the height of the tan δ peak decreases upon increasing nanofiller loading, another indication of the restrained chain mobility in the nanocomposites compared to the neat polymer. Besides, the lower tan δ for the nanocomposites suggests that when the stress is removed, the energy accumulated in the material is recuperated earlier than in the pure polymer. In addition, a broadening of tan δ is found upon raising the nanofiller concentration, which has been explained as a larger volume of the interface (Diez-Pascual and Diez-Vicente, 2016a).

12.6.5 ANTIBACTERIAL PROPERTIES Microbial infection of biomaterials is an extensive problem in surgical treatment because it may lead to implant release, arthrodesis, and even death. The growth of microorganisms is influenced by numerous factors, like the biomaterial chemistry, surface physical properties, and the extension of surgical invasion, to mention but a few. Amongst the most common bacteria involved in the contamination of biomaterials are Gram-negative E. coli and P. aeruginosa as well as Grampositive S. aureus and S. epidermidis. Thus, the antibacterial characteristics of PPF-based nanocomposites were explored against the indicated bacteria. The nanocomposites were initially sterilized and submerged in a B2.0 3 106 colony forming units (CFU)/mL broth, followed by incubation for 24 hours at 37oC; % finally, the number of bacteria colonies was counted and the antibacterial activity was obtained as: log(viable cell countcontrol/viable cell countcomposite), where a

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Table 12.3 Antibacterial Activity of PPF-Based Nanocomposites Against the Indicated Bacteria and Percentage of Cell Viability Versus NHDF for an Incubation Period of 1 day Nanofiller (wt.%)
PEG-GO (0.1) PEG-GO (0.5) PEG-GO (1.0) PEG-GO (2.0) PEG-GO (3.0) PEG-g-BNNTs (0.1) PEG-g-BNNTs (0.5) PEG-g-BNNTs (1.0) PEG-g-BNNTs (2.0) PEG-g-BNNTs (4.0)

S. aureus

S. epidermidis

E. coli

P. aeruginosa

Cell Viability (%)

0.07 0.66 1.26 1.92 2.02 2.43 0.23

0.08 0.74 1.38 1.83 2.14 2.34 0.18

0.03 0.33 0.86 1.45 1.70 1.96 0.54

0.08 0.39 0.77 1.33 1.62 2.08 0.44

99.4 96.9 94.8 90.1 91.6 85.7 98.1

0.67

0.56

0.88

0.71

97.3

0.92

0.89

1.24

1.13

95.1

1.16

1.09

1.52

1.50

90.6

1.87

1.72

2.06

1.99

87.7

beaker with no sample was used as a control. The average values obtained are summarized in Table 12.3. According to the ISO 22196:2007 regulation, to attain effective antimicrobial action the antibacterial activity must be higher than 2. Neat PPF does not show any biocide activity against the bacteria examined. For composites reinforced with the two types of fillers, the antibacterial action grows as the filler content rises. The nanocomposite with PEG-GO (2.0 wt.%) is effective against S. aureus and S. epidermidis and that with 3.0 wt.% loading is effective against all the bacteria investigated, whilst the rest of the nanocomposites incorporating this type of nanofiller are not effectual. On the other hand, only the nanocomposite with PEG-g-BNNTs (4.0 wt.%) has antibacterial activity against E. coli and P. aeruginosa. Thus, composites filled with PEG-GO have superior activity compared to the equivalent PEG-g-BNNT. The different behavior can be explained by the large surface area of GO, which results in a large contact area with the bacteria. Surprisingly, nanocomposites incorporating PEG-GO exhibit stronger effect against Gram-positive cells while those including PEG-g-BNNTs show stronger effects against Gram-negative ones. This suggests the existence of diverse antibacterial mechanisms, although this has still not been clarified. In addition, negligible differences were reported between the activity against S. aureus and S. epidermidis or E. coli and P. aeruginosa, signifying that the toxicity differences are linked to the nature of the cell wall (Cabeen and Jacobs-Wagner, 2005) as

12.6 Characterization of PPF-Based Bionanocomposites

well as to their different sizes and shapes. In particular, S. aureus and S. epidermidis are minuscule round bacteria, while P. aureginosa and E. coli are rodshaped, the former being considerably longer. More information about the antibacterial action was attained by means of the agar-diffusion method, which ascertains the microorganism sensitivity against determined antimicrobial agents, being the bacteria more susceptible as the zone of inhibition area increases. For such tests, each bacteria was grown in a nutrient agar overnight and wells were placed in the medium; the nanocomposites were loaded into the wells, incubated for 1 day at 37oC, and then the inhibition zone was determined. Fig. 12.9 shows photographs % of the inhibition zone against E. coli and S. aureus for PPF/PEG-g-BNNT composites (Diez-Pascual and DiezVicente, 2016b). Third generation antibiotics (cefixime and cefoperazone) were used as blanks against E. coli and S. aureus, respectively. In agreement with the results from the colony-counting method, the nanocomposites had better antibacterial activity toward Gram-negative cells and the activity was reduced as the nanofiller loading increased. Both bacteria tested are resistant to neat PPF, since it does not exhibit inhibition zone. The inhibition zone diameter grows with nanofiller content; the largest (19 nm) being for the nanocomposite with 4.0 wt.% loading, which had a similar size to that of cefoperazone against S. aureus. These facts confirm the susceptibility of the bacteria tested to these types of nanocomposites, and corroborate their efficiency as antibacterial agents. The antibacterial activity of graphene-based materials is well documented, although to date, the reasons for their biocide action are not completely clear.

FIGURE 12.9 Inhibition zone of PPF/PEG-g-BNNT nanocomposites on E. coli (left) and S. aureus (right) measured via the agar-diffusion technique. The wells corresponding to each nanocomposite loading are: (A) 4.0 wt.%; (B) 1.0 wt.%; (C) 0 wt.% (positive control); (D) 0 wt.% (PPF); (E) 0.5 wt.%; (F) 2.0 wt.%. Reprinted from Diez-Pascual, A.M., Diez-Vicente, A.L., 2016b. PEGylated boron nitride nanotube-reinforced poly(propylene fumarate) nanocomposite biomaterials. RSC Adv. 6, 79507
79519, with permission from the Royal Society of Chemistry.

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Several mechanisms have been proposed, including oxidative stress via the production of reactive oxygen species (ROS), cell membrane harm caused by direct GO-bacteria contact, or the entrapment of bacteria within the graphene layers (Akhavan and Ghaderi, 2010). Few studies have provided proof of ROS production by GO in bacterial systems (Krishnamoorthy et al., 2012). GO flakes can generate hydroxyl radicals that harm the peptide bonds of the bacterial wall and damage the cellular components, resulting in bacteria disruption. Gram-negative bacteria are more resistant to membrane spoil owing to their outer membrane, while the Gram-positive ones that lack an outer membrane are weaker. On the other hand, scarce studies on the antibacterial action of BNNTs have been published so far, and probable mechanisms could be the formation of ROS, the ability to perform endocytosis, and membrane rupture caused by the interiorization of the BNNTs in the cell membrane (Horva´th et al., 2011). Certainly, BNNTs are instinctively attracted to lipid bilayers and can go through the cell membranes by means of a lipid-mediated inclusion mechanism similar to that reported for CNTs (Corredor et al., 2013). As a result, lipid membrane destruction can be regarded as the main motive for the antibacterial action of nanocomposites incorporating BNNTs, therefore, Gram-negative cells with a peptidoglycan layer are more vulnerable to injury provoked by these inorganic nanotubes.

12.6.6 CYTOTOXICITY For biomedical applications, to attain an optimal integration of a biomaterial into the body, cytocompatibility through in vitro studies is the first goal that has to be accomplished. The most typical tissues suitable for interaction with bone tissue scaffolds are normal human dermal fibroblasts (NHDF), thus, these were selected to evaluate the cytotoxicity of PPF and its nanocomposites and the results are summarized in Table 12.3. As expected, neat PPF, which is completely biodegradable, is not toxic toward NHDF and exhibits a cell viability value close to 99%, in accordance with previous studies that demonstrated the excellent biocompatibility of PPF with varied cells (Fisher et al., 2001). All nanocomposites based on PPF can be regarded as nontoxic, since they have cell viability data higher than 85%. Nevertheless, a decreasing trend in this parameter is found with increasing nanofiller content, and nanocomposites with PEG-GO (3.0 wt.%) and PEG-g-BNNTs (4.0 wt.%) show the biggest drops, around 14% and 11% lower than PPF respectively. Not many groups have explored the in vitro cytotoxicity of GO and BNNTs and opposing conclusions have been drawn (Horva´th et al., 2011; Wang et al., 2011) since the toxicity of nanomaterials depends on many factors, like dimension, concentration, defects, synthesis method, surface modification, and cell type. Some studies have suggested that these nanomaterials can bring cell injury by means of different mechanisms; the foremost being the introduction of oxidative stress and DNA breakage, which results in cell disruption. The toxicity of the nanomaterials described herein is both dose- and time-dependent: they are

12.6 Characterization of PPF-Based Bionanocomposites

nontoxic when used in small amounts. More outstandingly, after coating with a biocompatible polymer like PEG, their toxicity was significantly reduced, with lethal doses of about 100 and 50 mg L21 respectively (Ciofani et al., 2012). As a result, cell viability information points out that the functionalization of GO and BNNTs with PEG strongly decreases their cytotoxicity to human cells. These are important results, since the exceptional in vitro cytocompatibility of nanocomposites based on PPF makes them perfect candidates for biomedical applications.

12.6.7 TRIBOLOGICAL PROPERTIES Biomaterials should have good tribological properties in terms of high wear resistance and low friction coefficient when sliding against body tissues in order to prevent inflammation and implant loosening. Accordingly, it is crucial to evaluate the tribological properties prior to long-term application of a biomaterial. In this regard, the coefficient of friction (μ) and the specific wear rate (Wsp) of PPF/ PEG-g-BNNT nanocomposites were measured at 24 C with a 100Cr6 steel ball of 6 mm diameter applying a load of  1N at a rotation rate of 120 rpm, and the results are shown in Fig. 12.10. The μ of PPF decreases gradually upon increasing nanofiller concentration by about 45% for the nanocomposite with 4.0 wt.% loading. This trend was attributed to the improvement in stiffness and strength, the

FIGURE 12.10 Coefficient of friction (solid bars) and specific wear rate (dashed bars) of PPF/PEG-gBNNT nanocomposites in relation to nanofiller concentration. Reprinted from Diez-Pascual, A.M., Diez-Vicente, A.L., 2016b. PEGylated boron nitride nanotube-reinforced poly(propylene fumarate) nanocomposite biomaterials. RSC Adv. 6, 79507
79519, with permission from the Royal Society of Chemistry.

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presence of regularly dispersed hard inorganic nanotubes that can inhibit abrasion and wear of the matrix during sliding, the high thermal conductivity of the BNNTs that facilitates heat dissipation in sliding contact, and their high shear modulus that makes then act as solid lubricant additives (Diez-Pascual and DiezVicente, 2016b). Thus, hexagonal boron nitride has been reported to be a good lubricator that reduces the μ of polymeric systems (Kalay et al., 2015). An analogous trend is found for the wear rate, which shows an approximately twofold fall for the nanocomposite with 4.0 wt.% PEG-g-BNNT. This strong improvement in wear resistance is associated to the exceptional mechanical properties of this nanocomposite and its strong PPF/PEG-g-BNNTs adhesion through polar and hydrogen bonding interactions; facts that result in the formation of a thin and uniform transfer film and consequently, improved tribological properties. It has been reported (Pawlak et al., 2009) that inorganic nanotubes immersed in a polymer matrix can act as ball-bearing elements and roll instead of slide between the composite surface and the counterpart, therefore, reducing the shear stress, μ, and Wsp. This rolling mechanism, the formation of a protecting film, the mending effect, and/or the polishing effect could explain the improvement in the tribological properties found for PPF/PEG-g-BNNT nanocomposites, which are greatly attractive traits for their application as medical implants.

12.7 CONCLUSION AND FUTURE PERSPECTIVES In this chapter, the preparation and characterization of biodegradable PPF/PEGGO and PPF/PEG-g-BNNT nanocomposites for tissue engineering applications were described. The composites were fabricated using a combination of ultrasonication and heat curing, and their morphology, hydrophilicity, biodegradation, cytotoxicity, thermal, mechanical, tribological, and antibacterial properties have been examined. The nanocomposites showed superior water uptake, biodegradation rate, protein absorption capability, stiffness, strength, and thermal stability than those of PPF. They also maintained sufficient stiffness and strength in a physiological environment and displayed antimicrobial action against pathogenic bacteria in humans, such as Gram-positive S. aureus and S. epidermidis as well as Gram-negative P. aeruginosa and E. coli. Nonetheless, they did not cause toxicity to human dermal fibroblasts. Composites filled with PEG-GO displayed better properties than those reinforced with PEG-g-BNNTs, pointing out that the noncovalent functionalization approach is better for improving the properties of PPF than the covalent method is; this is likely due to the oxidative process used for anchoring the PEG fragments onto the BNNT surface, which causes defects on their sidewalls, since these defects have negative effects on composite properties. These novel biomaterials demonstrate immense prospective in the field of tissue engineering. Owing to their good biocompatibility and tailorable mechanical properties, they can be used as orthopedic tissues, cancellous bone defect repair,

References

and soft tissue replacement, including muscles, tendons, ligaments, nerves, fat, and blood vessels, to mention but a few. Nonetheless, additional research is required to completely take advantage of the potential technological applications of these nanocomposites. Accordingly, in depth studies on structure
property relationships as well as the manufacturing and characterization of both the nanofillers, especially BNNTs, and the corresponding bionanocomposites at a largescale are required prior to their use in commercial applications. The development of the optimum formulation for each nanofiller system to meet specific requirements and to reduce the cost of the production of bionanocomposites are also desired. In addition, there are some safety concerns about the use of these nanocomposites as biomaterials. Additional investigation is still necessary to fully ensure their nontoxicity as well as the ecological safety of their use. On the whole, these bionanocomposites represent a brilliant prospect for a broad range of applications in the biomedical field.

ACKNOWLEDGEMENT Dr. Ana Dı´ez-Pascual wishes to acknowledge the Spanish Ministry of Economy and Competitivity for a “Ramo´n y Cajal” Postdoctoral Fellowship (RYC-2012-11110) cofinanced by the EU.

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3924.

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Diblock and triblock copolymers of polylactide and polyglycolide

13

Divya Sharma, Lindsey Lipp, Sanjay Arora and Jagdish Singh Department of Pharmaceutical Sciences, College of Health Professions, North Dakota State University, Fargo, ND, United States

13.1 INTRODUCTION The term resorbable comes from the word resorb, which means to be absorbed again. In the context of drug delivery, resorbable means to be broken down and assimilated in the body. Diblock and triblock copolymers composed of polylactide (PLA) and polyglycolide (PGA) are bioresorbable, which means that the brokendown parts of these polymers will get absorbed or dissolved in the body. Block copolymers are a specific type of polymer that constitute different blocks or sections of polymerized monomers. A diblock copolymer is composed of two different chemical blocks, such as PLA-poly(D,L-lactic-co-glycolic acid) (PLGA) and a triblock copolymer is composed of three different chemical blocks where each block has at least one feature absent in the adjacent sections, such as PLA-PLGAPLA. The basic units (monomers) of PLA and PLGA are lactic acid and glycolic acid. PLA is generally synthesized by the ring opening polymerization of two monomers, lactic acid and the cyclic diester lactide using a metal catalyst (e.g., stannous octoate). The polymer PLA exists in an optically active form (L-PLA) which is semicrystalline in nature and an optically inactive racemic form (D, LPLA) which is an amorphous polymer because of irregularities in its polymer chain structure. D, L-PLA forms a more homogenous dispersion of drug within a polymer matrix and is, therefore, the preferred choice over L-PLA for controlled drug delivery systems. Polyglycolide or poly(glycolic acid) (PGA) is a polymer formed by the polycondensation of glycolic acid or most commonly by the ringopening polymerization of the cyclic diester of glycolic acid; glycolide. PGA is hydrolytically unstable and degrades rapidly by random hydrolysis and cellular enzymatic activity, owing to the ester linkage in the backbone, to form glycolic acid which is consumed by cells via the citric acid cycle. Expeditious degradation leading to low mechanical strength usually limits its application as a biomaterial. PLA as compared to glycolic acid is more hydrophobic due to the presence of an extra methyl group resulting in its resistance to hydrolysis and degradation. Consequently, to optimize its degradation rate and pattern, PLA is often Materials for Biomedical Engineering: Thermoset and Thermoplastic Polymers. DOI: https://doi.org/10.1016/B978-0-12-816874-5.00013-X © 2019 Elsevier Inc. All rights reserved.

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FIGURE 13.1 Schematic representation of diblock and triblock copolymers of different types.

copolymerized with other degradable polymers, such as PGA and polyethylene glycol (PEG), which are comparatively hydrophilic in nature. Additionally, PEG serves as a protective layer against the immune system (Laredj-Bourezg et al., 2015). Biodegradable copolymers of ABA and a BAB triblock were introduced by MacroMed, where A represents the hydrophobic polyester block (PLA or PLGA) and B represents the hydrophilic (PEG) block. Some examples of biodegradable triblock copolymers include PLGA-PEG-PLGA; PLA-PEG-PLA; and mPEGPLGA-mPEG (Fig. 13.1). Due to these hydrophobic and hydrophilic moieties, these polymers have the ability to form temperature sensitive polymeric micelles. These polymeric micelles resemble natural carriers (such as viruses and serum lipoproteins) owing to a hydrophilic shell allowing them to circulate in the blood stream for a longer period of time unharmed by the immune system with a small size that prevents their uptake by the reticuloendothelial system (RES), and a hydrophobic core enabling protective encapsulation of drugs/proteins/peptides (Bonacucina et al., 2011). The main objective of this chapter is to summarize the history, synthesis, and characterization of PLA and PGA based diblock and triblock copolymers along with the application of these copolymers as resorbable drug delivery systems.

13.1.1 HISTORY OF POLYLACTIDE PLA is synthesized using two main monomers; lactic acid and the cyclic diester, lactide. Lactic acid was discovered in 1780 by a Swedish chemist, Carl Wilhelm Scheele, who isolated it as an impure brown sirup from sour milk (Kompanje et al., 2007). By the early 1880s lactic acid was commercially produced in the United States, marking the first step toward the study of lactic acid polymers. In 1845, PLA was first synthesized by the condensation of lactic acid

13.1 Introduction

(Jimenez, 2015), and later using reversible polymerization of cyclic esters by heating lactic acid under vacuum (Carothers et al., 1932). Soon after that, PLA had started being used commercially as a fiber material for resorbable sutures (Auras et al., 2010). PLA produced by these methods was expensive and of low molecular weight. Breakthrough research by Cargill Inc. in the early 1990s made acquainted the production of high molecular weight PLA using a commercially viable lactide ring-opening polymerization reaction. The direct condensation route was an equilibrium reaction which made it difficult to remove traces of water at high conversion stages in order to drive the reaction to a higher molecular weight. Additionally, the polymerization method used by Cargill Inc. involved synthesizing both lactide and PLA in the melt thereby avoiding the use of costly and unfriendly solvents. Firstly, a low molecular weight PLA prepolymer was produced by the continuous condensation reaction of aqueous lactic acid, following which a high molecular weight PLA was synthesized using a tin-catalyzed ringopening lactide polymerization reaction. Unreacted lactide was recycled to the beginning of the process by vacuum distillation. The major advantage of this process was the selectivity of the intramolecular cyclization reaction to add from a mixture of lactide stereoisomers using tin catalysis involving coordinationinsertion mechanism with more than 90% conversion and extremely low rate of racemization (Gruber and O’Brien, 2005).

13.1.2 HISTORY OF POLYGLYCOLIDE PGA was one of the very first resorbable polymers to be investigated for use as a biomaterial. It is synthesized from glycolic acid, which is a colorless, odorless, hygroscopic, and crystalline solid with high water solubility. PGA has been known since 1954 as a biodegradable, tough fiber-forming polymer largely used for forming synthetic absorbable sutures (Dexon) of high strength and modulus as well as medical implants (Gilding and Reed, 1979). PGA is a highly crystalline polymer (45% 55%) with a glass transition temperature B35 C and a high melting point in the range 225 230 C. It is insoluble in water and most organic solvents. Fluorinated solvents, such as hexafluoroisopropanol and hexafluoroacetone sesquihydrate, are unique in their capability of dissolving PGA allowing for its spinning or molding into cast films (Schmitt et al., 1971). Several methods have been investigated for the synthesis of PGA. The polycondensation of glycolic acid is the simplest way of synthesizing PGA; involving the heating of glycolic acid at 175 185 C to distill off water followed by continued heating at reduced pressure for a few hours to obtain the low molecular weight byproduct, glycolide. A ring opening polymerization method to synthesize PGA was invented by heating pure glycolide under nitrogen atmosphere in the presence of antimony, zinc, or tin containing compounds as catalysts. At present, stannous octoate is the most commonly used catalyst for this reaction. The reaction is carried out at a temperature below the melting point of PGA and the reactants are allowed to react for about 30 minutes to obtain high molecular weight PGA (Lowe, 1954). In another

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study, solid state polycondensation of halogenoacetates (e.g., sodium chloroacetate) under nitrogen atmosphere in a round bottom flask have also been used effectively to synthesize high molecular weight PGA (Schwarz and Epple, 1999).

13.1.3 SYNTHESIS OF DIBLOCK AND TRIBLOCK COPOLYMERS OF POLYLACTIDE AND POLYGLYCOLIDE Ring-opening polymerization is the most widely used method for the synthesis of diblock and triblock copolymers of PLA, PLGA, and PEG, where the hydrophobic A block is covalently linked to the hydrophilic B block by an ester linkage (Bret et al., 2011). Various authors have used ring-opening polymerization to synthesize different polymers of varying copolymer compositions. The scheme of synthesis remains analogous. Diblock copolymer (PEO:D, L-PLA) synthesis can be achieved by taking equal molar ratios of PEO and D, L-lactide in a round bottom flask. The solvent toluene and a nitrogen atmosphere are used to obtain an anhydrous atmosphere for the reaction, while stannous octoate is used as a catalyst. The reaction is carried out under reflux to prepare the diblock copolymer. Diblock copolymers can be coupled to synthesize a triblock copolymer (PEO:D, L-PLA:IPDI:D,L-PLA:PEO) using isophorone diisocyanate (IPDI) (a coupling agent) dissolved in toluene and refluxed with the diblock copolymer. The obtained copolymer is purified using fractional precipitation from methylene chloride using diethyl ether (Singh et al., 2007). Similarly, mPEG-PLGA diblocks can be synthesized and coupled using IPDI to prepare an mPEG-PLGA-mPEG triblock copolymer (Tang and Singh, 2009). On the other hand, the synthesis of PLA-PEG-PLA triblock copolymers is a one-step process with no intermediate step for coupling. Briefly, calculated molar ratios of PEG (initiator) and D,L-lactide are taken, and D, L-lactide is charged into a three-necked flask containing predried PEG in anhydrous toluene under nitrogen atmosphere. Once all the reactants are in a molten state, stannous octoate is added as the catalyst and the reaction is carried out at 120 C for 12 hours to synthesize the triblock copolymer of the desired copolymer composition. The copolymer obtained by this method can be purified by dissolving the crude copolymer in ice cold water followed by precipitation by heating. This purification step is repeated 2 3 times to remove unreacted monomers and impurities. The final product is freeze-dried to remove the residual water (Al-Tahami et al., 2011; Oak, 2012). Using the mentioned ring-opening polymerization method, copolymers of different block lengths can be achieved by varying the feed ratio of the monomers and the initiator. Studies have suggested that a larger hydrophobic block leads to sustained degradation of the resorbable polymer matrix resulting in controlled delivery of the incorporated drug over a long period. The ratio of hydrophilic and hydrophobic block lengths also affects the aqueous solubility and sol gel transition temperature of the respective copolymer. In different articles, Singh et al. varied the block lengths of both the hydrophobic and hydrophilic blocks while

13.1 Introduction

conserving the water solubility of the polymer as well as its injectability at room temperature, sol gel transition ability, and stability of the gel at 37 C (Chen and Singh, 2005,a,b, 2008; Chen et al., 2005; Al-Tahami, 2007; Tang and Singh, 2009; Al-Tahami et al., 2011). A different method was employed by Wu et al., to synthesize an MPEG-bPLA diblock copolymer using monomers MPEG and LA and carrying out the copolymerization reaction in an oil bath at 140 C for 48 hours. The precipitate obtained was cooled to room temperature and purified by dissolving in anhydrous methylene chloride followed by precipitating out the copolymer using ethyl ether. The MPEG-b-PLA copolymer obtained was then dried under vacuum and used as a macroinitiator; owing to the hydroxyl groups present on its backbone which can initiate the ring opening polymerization of cyclic poly(ethyl ethylene phosphate) (EEP), to generate methoxypolyethylene glycol-poly(D,L-lactide)-poly(ethyl ethylene phosphate) (MPEG-b-PLA-b-PEEP) triblock copolymers, in an additional synthesis reaction (Wu et al., 2011). A transesterification reaction between PLA with PEGNH2 was also explored by a group of authors to synthesize a PLA-b-PEG copolymer (Ta¸skin et al., 2012). Powder form PLGA-PEG-PLGA triblock copolymers can be synthesized by first preparing PLGA in powder form using a direct melt polycondensation method. Then to synthesize the PLGA-PEG-PLGA triblock copolymer in powder form, different proportions of PLGA can be added to a fixed amount of PEG with stannous octoate as the catalyst and the mixture heated in a two necked round bottom flask under nitrogen atmosphere to obtain a crude brown colored product (Gajendiran et al., 2013). Another study demonstrates the synthesis of a PLA/ PEG triblock and multiblock copolymers using an acyl halide-terminated PLA (PLA-diCOCl) prepolymer and anhydrous pyridine (Zhao et al., 2012). Thermosensitive star shaped block copolymers have been investigated for their application in injectable copolymeric drug delivery systems (Lee et al., 2009). The copolymers constituted of fixed molecular weights of PEG, varied mole ratios of D,L-lactide to glycolide (PLGA block), and overall feed ratios of D,L-lactide, glycolide, and three- or four-arm PEG. The monomers are added into a round-bottom flask under nitrogen atmosphere where the three- or four-arm PEG acts as a multifunctional initiator and stannous chloride acts as a catalyst. Bulk ring-opening polymerization is performed and the copolymer product obtained is cooled to room temperature and precipitated for purification using diethyl ether several times (Lee et al., 2009). A three-step synthesis mechanism was also invented to synthesize three-arm star-shaped PLGA-mPEG (3sPLGA-mPEG) and four-arm star-shaped PLGA-mPEG (4sPLGA-mPEG) copolymers using an armfirst method (Zou et al., 2012). In this method, a linear chain hydroxyl-terminated PLGA-mPEG diblock copolymer (LPLGA-mPEG) is synthesized by bulk ringopening polymerization followed by the carboxylation of dried trimethylolpropane (TMP) or pentaerythritol (PTOL) to produce CTMP or CPTOL with three or four carboxyl acid terminal groups using excess amounts of succinic anhydride (SA). Finally, an esterification reaction of two reactive precursors, L-PLGA-mPEG and

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CTMP or CPTOL, is performed using 1,3-dicyclohexylcarbodiimide (DCC) as a dehydrating agent and 4-(dimethylamino) pyridine (DMAP) as a catalyst to obtain a three-arm or four-arm star-shaped PLGA-mPEG block copolymer.

13.1.4 CHARACTERIZATION OF COPOLYMERS OF POLYLACTIDE AND POLYGLYCOLIDE 13.1.4.1 Structural composition analysis The confirmation of the completion of the polymerization reaction of PLA-PEGPLA and PLGA-PEG-PLGA triblock copolymers along with the molecular structure can be determined using Fourier transform infrared (FTIR) spectrometer in the frequency range 4000 1000 cm21 in absorbance mode. The characteristic peak of the carboxylic acid of PLA from 1700 to 1725 cm21 disappears and a new peak appears in the region of 1730 1750 cm21 due to the newly formed ester groups in the FTIR spectra. The characteristic peak for isophorone diisocyanate, used as a coupling agent, at 2175 cm21 can be used as an indicator of the completion of the coupling reaction in PEO:D,L-PLA:IPDI:D,L-PLA:PEO type coupled diblock copolymers (Singh et al., 2007). The characteristic signals of the PEG ether band and PLA ester carbonyl band can also be seen at 1086 and 1755 cm21 respectively, in a such a diblock or triblock copolymer (Ta¸skin et al., 2012). Additionally, both proton (1H) and carbon (13C) nuclear magnetic resonance (NMR) are established techniques to determine the chemical structure and structural composition of PLA and PGA block copolymers. The analysis may be performed using an NMR spectrometer operating at 300 or 400 MHz. The copolymer is dissolved in an organic solvent, such as deuterated dimethyl sulfoxide (DMSOd6) or deuterated chloroform (CDCl3), with tetramethylsilane (TMS) signal as the internal reference standard. Resonances in the B5.2 5.0 ppm ( O CH) and B1.5 1.4 ppm (CH3) ranges belong to PLA blocks. The main chain methylene signals of PEG (in a PLA-PEG-PLA triblock copolymer) usually show in the 3.7 3.3 ppm range. The α-methylene protons (PLA-COO-CH2) and hydroxylated methine ( CH) protons of lactyl end units appear together in the range of B4.3 4.1 ppm. If the copolymerization did not take place effectively, carboxylated end units of lactyl and methine protons of free lactic acid appear in the B5.0 4.0 ppm range and 4.03 ppm respectively, in a 1H NMR spectrum (Rashkov et al., 1996). Similar characteristic peaks were reported by several other authors with slight variations complying with the change in the type of block copolymer and the monomers in the copolymer chain backbone (Jeong et al., 1999; Tang and Singh, 2009; Al-Tahami et al., 2011). The structural composition, graft ratio, and the number average molecular weight (Mn) of the polymers can then be calculated by analyzing the integrated signals corresponding to chemical groups CH and CH3 of LA and CH2 of EG in an 1H NMR spectrum (Ta¸skin et al., 2012; Oak, 2012). The spectrum of 13C NMR has also been used to confirm

13.1 Introduction

the presence of PEG and PLA blocks by the characteristic peak of the methylene ( CH2) group of the PEG block at B71 ppm and the carbonyl ( CQO), methine ( CH), and methyl ( CH3) groups of the PLA block at B170, B69.4, and B17 ppm respectively (Oak, 2012). Gel permeation chromatography (GPC) is used to further determine the number average molecular weight (Mn), weight average molecular weight (Mw), and the molecular weight distribution (polydispersity index, PDI) of the synthesized copolymer. Polystyrene standards are used for calibration and tetrahydrofuran is a popularly used carrier solvent for the GPC analysis of PLA/PEG copolymers (Zhao et al., 2012; Oak, 2012).

13.1.4.2 Aqueous solubility and injectability Aqueous solubility and injectability are two lucrative properties making the copolymers of PLA/PGA versatile for drug delivery use; the main advantage being avoidance of toxic organic solvents. The concentration at which the copolymers are soluble below the gelation temperature is called the functional concentration (Rathi et al., 1998). The copolymer dissolves in cold water due to the PEG blocks keeping the copolymer in solution and the hydrophobic PLGA/PLA segments forming associative crosslinks. This happens due to the hydrogen bonding between hydrophilic PEG blocks and water making the copolymer soluble. This effect is dominant at lower temperatures. The copolymer concentration in water can be varied to allow injectability at room temperature and also to tailor the drug release profile. As the temperature increases the hydrogen bonding gets weaker and the hydrophobic forces in the hydrophobic PLGA/PLA blocks get strengthened and become dominant. In this way, a change in temperature leads to the reversible sol to gel transition of an aqueous copolymer solution of copolymers of PLA/PGA/PLGA (Bret et al., 2011). The transition temperature can be varied by changing the hydrophobic and hydrophilic block lengths of these copolymers. For injectable drug delivery application, it is desired that the copolymer solution incorporating the desired therapeutic should be injectable (sol form) at room temperature and on administration at physiological temperature (37 C) transform into a stable gel depot at the injection site (Tang and Singh, 2009). The total molecular weight of PLA and PGA copolymers for optimum solubility and reversible thermogelation should lie between 3500 and 4100 Da for ABA-type and 4000 4600 Da for BAB-type copolymers. For both types, the average molecular weight of hydrophilic block B (preferably PEG) should fall between 600 and 2200, while the overall weight percentage of the hydrophobic block relative to the hydrophilic block should be preferably high, between 65% and 78%. In BAB-type copolymers it has been found that the copolymer composition (ratio of PLA/PEG) and the total molecular weight of the copolymer have a striking effect on the release profile, especially for hydrophilic drugs (Rathi et al., 1998).

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13.1.4.3 Phase transition Response to stimulus is an innate property of living systems. The ability to design a system to manifest this property has been a starting point to several sterling researches and inventions. Thermosensitive copolymers of PLA and PGA undergo reversible in situ sol-to-gel transition in response to temperature changes. The transition mechanism of such aqueous copolymeric solutions is related to the presence of both hydrophilic and hydrophobic parts in their structure and usually to a lower critical solution temperature (LCST) in the aqueous solution (Chen and Singh, 2008). As the temperature increases above the LCST the equilibrium shifts from unimers to spherical micelles. The copolymer water interactions become thermodynamically unfavorable in comparison to water water or copolymer copolymer interactions, leading to dehydration of the solvated copolymer chains and finally to a transition into gel state (micelle packing). On increasing the temperature above the upper critical transition temperature (UCST), the polymer precipitates (Oak, 2012). The transition from gel to sol is related to the shrinkage of the hydrophilic component’s corona in the micelles owing to the effect of temperature on its solubility and the interaction of its chains with the hydrophobic hard core (Jeong and Gutowska, 2002). In copolymers containing PEG, the PEG chains orient themselves to align, forming the outer hydrophilic shell of the micelles facing the external aqueous environment. This layer of PEG acts as a barrier by reducing interactions with foreign molecules resulting from steric and hydrated repulsion. This results in the increased stability and shelf life of such systems (Makadia and Siegel, 2011; Bret et al., 2011). It has also been noted that PEG-PLGA-PEG triblock copolymers in aqueous solutions show increased polymer-polymer interaction as compared to polymer-solvent interactions. In other words, it has also been suggested that with an increase in temperature the polymeric micelles grow by increasing their diameter and eventually aggregate, thus, driving the sol gel transition (Bonacucina et al., 2011).

13.1.4.4 Thermal properties Thermogravimetric and differential thermal analysis (TG-DTA) of copolymers provides useful data to assess the influence of the copolymer composition on the degradation behavior. A TGA analyzer instrument is used for thermal characterization. Gajendiran et al. (2013) quantitatively assessed the degradation behavior of a PLGA-PEG-PLGA triblock copolymer using a thermal characterization method. According to their study, PLGA-PEG-PLGA triblock copolymers degrade in two steps with the loss of PLGA at B300 C and PEG at B410 C. Various changes in the degradation patterns with changing copolymer compositions were also reported. In a different study differential scanning calorimeter (DSC) was used to investigate the thermal properties of PLA-PEO-PLA triblock copolymers. It was found that the melting temperature (Tm) of PLA/PEO copolymers was lower than that of PEG alone, which ultimately has significant

13.1 Introduction

implications on the glass transition temperature (Tg) and crystallization peak (Tc) of these copolymers (Rashkov et al., 1996).

13.1.4.5 Crystallization behavior The effect of different copolymer compositions and architectures on the thermal properties and crystal structures of block copolymers MPEG-b-PLLA; PLLA-bPEG-b-PLLA; and four-arm PEG-b-PLLA has been intensely investigated (Zhou et al., 2015). DSC was used to scrutinize the thermal properties of the copolymers. The instrument was calibrated with pure indium and experiments were performed under nitrogen flow. Heating, cooling, and second heating scans were recorded for different MPEG-PLLA block copolymer compositions and the crystallinity of PLLA and PEG were calculated. Alongside this, wide angle X-ray diffraction (WAXD) measurements were also taken to study the effect of the chain connectivity, composition, and architecture of these copolymers on the crystal structures of PLLA and PEG. It was reported that melting point and crystallinity were affected by increasing molecular weight, arm length, and number of arms of PLLA in the MPEG-PLLA and PEG-PLLA block copolymers. This probably happens due to the formation of PLLA crystallites which causes the internment of PEG, resulting in increased difficulty for PEG to be packed into the crystal lattice. This suggests that varying the architecture and molecular weights of PLA/PEG block copolymers alters the consequential properties and exploring those will open new doors for the application of branched PLA/PEG block copolymers for controlled drug delivery applications (Zhou et al., 2015). WAXD data were also shown to support the fact that the formation of crystalline hydrophobic domains in PLLA gels resulted in a higher stiffness while in racemic PLA resulted in the formation of easily degradable amorphous hydrophobic domains due to the stereo random structure (Sanabria-DeLong et al., 2006). Thus, changing a simple chemical parameter, that is, stereo-regularity, can help tailor PLA containing block copolymers, such as PLA-PEO-PLA, for controlled drug release. Overall, it has been well investigated and proved that PLA and PEG blocks in a diblock or triblock copolymer affect the crystallization behavior of each other (Yang et al., 2006), and the gradually the increasing confinement of PEG is dictated by the crystallization of the PLA block (Huang and Paul, 2007). Another study supported this fact by suggesting that the crystallizability of PEO blocks depends on their length and can be reduced by copolymerization with PLA blocks (Rashkov et al., 1996).

13.1.4.6 Biocompatibility, cytotoxicity, and biodegradability In general, a system containing a drug suspended in the aqueous solution of PLA/ PGA/PLGA based diblock and triblock copolymers causes minimum toxicity and mechanical irritation to the surrounding tissues due to their inherent mucomimetic property, pliability of the gel, and biodegradation into lactic acid, glycolic acid, and ethylene glycol, which naturally dissipate from the body (Rathi et al., 1998; Bonacucina et al., 2011). PEG forms the hydrophilic shell of these copolymeric

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micelles in aqueous media making them nonimmunogenic, biocompatible, and soluble in water. It has also been documented that PEG of a molecular weight below 30,000 is easily eliminated by the body (Rathi et al., 1998). A plethora of data are available from cytotoxicity and biocompatibility studies on PLA/PGA based copolymers. 3-(4,5-dimethylthiazol-2-yl)-2,5- diphenyltetrazolium bromide (MTT) assay is a popularly used method to test the biocompatibility of a system in vitro. The principle of this assay is based on the ability of mitochondrial succinate dehydrogenase present in living cells to reduce this MTT dye to water-insoluble purple formazan crystals. The formazan crystals are then dissolved using isopropanol and the absorbance is measured using a plate reader. MTT assay is useful for assessing the subtle toxicity of systems which may not kill cells rapidly (i.e., within 24 72 hours) but may affect the metabolic and other functions of the cells necessary to maintain viability. A high absorbance relates to a high viability of the cells and hence, the low cytotoxicity of the sample tested. The cytotoxicity testing of PLA and PGA constituting block copolymers has been reported with insignificant difference from the control (cells incubated with growth medium only) affirming their biocompatibility (Oak and Singh, 2012a; Chen and Singh, 2008; Tang and Singh, 2009; Al-Tahami et al., 2011; Zou et al., 2012). In vivo biocompatibility can be tested by injecting animal models with these copolymeric drug delivery systems and comparing the histology of the injection site skin tissue using hematoxylin-eosin (H and E) stain to test for inflammatory responses, Masson’s trichrome staining to examine the vascularization, and Gomori’s trichome stain to test for collagen deposition, after specific time intervals. The safety and in vivo biocompatibility of PLA/PEG based copolymeric systems have been published in various articles. The overall results indicate that up to a week following subcutaneous injection of this copolymeric system, the infiltration of neutrophils and macrophages to the injection site occurs, demonstrating clear incidence of an acute inflammatory response, which subsides to a milder chronic inflammatory response at 30 days post administration with the presence of a few inflammatory cells, and finally at about 90 days closely resembles the control, indicating restoration to normal tissue with no signs of necrosis and/or chronic inflammation (Ma et al., 2014; Chen and Singh, 2008; Oak and Singh, 2012a). Biocompatibility of PLA and its copolymers for orthopedic, ophthalmic, otologic, skin, central nervous system, pulmonary system, parotid glands, urinary tract, and cardiovascular applications has also been established with a good safety profile compared to conventionally used devices and implants (Ramot et al., 2015). The in vivo biocompatibility testing of star-shaped block copolymers has also been similarly performed (Zou et al., 2012). The histology after 30 days showed almost complete restoration to normal tissue and no significant tissue necrosis, hyperemia, edema, hemorrhaging, or muscle damage. A Masson’s staining experiment showed that vascularization took place after 15 days, suggesting that starshaped PLGA-mPEG copolymers supported vascular in-growth, and overall they have good biocompatibility.

13.2 Resorbable Thermosensitive Polymers

PLA/PEG constituting copolymers degrade by nonenzymatic hydrolysis of ester bonds to nontoxic products which are naturally eliminated by the body (Zhao et al., 2012; Oak, 2012). 1H NMR and GPC can be used to determine the reduction in molecular weight of these copolymers while undergoing hydrolytic degradation. Copolymer composition affects the degree of gel hydration affecting the degradation rate of the copolymer, which in turn affects the permeability coefficient of the incorporated drug through the gel matrix (Jeong et al., 2000; Ma et al., 2014). The in vitro degradation of block copolymer hydrogels happens similarly; by the hydrolysis of ester bonds accompanied by the erosion of the gel in PBS solution at physiological temperature (Zou et al., 2012).

13.2 RESORBABLE THERMOSENSITIVE POLYMERS Temperature sensitive or thermoresponsive polymers are the most widely studied type of stimuli-sensitive smart polymers for drug delivery owing to the ease and benefit of exploiting changes in their state in response to physiological temperature. Numerous research papers provide fortifying evidence that drug delivery using thermosensitive polymeric systems is progressing at a rapid rate. Poly(N-isopropylacrylamide) (poly-NIPAAM) was the first most extensively studied prototype of thermosensitive polymers (Tang and Singh, 2009). It was synthesized in the early 1950s by free radical polymerization of N-isopropylacrylamide (Schild, 1992). However, due to its toxicity as well as low mechanical strength, poly-NIPAAM did not succeed for drug delivery applications (Bae et al., 1987). Later, ABA type of nonionic triblock copolymers (poloxamers), containing poly(ethylene oxide) as the hydrophilic block B and poly(propylene oxide) as the hydrophobic block A (poly(ethylene oxide)-co-poly(propylene oxide)-copoly(ethylene oxide) (PEO PPO PEO) copolymer, also called Pluronics) received the US Food and Drug Administration (USFDA) approval as a pharmaceutical excipient, but could not progress further as a pharmaceutical drug delivery system due to the nonbiodegradability of the hydrophobic PPO block and related toxicity (Wasan et al., 2003). In 1997, MacroMed Inc. replaced the nonbiodegradable block in Pluronics with a biodegradable and biocompatible PLA block to develop PEO PLA PEO thermosensitive triblock copolymer (Jeong et al., 1997). In an effort to optimize the phase transition behavior, degradation pattern, and hence the drug release kinetics, several copolymer compositions have been investigated over the years with encouraging results. Broadly, a mixture of the biodegradable/biocompatible copolymer and drugs/ proteins/peptides can be prepared by simple mixing in aqueous copolymer solution below gelation temperature to form a partially dissolved (colloidal state dispersion, such as suspension or emulsion) or completely dissolved drug delivery system which could be injected parenterally, administered topically/transdermally,

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and/or inserted into a cavity (ocular, vaginal, transurethral, rectal, nasal, oral, or aural). On administration, the formulation would undergo thermal gelation at physiological temperature (typically body temperature would be above the transition temperature) forming a depot and entrapping the drug in the polymer matrix (Zentner and Shih, 2004). The release from these copolymeric delivery system follows two mechanisms acting simultaneously: diffusion of the incorporated drug and degradation of the polymer matrix. Mostly, the initial release is diffusion-controlled and the later stage is a combination of both with degradation being dominant (Jeong et al., 2000). Thermosensitive copolymers of PLA and PGA retain water equivalent to B10% of the total weight of the hydrogel, which allows for the swelling of the gel depot and a diffusion pathway for the incorporated drug molecules. Water retention has been observed to vary with the ratio of the hydrophilic/hydrophobic content in the copolymer (Al-Tahami, 2007; Oak, 2012). It has also been reported that during the erosion of the hydrogel matrix (in the later phase), there is a preferential loss of hydrophilic segments (PEG-rich) rendering the remaining gel matrix hydrophobic with reduced water retention and swelling resulting in decreased copolymer degradation, ultimately leading to reduced drug release (Bret et al., 2011). Additionally, the drug release profile can also be altered by varying the copolymer concentration (Oak, 2012). Concentrations between 10% and 30% w/w are most preferred for drug delivery as lower concentrations were found to transition to form a weak gel, and higher concentrations are too viscous to be injectable. Optimization is required to reach a balance between a strong gel network and the desired release rate (Rathi et al., 1998). A model hydrophilic drug (ketoprofen) and a model hydrophobic drug (spironolactone) were tested by Jeong et al. using a PEG-PLGA-PEG thermosensitive copolymer to assess the release model of such a copolymeric system (Jeong et al., 2000). A domain structure was assumed with the drugs partitioning between the hydrophilic shell domain and the hydrophobic core domain. Drug release from the hydrophilic shell can be explained by diffusion and that from the hydrophobic core by the modified Higuchi equation. Thermosensitive copolymers made of PLA and PGA have the advantages of being easy to manufacture, soluble in water, showing avoidance of toxic organic solvents, simple formulation, ease of administration, controlled release of the incorporated drug, and ability to adjust copolymer composition for controlling the release period by modifying the degradation rate, the permeability of the matrix, and hence the drug release profile. Drug delivery systems using thermosensitive diblock and triblock copolymers of PLA and PGA will be discussed in detail next.

13.2.1 THERMOSENSITIVE POLYMER-BASED DRUG DELIVERY SYSTEMS The use of amphiphilic block copolymers for drug delivery was first proposed in the early 1980s (Pratten et al., 1985). The innovation of using PLA/PGA/PLGA/

13.2 Resorbable Thermosensitive Polymers

PEG copolymers for drug delivery applications lies in the simplicity of using these copolymers to deliver a wide variety of drugs, hormones, as well as sensitive proteins and peptides with efficacy. Sustained delivery of various such therapeutics is highly desirable as conventional drug delivery methods are far from ideal. Frequent subcutaneous, intramuscular, or intravenous injections at short intervals, daily application of patches which adhere poorly and/or cause irritation, poor oral bioavailability, and short half-life after parenteral administration, confront the need for a better controlled delivery system without toxicity (Chen and Singh, 2005a). Thermosensitive copolymers of PLA and PGA have shown good results both in vitro and in vivo for a large number of such therapeutics, with some currently in the clinical testing phase discussed in Section 13.2.2. Succinic anhydride terminated diblock copolymer methoxypoly(ethylene glycol)-b-poly(lactide) (mPEG-PLA-SA) has been investigated to synthesize 7-Ethyl10-hydroxy camptothecin (SN38) drug conjugated polymeric micelles (Lu et al., 2016). SN38, an active metabolite of irinotecan is a potent topoisomerase I inhibitor. Its clinical applicability as an antineoplastic drug is limited by its hydrophobicity and the instability of the lactone ring in its structure at a physiological pH. Drug conjugates with amphiphilic diblock copolymers allow for the formation of polymeric micelles as drug carriers. Advantages of this micellar drug delivery system include passive accumulation of polymeric micelles in solid tumors via enhanced penetration and retention effect (EPR), increased therapeutic efficacy, reduced side effects, less frequent drug administrations, and improved patient compliance. The chain lengths of mPEG and PLA were shown to have a large effect on the particle size of the drug conjugated micelles as well as antitumor efficacy, both in vitro and in vivo. Polymer drug conjugate micelles were found to be less toxic and more efficacious drug delivery systems for cancer treatment. Enhanced controlled release properties were also depicted by this drug delivery system, which were mainly due to the shielding effect of the hydrophilic mPEG shell against plasma proteins, thereby reducing clearance via the mononuclear phagocyte system; while the hydrophobic core (due to PLA) showed the ability to incorporate hydrophobic drugs and allowing for their controlled release. Similarly, another study reported the solubilizing efficacy of typical amphiphilic block copolymers by studying the poorly water soluble drugs, paclitaxel and cyclosporin A (Rathi et al., 1998). In another study, the ocular pharmacokinetics of dexamethasone acetate was evaluated in rabbits using the microdialysis method with 20% w/w PLGA-PEGPLGA copolymer solution and compared to regular eye drops. A sevenfold higher maximum serum concentration (Cmax) and a 7.89-fold larger area under the curve (AUC) was obtained with the thermosensitive in situ gelling copolymer, thus, validating enhanced corneal permeability, prolonged precorneal retention, improved bioavailability, and higher drug efficacy (Gao et al., 2010). A group of researchers also suggested that various additives, such as sugars, surfactants, salts, amino acids, proteins, and other substances, can be readily incorporated in these block copolymers as and when required to modify the release characteristics and/or

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stability of the drug compound (Rathi et al., 1998). A long acting formulation of exendin-4 (exenatide, EXT) incorporated in a PLGA-PEG-PLGA triblock copolymer was tested with and without excipients (zinc acetate, PEG, and sucrose) to control the burst release of EXT, both in vitro and in vivo, with promising results (Li et al., 2013). EXT is an incretin-mimetic polypeptide established to enhance glucose-dependent insulin secretion for the treatment of type II diabetes. Due to the viscous environment of the gel, the hydrolytic instability of the polypeptide significantly decreased. The addition of excipients reduced the burst release with zinc acetate showing the best effect. In other studies, controlled release of insulin has been widely studied using this delivery system injected subcutaneously. Regel (PLGA-PEG-PLGA) polymer was used for the controlled delivery of recombinant human insulin for the basal requirement of insulin for up to 15 days (Kim et al., 2001). Meanwhile, Zentner et al. studied the release of paclitaxel, porcine growth hormone, glycosylated colony-stimulating factor, and recombinant hepatitis B surface antigen, using Regel polymer with positive influence on drug effectiveness and stability (Zentner et al., 2001). The controlled delivery of levonorgestrel, testosterone, and growth hormone has also been investigated from PLGA-PEGPLGA thermosensitive copolymer-based delivery systems. The effects of varying block lengths on release profiles were observed and conclusions were drawn for the effect of different drug types on the release profile and duration of drug release (Chen and Singh, 2005a,b, 2008). In order to simultaneously utilize the benefits of PLGA-PEG-PLGA, in terms of ease of formulation, localized administration, biodegradability, low systemic toxicity, and sustained drug delivery in combination with drug therapy, in an effort to improve the anticancer efficacy of drugs against osteosarcoma, a localized codelivery system of PLK1shRNA/PEI-Lys complexes and doxorubicin (DOX) suspended in a PLGA-PEG-PLGA thermosensitive hydrogel was developed (Ma et al., 2014). The delivery system allowed for the sustained codelivery of the incorporated drugs with no cytotoxicity, biocompatibility, and significant synergistic antitumor efficacy. Moreover, localized delivery to the tumor was beneficial in reducing systemic toxicity as observed by ex vivo histological analysis of major organs in Saos-2 xenograft models. Furthermore, the controlled delivery of proteins and peptides is a highly challenging effort owing to low half-life, implicit instability, and structural constraints. These are also some preeminent reasons that render basal level insulin delivery to type I diabetes patients a daunting task. Multiple frequent injections or round the clock insulin pumps are conventionally used currently in order to maintain normoglycemia. The delivery of sensitive proteins and peptides has been extensively studied using PLA/PLGA based triblock thermosensitive copolymers. PLGA-PEG-PLGA thermosensitive triblock copolymers showed a controlled release of different proteins for B2 weeks (Kim et al., 2001; Singh et al., 2007). This copolymer system demonstrated high burst release of hydrophilic drugs like insulin owing to the higher hydrophilic GA content in the copolymer backbone. PLA being more hydrophobic than PLGA was hypothesized to undergo a slower

13.2 Resorbable Thermosensitive Polymers

degradation owing to retarded hydration, swelling, and hydrolysis, and was further investigated for the controlled basal delivery of insulin. PLA-PEG-PLA triblock copolymers showed significantly lower burst release with desirable zeroorder release profiles over a period of 2 3 months (Al-Tahami et al., 2011). Later, by incorporation of chitosan zinc insulin complex into a PLA-PEG-PLA copolymer, a controlled basal insulin delivery of B84 days was obtained in vitro (Oak and Singh, 2012b). In an additional study, the biocompatibility of the delivery system and efficacy of the released insulin was successfully confirmed in vivo using a streptozotocin-induced diabetic rat model (Oak and Singh, 2012a). Simultaneously, an additional advantage observed with these block copolymers is the protection of the incorporated drugs from chemical degradation which is extremely helpful for easily chemically degraded drugs as well as sensitive protein and peptide-based drugs. Chen et al. (2005) studied the release profile of the model protein, lysozyme, using PLGA-PEG-PLGA thermosensitive copolymers of varying block lengths and aqueous copolymer concentrations. The controlled delivery of lysozyme was reported in a biologically active form, with significant lowering of burst release with increasing copolymer concentration (Chen et al., 2005). Similar studies were done with an mPEG-PLGA-mPEG copolymer, and the effect of extending the PLGA block resulting in decreased degradation and controlled release of the protein for a longer duration was reported (Tang and Singh, 2009). Controlled delivery of salmon calcitonin, a polypeptide hormone for the prevention and management of osteoporosis, was also investigated using an mPEG-PLGA-mPEG triblock copolymer in vitro in a female rat model. Calcitonin suspended in 40% w/v aqueous copolymer solution administered subcutaneously was seen to protect the rat from methylprednisolone acetate induced osteoporosis for up to 40 days (Tang and Singh, 2010). Representative examples of sustained release depot based drug delivery systems of PLA and PGA diblock and triblock copolymers are summarized in Table 13.1.

13.2.2 COMMERCIAL AND INVESTIGATIONAL EXAMPLES Long-term controlled delivery of hydrophilic and hydrophobic drug substances via the parenteral route is an attractive approach. PLA and PEG diblock and triblock copolymers are profoundly investigated for this purpose due to their myriad benefits. Ensuing a great deal of success in veterinary medication (Matschke et al., 2002), there is abundant appreciation of the potential for various applications of these copolymers. Substantial investigations are being carried out and several of them have made it to clinical trials. A sterile, lyophilized micellar formulation of paclitaxel Genexol-PM (Cynviloq) using a PLA-PEG diblock copolymer has been approved by the USFDA to be marketed in Europe and Korea (Pillai, 2014). In this copolymeric colloidal carrier, PEG served as a nonimmunogenic outer shell while the PLA in the hydrophobic core solubilized the hydrophobic drug. The maximum tolerated dose (MTD) of paclitaxel and its biodistribution in the liver, spleen, kidneys,

463

Table 13.1 Representative Examples of Depot-Based Drug Delivery Systems of Polylactide and Polyglycolide Diblock and Triblock Copolymers Copolymer

Drug or Active Ingredients

Major Effects

PLGA-PEGPLGA

Lysozyme

mPEG-PLA

7-Ethyl-10-hydroxy camptothecin

• • • •

PLGA-PEGPLGA

Cyclosporin, paclitaxel

PLGA-PEGPLGA

Dexamethasone acetate

PLGA-PEGPLGA

Exendin-4

PLGA-PEGPLGA PLGA-PEGPLGA (Regel)

Recombinant human insulin Paclitaxel, pGH, G-CSF, insulin, rHbsAg

• • • • • • • • • • • • •

PLGA-PEGPLGA

Levonorgestrel, testosterone, growth hormone

PLGA-PEGPLGA

PLK1shRNA/PEI-Lys complexes and doxorubicin

PLA-PEG-PLA

Insulin, zinc-insulin hexamers, chitosan-zinc-insulin complex

• • • • • • • •

mPEG-PLGAmPEG

Lysozyme, salmon calcitonin

• •

Increasing the PLGA block lengths of copolymers decreased initial burst release. Increasing copolymer concentration reduced the rate of drug release. Self-assembling micelles forming mPEG-PLA-SN38 conjugates were synthesized. Passive accumulation of polymeric micelles in solid tumors via enhanced penetration and retention effect was observed in vitro and in vivo. Reduced toxicity and increased anticancer efficacy of the system. Improved solubilization of poorly water-soluble drugs. Increased chemical stability. Enhanced corneal permeability and prolonged precorneal retention. Increased Cmax and AUC. Improved bioavailability, and higher drug efficacy. Increased stability. Possible addition of excipients reduced burst release. Controlled basal insulin release observed up to 15 days in vitro and in vivo after single s.c. injection. Reduced clearance of paclitaxel after direct intratumoral injection with minimal distribution into any organ. Controlled release of paclitaxel for B50 days. Controlled release of equivalent amount of pGH, insulin and G-CSF after single s.c. administration compared to daily i.v. conventional therapy. Regel/rHBsAg increased rHBsAg-specific antibody titers by 6 times compared to commercial vaccine Engerix-B. Increasing the hydrophobic PLGA block length of copolymers significantly decreased the release rate. Controlled zero-order in vitro release was observed. Enhanced absolute bioavailability of pGH compared to s.c. aqueous pGH solution. Synergistic antitumor efficacy of co-incorporated drugs. Reduced systemic toxicity owing to localized tumor delivery. Optimization of drug release rate by varying copolymer composition and aqueous copolymer concentration. Significantly lower burst release and controlled zero-order release profile of the system. Controlled basal insulin delivery in vitro and in vivo in chemically and structurally stable form. Controlled release of the protein for a longer duration by extending the PLGA block resulting from decreased degradation rate of the copolymer matrix. Protection of in vivo animal model from methylprednisolone acetate induced osteoporosis for up to 40 days.

References Chen et al. (2005) Lu et al. (2016)

Rathi et al. (1998) Gao et al. (2010)

Li et al. (2013) Kim et al. (2001) Zentner et al. (2001)

Chen and Singh (2005a,b, 2008)

Ma et al. (2014) Al-Tahami et al. (2011), Oak and Singh (2012a,b)

Tang and Singh (2009, 2010)

13.2 Resorbable Thermosensitive Polymers

lungs, heart, and in tumors were both found to be increased by two to threefold in preclinical studies. The antitumor efficacy was significantly improved compared to free paclitaxel. Clinical studies have demonstrated a better safety profile, higher efficacy, and better response rates of Genexol-PM in patients with metastatic breast cancer and advanced pancreatic cancer. Combination chemotherapy of Genexol-PM with cisplatin allowed for the administration of higher doses of paclitaxel, and showed significant results. Genexol-PM also increased the response rates for patients who were not responsive to conventional paclitaxel therapy. It is also considered a potentially effective treatment alternative for gemcitabineresistant pancreatic ductal adenocarcinoma based on promising in vivo data. Genexol-PM has completed phase I and phase II trials as a treatment strategy in metastatic breast cancer, nonsmall cell lung carcinoma, pancreatic cancer, ovarian cancer, and bladder cancer. Studies are underway for the treatment of several other diseases, as well as phase III and phase IV studies in recurring breast cancer patients (Jain et al., 2016). Additionally, a paclitaxel incorporated PLGA-PEG-PLGA triblock copolymeric formulation based on MacroMed’s proprietary ReGel technology, called OncoGel, was investigated for local tumor management. ReGel is a water soluble thermosensitive copolymer designed to undergo reversible phase transition from an injectable low viscosity solution incorporating a drug of choice (sol-state) between 2 C and 15 C, to a controlled release gel depot at physiological temperature (37 C) (Elstad and Fowers, 2009). Phase I clinical trials of OncoGel on patients with inoperable solid tumors showed mixed results. A paclitaxel dose of up to 2.0 mg/cm3 was well tolerated and the drug remained localized at the injection site. However, pain, injection site bruising, redness, irritation, muscle spasm, and postprocedural discharge were observed as major side effects. In another study, OncoGel demonstrated disappointing results in a phase IIb designed model to determine its impact on presurgical potential in patients with esophageal cancer (Elstad and Fowers, 2009; Jain et al., 2016). Alongside, a phase I/II dose escalation study for local injection of OncoGel in patients having recurring glioma was terminated within 8 weeks owing to dose-limiting toxicities with serious vascular adverse effects mainly subdural hematoma (Clinicaltrials.gov, 2017a,b).

13.2.3 LIMITATIONS OF THERMOSENSITIVE POLYMERS Thermosensitive copolymers are simple and elegant drug delivery systems which are easy to formulate and on administration at body temperature show instantaneous sol gel transition. Nonetheless, it is advised to keep the aqueous copolymer solution at 4 C to maintain good injectability and low viscosity of the system as the viscosity of the system generally tends to increase as the temperature approaches room temperature. This may be a possible hurdle that will need discretion while taking such drug delivery systems to a clinical setting (Gong et al., 2009; Vaishya et al., 2015). Alongside this, drugs with high water solubility and small size readily diffuse from these copolymer matrices because of the highly

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porous microstructure, low degree of crosslinking, and increased hydration and swelling of these copolymers. Studies have also reported that an initial burst release occurs owing to the drug being located close to the surface of these copolymeric delivery systems (Oak, 2012). Numerous studies have also reported irreversible protein aggregation in these delivery systems resulting in incomplete protein release in vitro and less than 100% bioavailability in vivo. Though, this issue can be resolved to some extent by decreasing the drug loading (Vaishya et al., 2015). Furthermore, the addition PEG has been shown to limit the encapsulation efficiency of various drugs and proteins, even when adopting the most pertinent formulation techniques. This effect is suspected to be an effect of steric interference and possible drug/protein polymer interactions (Bret et al., 2011).

13.3 RESORBABLE NANOPARTICLES Nanoparticles have become an extremely popular drug carrier over the years for several reasons. Polymeric nanoparticles offer a lot of potential variations in composition, which plays into their ability to be fine-tuned for a specific therapeutic or delivery target. Furthermore, different methods can be used for incorporating therapeutics to increase entrapment efficiency while maintaining the stability of the system. Generally, polymeric nanoparticles range in size from 10 to 100 nm and can vary structurally as core shell micelles, nanospheres of polymeric matrices, and polymeric shells surrounding aqueous cores as nanocapsules (Roney et al., 2005). Resorbable polymers, such as lactide and glycolide, are useful candidates when designing nanoparticles. Diblock and triblock copolymers based upon PLA and GLA have proven to be biocompatible. In addition, since the formation of nanoparticles relies heavily on hydrophobic forces, the ability to alter the hydrophobic and hydrophilic portions of the amphiphilic block copolymers allows for precise designing of nanoparticles with modifiable physicochemical characteristics, such as size, charge, and entrapment efficiency. This can be done by tailoring their composition either by changing the lactide to glycolide ratio and/or by changing the molecular weight of the hydrophilic or hydrophobic blocks. PEG, polymethacrylate, polyethyleneimine (PEI), and polyethylene oxide (PEO) are commonly used as the hydrophilic blocks of nanoparticles based upon PLA, PGA, and PLGA. (Park et al., 2005)

13.3.1 NANOPARTICLE PREPARATION AND CHARACTERIZATION TECHNIQUES The general polymer synthesis is the same as previously discussed. There are multiple methods that can be employed for nanoparticle formation. Factors such as the hydrophobicity and stability of the therapeutic to be entrapped are considered when deciding which preparation method to use. Each method has potential

13.3 Resorbable Nanoparticles

pitfalls that should be kept in mind, especially when optimizing size and/or entrapment efficiency. Additionally, surface modification is frequently used to target the delivery of nanoparticles to specific tissues or to further stabilize the nanoparticles. It should be noted that some preparation methods may facilitate surface modification better than others. Some of the common preparation methods include emulsification, phase separation, dialysis, and spray drying. In the single emulsion or emulsion/solvent evaporation method the copolymers are dissolved in an organic solvent along with the hydrophobic drug or therapeutic. The organic solvent solution is then emulsified in an aqueous solution often by the aid of a surfactant. Finally, the organic solvent is evaporated and nanoparticles can be isolated (Park et al., 2005). Like with single emulsion, the precipitation solvent diffusion technique produces nanoparticles using an organic solvent for the dissolution of copolymers and hydrophobic therapeutics, but in this method the organic solvent is miscible in water. Upon diffusion of the organic solvent into the aqueous phase, nanoparticles precipitate immediately and are recovered by evaporation of solvent or dialysis against water. A double emulsion technique can be employed by preparing a water in oil in water (W/O/W) system, in which a primary water in oil emulsion is added to an aqueous solution that often contains a surfactant, such as polyvinyl alcohol (PVA) or tween (Park et al., 2005). The primary emulsion may be used to solubilize hydrophilic therapeutics while the organic solvent solubilizes the copolymers. Upon homogenization, a W/O/W emulsion is achieved and subsequent stirring followed by filtration yields nanoparticles. Due to the amphiphilic nature of these block copolymers, self-assembling nanoparticles can be designed. As described previously, the hydrophobic blocks of a diblock or triblock copolymer in aqueous solution will isolate themselves from water to form a hydrophobic core, while the hydrophilic blocks will interact with water to form a hydrophilic shell of the nanoparticle (Fig. 13.2). Additionally, the spray-drying technique requires spraying the drug and copolymer solution or suspension in organic solvents into a stream of heated air. This technique produces particles with advantages, such as small size, reproducibility, and milder preparation conditions. However, product loss is a major disadvantage of this method and has been addressed by developing a double-nozzle to simultaneously spray a mannitol solution, which coats the particles and prevents aggregation (Park et al., 2005). The desired properties of nanoparticles require characterization techniques unique to this type of drug carrier. Some common nanoparticle characterization methods include determining the particle size, entrapment efficiency, and zeta potential. These characteristics are easily detected spectroscopically using light diffraction with a particle sizer, by analyzing the drug concentration of disrupted nanoparticles using HPLC or any other established analysis method for drug/protein and electrostatic interactions within a dispersion respectively (Kaszuba et al., 2010; Park et al., 2005).

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CHAPTER 13 Diblock and triblock copolymers of polylactide

FIGURE 13.2 Schematic representation of diblock and triblock (BAB and ABA type) copolymers showing their auto-arrangement to form micelles or nanoparticles in an aqueous system.

13.3.2 RESORBABLE NANOPARTICLES-BASED DRUG DELIVERY SYSTEMS Nanospheres of PLA and PLGA based diblock and multiblock copolymers have been intently studied for the effect of varying polymer composition and theoretical drug loading on particle size and encapsulation efficiency (Peracchia et al., 1997). Nanospheres were obtained using the emulsification/solvent evaporation technique using methylene chloride or a methylene chloride to chloroform ratio of 1:1 as the organic phase in which the copolymer and drug (lidocaine or prednisolone) were dissolved. The organic phase was subsequently poured into an aqueous phase of double distilled water followed by mixing and sonication to produce the desired emulsion. Nanospheres were formed as a precipitate and collected via centrifugation. Purification was performed by washing twice with water followed by lyophilization to obtain the final nanoparticle product. Regardless of drug, increasing the molecular weight of PEG from 5 to 12 to 20 kDa and the theoretical drug loading from 20% to 33% w/w, when prepared using a PEG-PLGA diblock copolymer, showed a slight change in drug loading and particle size. When investigating multiblock (PEG)3-PLA nanoparticles, the size of the particles increased with increasing chain length of PEG. In addition, the encapsulation efficiency was reduced in the multiblock nanoparticles when compared to the diblock nanoparticles. The release profile for the nanoparticles prepared from diblock copolymers showed a typical fast release initially followed by controlled

13.3 Resorbable Nanoparticles

release, in contrast to the nanoparticles prepared using multiblock copolymers, which showed reduced burst release and a drastically slow drug release rate overall. The authors also highlighted the additional influence of an external PEG coat on the characteristics of drug release (Peracchia et al., 1997). In another study, poly(ester-anhydride) copolymer nano- and microspheres were prepared using a double-emulsion/solvent evaporation technique and characterized with additional attention given to the surface modification potential of these spheres. Surfaces were labeled with cystamine, which was quantified by reducing cystamine with Ellman’s reagent (5,5-dithio-bis-(2-nitrobenzoic acid) (DTNB). Microspheres and nanospheres showed labeling of 0.20 35 and 0.6 100 μmol/g respectively. Furthermore, increased labeling was observed with increasing polyanhydride composition (Pfeifer et al., 2005). A pivotal study was done on targeted self-assembling nanoparticles, evaluating their ability to release drugs in a controlled manner to target prostate cancer cells while evading the immune system by utilizing stealth properties (Gu et al., 2008). The nanoparticles were designed using PLGA, PEG, and the A10 aptamer (Apt). Apt targets the nanoparticles to prostate cancer cells that display prostate-specific membrane antigen (PSMA) on their surface enabling the nanoparticles to bind specifically to these cells and, thus, be endocytosed. This study demonstrates how nanoparticles can be tuned to obtain certain enhanced characteristics. The PLGAb-PEG-b-Apt triblock copolymer was engineered to self-assemble in water by hydrophobic forces driving the PLGA to form a hydrophobic core, while the PEG formed the shell with Apt protruding into the aqueous environment. The incorporation of PLGA-PEG at varying ratios provides a means to alter the concentration of Apt, which optimizes the stealth ability conveyed by the PEG moieties to increase circulation time and decrease accumulation in the liver, while retaining the targeting abilities conveyed by the Apt. Additional optimization was achieved through the tunability of PEG with the size of the nanoparticles being directly proportional to the size of PEG used while being independent of PLGA molecular mass. However, when optimizing the release of docetaxel, the opposite effect was observed and raising the PLGA molecular mass, but not PEG, prolonged the release rate. Cell specific targeting was confirmed in vitro using LNCaP cells in comparison to PC3 cells. In vitro and in vivo optimization of increasing Apt density was found to be 5% for maximum tumor targeting, conversely, a reduction in Apt density results in a lower retention within the liver and should be considered to obtain maximum targeting and minimal retention. In another study of a diblock copolymer comprised of polymethacrylate and PLGA was used to deliver plasmid encoding interleukin-10 to prevent autoimmune disease type 1 diabetes (Basarkar and Singh, 2009). In this study, nanoparticles delivered unique payloads and harnessed the benefits of a polymer other than PEG. Specifically, Eudragit E100 was used, which is a copolymer of n-butyl methacrylate, 2-dimethylaminoethyl methacrylate, and methyl methacrylate (1:2:1) with an average molecular weight of 150 kDa. Double emulsion nanoparticle preparation resulted in PLGA-E100 nanoparticles with a zeta potential of

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58.5 6 2.1 and 42.7 6 1.1 mV before and after plasmid loading respectively, which immensely exceeds that of nanoparticles prepared with PLGA alone with a zeta potential of 4.2 6 0.7 and 1.9 6 0.2 mV before and after loading respectively. The size of the nanoparticles was B800 nm and unaffected by composition. To prepare the nanoparticles, phosphate buffered saline was added to dichloromethane containing copolymers and sonicated to produce a primary emulsion. An aqueous solution of the cationic surfactant cetyltrimethylammonium bromide (0.5% w/v) formed the secondary emulsion upon the addition of the primary emulsion followed by homogenization. Evaporation of dichloromethane, centrifugation, and washing allowed for the isolation and finally the lyophilization of the nanoparticles. Plasmid loading and cellular internalization was facilitated by the cationic polymer and an additional benefit of the E100 was proven to be its buffering ability as demonstrated by acid titration. The ability of the nanoparticle to provide buffering inside the endosome is crucial to the protection and endosomal escape of intact plasmid to provide maximum transfection and, hence, protein expression. The destabilization of endosomes due to the presence of E100 was confirmed via monitoring of the endosomal pH, and higher IL-10 levels of expression were observed. In vivo, mice given low-dose streptozotocin (STZ) treatment developed immune infiltration of the pancreas and while inflammation was reduced with passive treatment with IL-10 plasmid or IL-10 loaded PLGA nanoparticles, only the PLGA-E100 nanoparticles loaded with IL-10 showed the ability to completely protect the pancreas. Therefore, harnessing the cationic nature of E100 allowed for the development of nanoparticles capable of condensing plasmid DNA, cellular uptake, and endosomal escape to accomplish transfection and the expression of protein to protect against autoimmune inflammation of the pancreas. Pagar and Vavia (2013) provide another example of a unique nanoparticle design in which L-lactide is polymerized onto the cyclodepsipeptide, cyclo(GlcLeu) (Pagar and Vavia, 2013). An oil in water single emulsion was used to prepare the nanoparticles. The polymer and the drug, rivastigmine tartrate, were dissolved in methylene chloride and added to the aqueous phase containing 0.5% PVA while stirring. Homogenization and then sonication were employed to reduce the droplet size before evaporating the organic phase. The collected nanoparticles were lyophilized and the formulation variables were analyzed for optimization purposes. It was found that increasing the polymer concentration produces lager sized nanoparticles with increased entrapment efficiency. As can be expected, increasing the concentration of PVA in the aqueous phase reduces particle size, but with reduced entrapment efficiency. The authors hypothesized that an increase in entrapment efficiency should occur with an increase in the amount of drug. However, the results stated that increasing the drug amount increased the nanoparticle size, but decreased the entrapment efficiency. This information was utilized and an optimal drug-to-polymer ratio was achieved, which was found to be 1:5, in order to achieve maximum drug entrapment with the smallest particle size. Another parameter to consider is the organic-to-aqueous phase ratio.

13.3 Resorbable Nanoparticles

Table 13.2 Representative Examples of Resorbable Nanoparticle-Based Drug Delivery Systems of Polylactide and Polyglycolide Diblock and Triblock Copolymers Copolymer

Design Attributes and Key Findings

Reference

PEG5K-PLGA PEG12K-PLGA PEG20K-PLGA (PEG5K)3-PLGA (PEG12K)3-PLGA (PEG20K)3-PLGA PLA:PSA

• Minimal effect of varying polymer composition and theoretical drug loading on particle size and encapsulation efficiency. • Observed additional influence of an external PEG coat on the characteristics of drug release.

Peracchia et al. (1997)

• Increasing polyanhydride content allows for increased surface labeling. • Increasing Apt density to 5% was found to be optimum for maximum tumor targeting. • Reduction in Apt density results in a lower retention within the liver. • E100 produced increased IL-10 transfection and expression to protect the pancreas against autoimmune inflammation. • Optimization of drug to polymer ratio will allow for maximum drug loading and minimal size. • Organic: aqueous phase ratio, cryoprotectant used, and sonication time were also found to affect nanoparticle characteristics.

Pfeifer et al., (2005) Gu et al. (2008)

PLGA-b-PEG-b-Apt

PLGA-E100

L-Lactide-

depsipeptide

Basarkar and Singh (2009) Pagar and Vavia (2013)

The size of nanoparticles can be influenced via this parameter and increased particle size was seen with increased organic phase volume. Once again, the optimal ratio for small particle size and high entrapment was 1:5. Other parameters, such as cryoprotectant used and sonication time, can also have an effect on the characteristics of such nanoparticles. Armed with this knowledge, an optimized nanoparticle was produced with a particle size of 142.2 6 21.3 nm and an entrapment efficiency of 60.72% 6 3.72%. Representative examples of resorbable nanoparticle based drug delivery systems of PLA and PGA diblock and triblock copolymers along with their design attributes and key findings are summarized in Table 13.2.

13.3.3 COMMERCIAL AND INVESTIGATIONAL EXAMPLES The first polymeric nanoparticle, Genexol-PM, has reached phase II trials in the United States. Genexol PM is a self-assembling nanoparticle of mPEG-PLA entrapping paclitaxel for the treatment of nonsmall-cell lung cancer, metastatic breast cancer, and gynecologic cancer. Genexol-PM has a particle size of

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23.91 6 0.41 nm and a zeta potential of 28.1 6 3.1 mV. First-order controlled release kinetics were observed in vitro with B65% of entrapped paclitaxel released within 24 hours and 95% released at 48 hours. When examined in nonsmall-cell lung cancer patients, 46.5% were responsive with favorable antitumor activity, but a high frequency of toxicities was reported. Likewise, phase II trials for metastatic breast cancer showed promising results with an overall response rate of 58.5%, however, associated toxicities were frequent. Nevertheless, gynecologic cancer phase I trials are underway. The consensus surrounding this formulation is that the nanoparticles can prevent some exposure of paclitaxel to normal tissue and potentially increase the dosing schedule and duration in comparison to current treatment with paclitaxel. The lower toxicity in this formulation warrants continued research and clinical trials (Lee et al., 2008; Ahn et al., 2014; Werner et al., 2013; Clinicaltrials.gov, 2017a,b). Meanwhile, docetaxel is also being investigated in an mPEG-PLGA nanoparticle formulation, Docetaxel-PNP. Patients with advanced solid malignancies are participating in the determination of the MTD as well as an evaluation of safety and pharmacokinetic profile (Wang et al., 2014; Clinicaltrials.gov, 2017a,b).

13.3.4 LIMITATIONS OF RESORBABLE POLYMERIC NANOPARTICLES As with any delivery system, nanoparticles have some pitfalls that should be taken into consideration. The limitations given previously for PLA, PGA, and PLGA based copolymers can be apparent when used in nanoparticle formulations, but there are also limitations that are unique to nanoparticles. In general, the variations and optimizations that have been discussed are engineered to overcome the potential limitations of these drug carriers. The major limitations include, but are not limited to, obtaining the desired size, entrapment efficiency, stability, and pharmacokinetics. Nanoparticles of small, uniform sizes (,1000 nm, but preferably 50 300 nm) are desirable due to their ability to reach deep tissues via diffusion while evading the immune system. Stealth properties, like PEG coating, will also allow for prolonged circulation by preventing immune system recognition and removal. Entrapment efficiency is extremely important since the efficacy of many drugs are dependent on the amount of drug that reaches its target. Hydrophilic drugs and small molecules can be particularly troublesome in this aspect and researchers frequently inquest to determine whether nanoparticle formulations would be feasible. Stability can vary depending on the copolymers and surfactant used as well as their concentrations. Drug leaking and premature release are also major hurdles when developing nanoparticle formulations. Finally, pharmacokinetics goes hand in hand with stability, which can ultimately influence important factors of the release profile, such as burst release and the duration of sustained release (Olivier, 2005).

References

13.4 CONCLUSIONS AND FUTURE PERSPECTIVES PLA and PGA based diblock and triblock copolymers have been shown as excellent drug/protein delivery carriers for easy administration and controlled drug delivery. The biodegradability and biocompatibility of these copolymer systems has attracted a lot of attention over the years for their wide application optimized for the delivery of both small and large molecules with good safety profiles. The striking potential of these copolymeric delivery systems is the ability to be tailored in relation to the therapeutic incorporated, by increasing or decreasing the hydrophilic/hydrophobic ratio resulting in accelerated or decelerated degradation for shorter or longer duration of drug release and can be further exploited. For longer drug release periods synthesizing a polymer with high degrees of crystallinity can also be considered. Additionally, chemical alterations in the copolymer backbone can be explored for polyelectrolyte complex formation with charged drug molecules in a way that modifies the release pattern or enhances the stability of the delivery system. The incorporation of additives can be potentially tested with the delivery systems for their effect on drug delivery or for a combination therapy approach (Al-Tahami, 2007; Oak and Singh, 2012a). Overall, these copolymers can be formulated into carriers at multiple scales, such as depots, microspheres, nanoparticles, as well as implants. These systems have the ability to incorporate a wide range of therapeutics of diversified intrinsic characteristics for their controlled delivery in a chemically and structurally stable form over varying time periods with different possible routes of administration. (Bret et al., 2011). Further studies will be effective in making this delivery system an ideal approach to administer a large number of protein and peptide based drugs at a controlled rate for a longer duration.

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Characteristics of polymeric materials used in medicine

14

Ernesto David Davidson Hernandez1 and Jacobo Rafael Reyes-Romero2 1

Tecnicatura Superior Universitaria en Palenteologia, Universidad del Chubut, Rawson, Repu´blica Argentina 2Escuela Ba´sica, Facultad de Ingenierı´a, Universidad Central de Venezuela, Caracas, Venezuela

14.1 INTRODUCTION Biomaterials are substances or a combination of substances of natural or synthetic origin, designed to act interracially with biological systems in order to replace any tissue, organ, or function of the human body. Another way of defining a biomaterial is as a material of nonbiological origin that is used in the manufacture of devices that interact with biological systems and that are applied in different areas of medicine. The science of biomaterials is made up of fundamental pillars concerning branches of science, engineering, and medicine. More specifically, this consists mainly of biology, materials science, tissue engineering, and biomedicine. Fig. 14.1 presents a scheme of organizations of each of these disciplines. Biomaterials must possess certain characteristics in order to be used in the medical field. The porosity of these is important, since pore size and microstructure influence axonal growth, motility, morphology, and cell adhesion (Mata et al., 2009), as well as the space between the pores as it is involved in cell adhesion and the rate at which the cells propagate (Mitragotri and Lahann, 2009). Also is important to consider the degradation mechanism of polymers because a biomaterial has to replace the natural components of living tissue. This property influences cell migration, proliferation, differentiation, and even cellular morphology (Carvalho et al., 2013). In addition, the elasticity of a biomaterial should be appropriate according to the tissue in question, since the biomaterial should not to deform easily or lose its structure, which also involves in the organization of cells (Mitragotri and Lahann, 2009). Biocompatibility is a fundamental aspect because each organism reacts differently to an implant. This means that the response of the immune system to a

Materials for Biomedical Engineering: Thermoset and Thermoplastic Polymers. DOI: https://doi.org/10.1016/B978-0-12-816874-5.00014-1 © 2019 Elsevier Inc. All rights reserved.

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FIGURE 14.1 Main axes that make up the science of biomaterials.

foreign body or antigen changes from one organism to another. When a biomaterial is referred to as being biocompatible it is thought that the biomaterial does not cause adverse effects within an organism. This means, that it is an inert, nontoxic material and that it is accepted by the organism (Wang, 2013). Finally, the compartmentalization of some biomaterials is frequently used for the release of molecules in the tissue where they are implanted; generally used for the release of drugs or growth factors that aid in tissue recovery (Wang, 2013; Palakurthi et al., 2013). Historically, applications of biomaterials in the medical field date back to 1860 with the introduction of aseptic techniques, and some unique biomaterials used include steels and alloys, extending their applications from bone repairs to drug delivery systems (Collet Gonzalez, 2004). At first, the search for biomaterials was carried out in a purely empirical way. This then changed profoundly to the point that the science and engineering of biomaterials can now be defined as interdisciplinary activities (Abraham et al., 1998). In other words, the development of biomaterials requires specialists in the different branches of medicine, engineering, and pure sciences. This interdisciplinary challenge allows advances in biomedical sciences and tissue engineering, considering the use of a simple biomaterial, drug use, living cells, and hybrid biomaterials (consisting of drugs and living cells). Another important group of biomaterials is the so-called intelligent biomaterials, which respond to signals from the biological environment. All of them make up a series of devices for mass and daily use in hospitals, clinics, and other healthcare centers. The most commonly used are syringes, bandages,

14.1 Introduction

catheters, serum and blood bags, waste containers, as well as sophisticated pieces that are applied to promote tissue regeneration or organ replacement (Abraham et al., 1998). One of the first polymers used as a biomaterial was polymethyl methacrylate (PMMA) during World War II for the purpose of repairing the human cornea (Hernandez Martı´n, 2012). Dr. Sir Harold Ridley invented the first intraocular lens manufactured in acrylic, and performed the first implants in patients with cataracts (Duffo, 2011). Occasionally, due to a malfunction of the eye lens (cataracts), it is necessary to surgically remove and surgically implant an intraocular lens to correct vision. Well, PMMA is also used to make such an intraocular lens (Hernandez Martı´n, 2012). The use as an artificial cornea or keratoprosthesis of this biomaterial, commercially known as “Plexiglas,” was first discovered during World War II. Later, polyethylene was used as an alternative to metal catheters in the 1950s and 1960s, and then in the 1960s acrylic was used as bone cement in hip and knee replacements. From then on the use of polymers has seen tremendous growth in the field of biomedicine. For example, PMMA is mainly used as bone cement in hip and knee replacements, and UHMWPE of medical grade is the main component that forms the articulating surfaces of hip and knee joints (Ramakrishna et al., 2001). In addition, the medical grade UHMWPE is also used in the spine as a convex plate interchangeable in the type of cervical prosthesis prodisc C and as a bearing that fixes chromocobalt in the prosthesis of pourus coated motion (PCM) (Van Dijk et al., 2003; Phillips and Garfin, 2005). The most important properties of polymers that are used as biomaterials are: low density, high molecular weight and biocompatibility, nontoxicity, easy sterilization, excellent mechanical properties that support the application until the tissue is scarred, absorptivity, and slow degradation. These are suitable for use as: prostheses, joints, implants, equipment and surgical instruments, and in elements such as bone cements, membranes, and components of artificial organs in order to replace others materials. The use of natural substances is of great importance to protect the biomaterials used within living tissue from oxidation, specifically for medical grade UHMWPE used in orthopedics. The antioxidant mechanism is based on the annihilation of free radicals that circulate in the body, mainly radical peroxides. Investigations reported by many authors show the effectiveness of polyphenols from vitamin E and vitamin C in synthetic polymers, such as medical grade UHMWPE, to prevent oxidation cascades, such as occurs in humans (Oral et al., 2006; Davidson et al., 2010, 2011). Polyolefins, mainly medical grade UHMWPE, are the most important polymers used in orthopedics. Its main applications are: sliding surfaces for artificial joints. Polyethylene (PE) can undergo oxidation, especially gamma sterilization, which increases hydrophilicity, recrystallization, and makes the polymer more brittle (Maitz, 2015).

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Since 1962, medical grade UHMWPE has been a component of prostheses and implants of the hip and knee. The most important characteristics or properties of this material are: low density, high molecular weight (in the range of 2 6 million g mol21), nontoxicity, natural chemical composition, simple structure, low water absorption, excellent chemical and physics properties, and high resistance to ionizing radiation. It increases mainly in an inert atmosphere, which is an important property of polyethylene. Because of these and other properties, this polymer has been used on both an industrial and medical level. For example, in the space industry it is used as aerospace costume design due to the fact it has carbon and hydrogen in its structure (Stephens et al., 2005). Currently many researchers are trying to modify the structure and properties of medical grade UHMWPE in vitro in order to use this material in minimizing the intolerance of living tissue. Many studies report excellent results in UHMWPE samples as bearing artificial joints and joint prostheses components (Maitz, 2015; McKeen and Lawrence, 2014). However, the average lifetime of prostheses is near to 15 years. In other words, UHMWPE is susceptible to wear, exhibiting an aseptic loosening of the prosthesis or implant, which it is the main cause of failure in UHMWPE components used in orthopedics. Due to this situation many researchers have been attempting improve the properties of this material. Investigations have been reported that the combinated use of ionizing radiation, in both inert and air atmosphere, thermal treatments and use of natural substances result successful, mainly with the use of ionizing radiation in inert atmosphere (Davidson et al., 2010, 2011; Kim et al., 2006; Oral et al., 2006). On the contrary, is a well-known the fact that the radiation in air atmosphere results in oxiadtion and degradation processes. In addition, studies by Davidson et al. (2011) on UHMWPE medical grade samples irradiated in air and stored in simulated body fluid (SBF) showed a predominance of degradation processes. The radiation in atmosphere air lead to the oxidation process and subsequent chain breaking mechanism of UHMWPE samples due to peroxide radicals (ROO ) being formed and transformed into carbonyl groups. On the other hand, in an inert atmosphere a predominance of crosslinking mechanisms is observed, thus, forming long chains leading to a high molecular weight. Other studies have been reported that the combinated effects of ionizing radiation in inert atmosphere, thermal treatments of annealing and remelting andh vitamin E as transport media as storage substance improve the crosslinking and decreasing the chain break mechanism. The results showed an increase in crosslinking formation and an improvement in chemical and physical properties (Davidson et al., 2011; Oral et al., 2006; Kim et al., 2006). The thermal treatments as annealing and remelt led to a decrease the residual stress during the processing of UHMWPE, in order to avoid structural failure. The use of natural substances, such as vitamin E, vitamin C, aloe vera, etc., as a means of transport and storage reduces the oxidation processes after of irradiation. The mechanism is based on a neutralization reaction between the phenol and peroxide radicals (ROO∙) yielding hydroperoxide and phenyl radical, which are

14.2 Applications of Biomaterials

FIGURE 14.2 Mechanism of reaction for peroxide and fenoxi radical.

more stable and keep for a long time in polymers, thus, retarding the oxidation cascades (Oral et al., 2006). This mechanism is described in Fig. 14.2. The importance of crosslinking is that it optimizes the properties of UHMWPE. In other words, modifying the crystal structure of the polymer may improve its properties in vitro. The changes observed could be beneficial to reduce the problem of the wear of joints, bearings, and others components of prostheses and implants once placed in living tissue. In this chapter, the types of biopolymers, mainly medical grade UHMWPE, characteristics, applications, behavior in vitro, and behavior in living tissue are described. In addition, is analyzed the behavior of biopolymers and bioplastics and its relation with nanomatarials. Moreover, the mechanisms of failure in orthopedic prostheses inside living tissue are discussed. Finally, a forward projection in the use of these materials is given.

14.2 APPLICATIONS OF BIOMATERIALS Most synthetic polymers that have been used as biopolymers are thermoplastic. Thermosetting polymers are also used but to a lesser extent. It refers to materials that are used to produced syringes, serum or blood bags, hoses or flexibles tubes, adhesives, clips, elastic bands, sutures, bandages, functional imaging (positron emission tomography), etc. The materials used are of synthetic origin and are nonbiodegradable, such as polyethylene, mainly UHMWPE and HDPE, polypropylene, polyvinylchloride, polymethyl methacrylate, polycarbonate, polystyrene, nylon, etc. (Gonza´lez, 2004). For example, nylon and polypropylene are used as suture materials and PVC is used in tubes and bags for storing blood and pharmaceuticals, as well as in antibacterial membranes (Science in School, Polymer in Medicine, 2011). Moreover, thin films and coatings of PVC are use as storage bags and packaging surgical blood and others solutions; parts esophageal segment

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arteries, biodegradable sutures, implants joints parts fingers acetabular hip and knee, among others (Science in School, Polymer in Medicine, 2011). Polyethylene terephthalate (PET) is used for large diameter vascular grafts. Polymethyl methacrylate (PMMA) is used as a cement material for the femoral fixation of intraocular lens and for hip prosthesis. Polytetrafluorethylene (PTFE) is used in membranes for vascular periodontal ligament prosthesis as grafts. Various hydrogels are beginning to be used in applications such as: contact lens due to low protein adsorption and ease lubricity. Polyurethanes are an example of materials with excellent resistance to fatigue; and they are, therefore, used in artificial hearts (Blanco Rebeca Infante et al., 2015). PET is used in functional imaging (positron emission tomography) (Go´nzalez et al., 2002). Coating polymers (silicones, hydrogels, and fluorocarbons) are used for many cardiovascular applications. Bioreadsorbable materials are interesting because they are eliminated without further surgery. Generally, they are materials that degrade without being toxic to the body and are then eliminated. The most polymer materials commonly used, like bioreadsorbible biopolymers and hidroxy acids, are degraded to half of their mass in a few months (Blanco). Polyvinyl alcohol (PVA) is used in drug delivery systems; while polyacrylamide is used in medical diagnoses. Poly(lactic acid) (PLA) has become an indispensable material in the medical industry, where it has been used for 25 years due to its biodegradable and bioabsorbable properties (which means that it can be assimilated by the biological system). Its features and absorbability make PLA an ideal implant material for bone or tissue, orthopedic surgery, ophthalmology, orthodontics, controlled cancer drugs release, and for sutures (eye surgery, breast surgery, and abdomen) (Guerra et al., 2016). In addition, poly-L-lactic acid is a suitable biodegradable polymer for use as an implant material, mainly in screws for bone fractures, since it promotes bone regeneration, as confirmed by tests on rabbits and gross and histologic analyses of specimens studied. Its mechanical properties can be improved with the use of alcohols as coinitiators or by copolymerization with e-caprolactone (Zhang et al., 2013). Poly(glycolic acid) or polyglycolide (PGA: poly(hydroxyacetic)) is important both industrially and in the field of medical biodegradable linear polyesters. This area is focused on wound closure, such as surgical suture material, as well as bone fixation devices, such as rods, plates, or screws, and likewise on implants to replace bone fragments and on drug delivery systems. Natural biopolymers or macromolecules are extremely important substances for use in biomedical systems as they are commonly synthesized by living organisms. They are also important for use in new medical disciplines, such as tissue engineering, as biopolymers also include synthetic materials with the particularity of being biocompatible with living organisms (usually with human beings). The main families of natural biopolymers, are proteins (fibroins, globulins, amino acids), polysaccharides (cellulose, alginates, etc.), and nucleic acids (DNA, RNA, etc.); although others more unique, such as polyterpenes, among

14.2 Applications of Biomaterials

which natural rubber is included, as well as polyphenols (such as lignin) or some polyesters, such as polyhydroxyalkanoates produced by certain bacteria, due to it forming most of the Earth’s biomass. The polyhydroxyalkanoates is a polysaccharide. This biopolymers can be use in the diet of humans as dietary fiber, this works when is mixed with feces and this result give us digestion, defecation and prevents bad gases (Institute of medicines, 2005). Chitosan is a natural polysaccharide that is biodegradable, biocompatible, nontoxic, and has low immunogenicity so it is of great interest in the medical field for use as a biopolymer material. This substance is extracted from the shells of shrimps and prawns (Lo´pez Rubio, 2015). It is the second most abundant biopolymer available after of cellulose. In addition, for more the 20 years many researchers have published work on chitosan as a potential drug delivery system (Bernkop-Schnu¨rch and Sarah Du¨nnhaup, 2012). The chitosan in contrast with others polysaccharides having a monograph in a pharmacopeia, chitosan has a cationic character because of its a primary amino groups. These primary amino groups are responsible for properties such as controlled drug release, mucoadhesion, in situ gelation, transfection, permeation enhancement, and efflux pump inhibitory properties (Bernkop-Schnu¨rch and Sarah Du¨nnhaup, 2012). DNA is a polymer composed of monomers called nucleotides, which are transcribed in the cell in the form of a small chain of ribonucleic acid: messenger RNA. A study on women with advanced triple negative breast cancer found that those who received a poly (ADP-ribose) polymerase (PARP) inhibitor called iniparib along with chemotherapy had a longer survival rate than those who only received chemotherapy. Other small and preliminary studies showed some positive results using another PARP inhibitor, olaparib, in combination with chemotherapy in cases of triple negative breast cancer. The results were presented at the 2010 annual meeting of the European Society for Medical Oncology (ESMO) (Breasttcancer.org, 2010). Agarose is a thermoreversible natural biomaterial that is obtained from red algae (Jain and Bellamkonda, 2007). The term “thermoreversible” refers to a substance that can be converted from gel phase at room temperature to liquid phase at an elevated temperature above ambient temperature and which can be reconverted to gel phase when cooled to room temperature. The grievance has the same of a disaccharide composed of 3,6-anhydro-α-L-galactose and β-D-galactose (Wong et al., 2004). Another interesting and important group that is widely used both industrially and in medicine is referred to as “bioplastic.” These plastics are derived from plants, such as from soybean oil and corn or potato starch products, unlike conventional plastics which are derived from oil. Its origins go back to 1926, when scientists from the Pasteur Institute in Paris were able to produce polyester from the bacterium, Bacillus megaterium. The main characteristic of bioplastics that make them attractive for use is that they are based on natural resources and not on petroleum as in the case of synthetic polymers. This is of great interest because it reduces the energy consumption involved in their production and it eliminates the emission of gases that

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cause the greenhouse effect. In addition, producing plastics from biomass implies independence from petroleum. Major bioplastics are described here: Starch-based bioplastics are important not only because starch is the least expensive biopolymer, but because it can be processed by different methods used for synthetic polymers, like film extrusion and injection molding (Jogdand, 2014). Soybean have been revived, recalling Ford’s early efforts. In research laboratories it has been shown that soy protein, with and without cellulose extenders, can be processed with modern extrusion and injection molding methods (Jogdand, 2014). Many water soluble biopolymers, such as starch, gelatin, soy protein, and casein, form flexible films when properly plasticized. Although such films are regarded mainly as food coatings, it is recognized that they have the potential to be used as nonsupported stand-alone sheeting for food packaging and other purposes (Lo´pez Rubio, 2015). Starch is found in corn (maize), potatoes, wheat, tapioca (cassava), and some other plants. The annual world production of starch is well over 70 billion pounds, with much of it being used for nonfood purposes, like making paper, cardboard, textile sizing, and adhesives (Jogdand, 2014). Casein, commercially produced mainly from cow’s skimmed milk, is used in adhesives, binders, protective coatings, and other products. Soy protein and zein (from corn) are abundant plant proteins. They are used for making adhesives and coatings for paper and cardboard. Cellulose is the most abundant renewable material on Earth and is widely used in various industries, such as the paper and textile industries. Cellulose is formed by the union of glucose molecules by β-1,4 glycosidic bonds. In addition, it has a linear structure in which multiple hydrogen bonds are established between the OH group of the glucose chains, which play an important role in determining the strength and rigidity of the cellulose structural support (Bastioli, 2001). Cellulose represents 40% of the organic matter on the planet. Overall, bioplastics obtained from cellulose, either pure or in mixtures, are used in toys, sports equipment, medical applications, interior decoration, automobiles, and construction (Pacheco et al., 2014; Mutlu Hatice and Meier Michael, 2010). More than 150 different types of PHA have been studied, the analysis have been shown, the most representative being polyhydroxybutyrate (PHB), which accumulates in bacteria such as Alcanigenes eutrophus and Azotobacter vinelandii (Lakshman and Shamala, 2003). These polymers are used for the manufacture of cosmetic container products, in feminine hygiene products, utensils, packaging products, and bags (Valero et al., 2013). Collagen is the most abundant protein found in mammals. Gelatin is denatured collagen, and is used in sausage casings, capsules for drugs and vitamin preparations, and other miscellaneous industrial applications, including photography (Jogdand, 2014). Polyesters are now produced from natural resources, like starch and sugars, through large-scale fermentation processes and used to manufacture waterresistant bottles, eating utensils, and other products (Jogdand, 2014).

14.3 UHMWPE Behavior Under the Action of External Factors

Triglycerides have become the basis for a new family of sturdy composites. With glass fiber reinforcement they can be made into long-lasting, durable materials with applications in the production of agricultural equipment, the automotive industry, construction, and other areas. Fibers other than glass can also be used in this process, like fibers from jute, hemp, flax, wood, and even straw or hay. If straw could replace wood in composites that are currently used in the construction industry, it would provide a new use for an abundant, rapidly renewable, agricultural commodity and at the same time conserve less rapidly renewable wood fiber. Polyamide 11 is a polymer that although it comes from natural resources, is not biodegradable as are biopolyesters, such as bio-PET or bio-PE. Polyamide 11 or nylon 11 comes from the degradation of castor oil. Its properties include water resistance and high temperatures, which is reason it is used in electrical cables and in the automotive industry (Pacheco et al., 2014). Polyphenols are used as antioxidants to prevent premature aging of the body. The advantage of these biopolymers is that they are present in many foods and natural substances that are regularly used. It is present in substances such as vitamin E and aloe vera, and in beverages such as coffee, tea, beer, as well as in foods such as chocolate, nuts, olive oil, red wine, legumes, grains, etc. Biopolyethylene represent a type of renewable polyethylene obtained from the polymerization of bioethanol, possessing a similar structure to polyethylene, it is a nonbiodegradable compound, but to have the same characteristics as that obtained from oil becomes multipurpose material.

14.3 UHMWPE BEHAVIOR UNDER THE ACTION OF EXTERNAL FACTORS UHMWPE has previously been described here as a material that possesses excellent chemical and physical properties. In addition, it presents ideal characteristics as a biomaterial, and for this reason has been an important material used in orthopedics since 1962. Many hip and knee prosthesis components, such as joints, bearing, femoral heads, and friction pairs are made of medical grade UHMWPE. Fig. 14.3 shows a prototype of a hip prosthesis made with this material (Davidson, 2012). Presently in hip prostheses, UHMWPE is used instead of hydroxyapatite in order to form friction pairs with titanium and its alloys and/or tantalum. UHMWPE is located in the smooth part corresponding to the UHMWPE is in contact with the femoral head of Titanium in order to reduce the wear prostheses. Titanium is placed directly in contact with the bone of the femur because it is biocompatible and has a Young’s modulus of 110 GPa, while that of bone is between 20 and 30 GPa (Oldini, 2014). This means that the elastic modulus of titanium is five times higher than that of bone. In other words, this difference creates a high metal stiffness in the osseous fabric and so here, titanium, its alloys, and alloys of

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FIGURE 14.3 Hip prostheses prototype top view. Doctoral Thesis Davidson Ernesto, 2012.

other steels, generate the phenomenon of “stress shielding” (bone retraction because of lack of bone work). For this reason, other alloys containing aluminum and lighter metals are being investigated and evaluated (Oldini, 2014). In the case of knee prostheses, currently the most common options for knee replacement usually include metal alloys (stainless steel, Co Cr, and TI), UHMWPE, alumina (Al2O3), or Zirconia (ZrO2) (Plaza Torres and Aperador, 2015). However, the use of UHMWPE as a replacement for knee and hip prostheses is increasing rapidly each day, to the point that projections for 2030 are estimated to be 100% for patients over 65 years of age and 26% for patients less than 65 years (Kurtz, 2015). Due to the importance of UHMWPE and its increasing clinical use in orthopedics over more than five decades, many investigations using different systems have been reported. Studies show the use of ionizing radiation, mainly with gamma rays, in both inert and oxidative atmospheres and less frequently with neutrons, electrons, and heavy ions have been considered. Moreover, other works show the use of sterilization methods with peroxide and ethylene oxide, thermal treatments such as annealing and remelting, and the use of natural substances, mainly vitamin E, vitamin C, animal serum, and SBF. The majority of studies have shown encouraging results, mainly those involving ionizing radiation in an inert atmosphere, annealing and refining thermal treatments, and the use of natural substances for storage and transportation. There have also been reports of interesting results with the use of bovine serum, sterilization methods with peroxides and ethylene oxide, and the use of neutron and electron radiation. In an investigation, Davidson (2012) reported that ionizing radiation combined with gamma rays at different integral doses in an inert atmosphere, with storage in vitamin E, and annealing and remelting, showed improvements in the mechanical properties, wear resistance, and coefficient of friction of UHMWPE medical

14.3 UHMWPE Behavior Under the Action of External Factors

grade samples. In Fig. 14.4A and B the results of these analyses are shown. Davidson (2012) concluded that the combined effect of irradiation with gamma rays, storage and transport in vitamin E, and thermal annealing and melting treatments favor the mechanism of crosslinking. In other words, the combination of

FIGURE 14.4 (A) Wear versus integral dose for UHMWPE-Gur-1050 samples. (B) Coefficient of friction versus integral dose for UHMWPE-Gur-1050 samples. Doctoral Thesis Davidson, 2012.

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these external factors allows for the modification of the structure of UHMWPE in order to reduce the mechanisms that lead to the formation of oxidation cascades. This result shows for the first time the effectiveness of aloe vera as a storage substance in medical grade UHMWPE samples. The importance of this finding is that it will allow the use of a natural substance directly extracted from aloe vera. In other words, it is beneficial to reduce operational costs and generate a positive environmental impact, as it avoids the use of synthetic compounds in the design, development, and manufacture of prostheses. Moreover, the most relevant of these results is that it reduces the wear of the material to a large degree; it is of great interest for the design and manufacture of implants, prostheses, friction pairs, and tribological systems using UHMWPE, as wear is the most important factor that reduces the lifespan of prostheses. In other results, Davidson (2012) showed changes in the thermal properties of UHMWPE-Gur-1050 samples irradiated with gamma rays in an inert atmosphere using vitamin E as a storage substance and heat treatment for annealing (See Table 14.1). Davidson explains that the results obtained in the UHMWPE samples analyzed under the established conditions are due to the predominant mechanism of crosslinking. There the alpha tocopherol present in the vitamin E acts as neutralizing substance to the radical peroxides, thereby drastically reducing oxidation cascades. Moreover, the annealing prevents the formation of residual stresses, which favors the improvement of the mechanical properties (Davidson et al., 2011). Additionally, the decrease observed in the melting and crystallization temperatures and the decrease in the degree of crystallinity are caused by crosslinking, which leads to an increase in molecular weight and, therefore, longer and heavier

Table 14.1 Thermal Properties for UHMWPE-GUR-1050 Samples Properties Crystallinity Grade (%) ( 6 1) (kGy) 0 50 100 0 50 100 0 50 100

Control Sample 25 C 45

137

114

UHMWPE GUR 1050 120 C Vitamin E

UHMWPE GUR 1050 130 C Vitamin E

UHMWPE GUR 1050 140 C Vitamin E

UHMWPE GUR 1050 145 C Vitamin E

35 30 32 137 135 133 113 113 113

31 30 30 138 134 132 113 113 113

31 32 28 141 133 132 113 113 113

30 31 30 138 133 133 114 112 113

Source Davidson, E., 2012. Estudio y Análisis de las Propiedades del Polietileno de Ultra Alto Peso Molecular (PEUAPM-GUR-1050) de Grado Médico, para el Desarrollo y Fabricación de Prótesis Acetabulares a Nivel Nacional. Doctoral Thesis Doctoral.

14.3 UHMWPE Behavior Under the Action of External Factors

chains. There the presence of crosslinks and branches and chains of different sizes cause a decrease in thermal properties (Davidson et al., 2011). In addition, these long chains prevent molecular packing by preventing crystallization because they generate a smaller amount of crystals with more imperfections and as a consequence they produce a decreased degree of crystallinity. The application of thermal treatments inhibits the formation of crystals. Crosslinking, on the other hand, prevents the recrystallization of molten UHMWPE (Kim et al., 2006; Davidson et al., 2011). Therefore, the crystallinity in the irradiated samples stored in vitamin E and thermally treated is lower than in the virgin UHMWPE samples. This result agrees with those obtained in the analysis of the degree of crosslinking and that of the mechanical properties. From these results, the combined effects of irradiation with gamma rays, heat treatments of annealing, and storage in vitamin E, allow for the crystalline structure of UHMWPE to be modified, favoring the mechanisms of crosslinking and reducing the mechanisms of oxidation cascades. Likewise, the increase observed in mechanical properties, mainly in the wear resistance, is an important contribution to the development of UHMWPE prostheses and implants used in orthopedics. Other researches have reported successful and interesting results regarding the effects of natural substances, uses of ionizing and nonionizing radiation, and thermal treatments (external factors) on the structure and properties of medical grade UHMWPE. In their investigations they show how the changes in the structure of the materials modify their mechanical and thermal properties. In other words, their results show how the combined effects of these external factors modify the crystal structure of UHMWPE by avoiding oxidation cascades, thus, increasing the formation of crosslinking. Bracco and Oral (2011), studied the effect of vitamin E on medical grade UHMWPE samples for total joint implants The purpose in their review was to summarize preclinical research on the development and testing of vitamin Estabilized UHMWPEs for total joint implants. In their methodology, they conducted searches in PubMed, Scopus, and the Science Citation Index to review the development of vitamin E-stabilized UHMWPEs and their feasibility as clinical implants. They compared the behavior of UHMWPE irradiated and stabilized with vitamin E, irradiated UHMWPE, and UHMWPE fused after irradiation with gamma rays. They observed that UHMWPE samples stabilized with vitamin E, showing oxidation resistance and superior mechanical properties; nevertheless they had values of resistance equivalent to those of wear. They concluded that vitamin E-stabilized UHMWPE offers a joint arthroplasty technology with good mechanical, wear, and oxidation properties (Puertolas et al., 2010). Vitamin E (or a-tocopherol) is an alternative to thermal treatments to achieve the oxidative stability of gamma or electron beam irradiated UHMWPE used in total joint replacements. Their purpose was to study the effects of vitamin E on the molecular dynamics and microstructural properties of UHMWPE samples. They started with the hypothesis that the antioxidant would plasticize the UHMWPE. Vitamin E was incorporated

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into UHMWPE at different concentrations by diffusion and blending, and detected by ultraviolet and infrared spectroscopies at 500 and 4000 ppm respectively. A dynamic mechanical thermal analysis was used to characterize the influence of this antioxidant on the relaxations of the raw material. DSC and TEM served to characterize thermal and microstructure properties respectively. They obtained the results that vitamin E concentrations above 3% (by weight) significantly reduced the degree of crystallinity and increased the melting transition temperature of raw UHMWPE. Additionally, an increase in the concentration of a-tocopherol introduced and/or strengthened the beta relaxation, which was also gradually shifted toward lower temperatures and had rising energies up to 188 kJ mol21. Moreover, the gamma relaxation remained unaltered upon the addition of vitamin E. Finally, they concluded that no plasticizing effects of vitamin E on the molecular dynamics of UHMWPE could be confirmed from mechanical spectroscopy data. However, the relaxation was modified in intensity and location due to the changes in the degree of crystallinity introduced by the incorporation of vitamin E. Rı´os et al. (2013) studied the behavior in microstructure, oxidation behavior, and mechanical properties of UHMWPE-Gur-1050 irradiated with gamma rays and annealed with respect to post annealed material. Changes in the thermal transitions, degree of crystallinity, and thickness of the crystals were analyzed by differential scanning calorimetry (DSC). Additionally, they used transmission electron microscopy to analyze the crystalline morphology. They finally adopted a TGA technique to evaluate the resistance to thermooxidation and the induced changes in crosslinking by the effect of irradiation with gamma rays. The different results obtained show that sequentially crosslinked UHMWPE exhibited improved thermooxidation resistance and thermal stability compared to UHMWPE annealed after irradiation. In addition, the mechanical behavior, including fatigue and fracture toughness, of these materials were generally comparable regardless of the annealing strategy. Therefore, the sequential irradiation and annealing process could provide increased oxidation resistance, but not a significant improvement in mechanical properties compared to that of a single radiation dose and the subsequent annealing procedure. They concluded that the annealing treatments improved the oxidation resistance compared to the results obtained post-irradiation for the UHMWPE-Gur-1050 samples analyzed. Microstructural characterization shows that both the crosslinked and pristine UHMWPE have the same crystal thickness and degree and crystallinity. While the thermogravimetric behavior confirms that the crosslinked UHMWPE has the highest crosslinking density.

14.4 BEHAVIOR OF MEDICAL GRADE UHMWPE IN LIVING TISSUE Medical grade UHMWPE` for use in living tissue is an excellent material due to is bioinert property, high molecular weight (between 2 and 6 million), low weight,

14.4 Behavior of Medical Grade UHMWPE in Living Tissue

easy sterilization, resistance to ionizing radiation, simply structure, mechanical and chemical strength. These properties have led to polyethylene being the material of choice for use in prosthetics and joint replacement. Normally prostheses and replaced components have an average lifespan of 15 20 years in living tissue. However, the medical grade UHMWPE once placed in the living tissue, is present failed for aseptic loosen. This mean that appearance of debris (microparticles due to wear of material by friction with the titanium) lead a failed for asptic loosen in a period from 15 to 20 years. This is caused by two mechanisms: osteolysis and debris. Fig. 14.5 shows both osteolysis and debris mechanisms. Osteolysis or Gorham’s disease is when the bones wear as result of an inflammatory response or a lack of growth of bone cells by obstruction caused by prostheses or implants. On the contrary, detritus involves the formation of microparticles of polyethylene around tissue causing postoperative problems. Both mechanisms cause the loss of longevity of prostheses and components. Osteolysis is a rare disease characterized by bone destruction and resorption. Of unknown pathogenic mechanism, it causes anatomical alterations and leaves variable functional sequelae that depend on the location and intensity of lesions (Ferna´ndez Tejada et al., 2015). The types of osteolysis are idiopathic and secondary. Idiophaticy osteolysis is due to by reasorption of bone. The forms of idiopathic osteolysis are very rare diseases, characterized by destruction and bone resorption. The bones, apparently normal at the beginning, undergo progressive

FIGURE 14.5 Wear mechanism for UHMWPE-Gur-1050 samples: (A and B) osteolysis; (C and D) debris.

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destruction until they disappear partially or in their entirety. The secondary osteolysis is due to a variety of diseases in the bone tissue. It is a heterogeneous group of diseases that can affect very different bones of the organism. The final result, depending on the location and intensity of the affected bones, shows a wide range of variability, from patients with intense damage to others with few functional repercussions (Ferna´ndez Tejada et al., 2015). The mechanism of osteolysis has a negative impact on prostheses and implants. The appearance of this mechanism as well as debris leads directly to the failure of prosthesis as it is the cause of aseptic loosening, which reduces the useful lifespan of the prosthesis. According to Gallo et al. (2013), the process of aseptic loosening is initially governed by factors such as implant/limb alignment, device fixation quality, and muscle coordination/strength. Later large numbers of wear particles detached from total knee arthroplasties (TKAs) and trigger and perpetuate particle disease, as highlighted by the progressive growth of inflammatory/granulomatous tissue around the joint cavity (Gallo et al., 2013). Aseptic loosening secondary to periprosthetic osteolysis is the leading cause of revision procedures, especially in patients with a hip or knee total joint replacement (TJR), in the long-term. In the context of joint replacements, osteolysis refers to bone destruction as seen on conventional radiographs and corresponds to bone defects seen during revision surgery (Nich et al., 2014). In a large series of total hip arthoplasties (THAs), Lu¨bbeke et al. (2011) reported femoral osteolysis in up to 24% of cases in the decade following the procedure, with more active patients at increased risk of developing osteolytic lesions. In other investigations of TKA, osteolysis has been found in 5% 20% of cases at follow-up times ranging from less than 5 15 years (Fehring et al., 2004). As a result, up to 15% of patients are likely to be revised for aseptic loosening in the decade following a total joint arthroplasty (TJA). Although arthroplasty may be successful because of the availability of biocompatible materials. However, its half-life is diminished by inadequate fixation, mechanical loss over time, or biological loss due to osteolysis, the latter being produced as a tissue response to the wear particles of the implanted material located in the bone prosthesis interface (Astillero, 2012). In another study by Schwarz (2016), aseptic loosening in TJRs were observed due to periprosthetic osteolysis. Investigations over the past two decades have elucidated a central mechanism for osteolysis, in which wear debris generated from implants stimulate inflammatory cells, thus, promoting osteoclastogenesis and bone resorption (Schwarz, 2016). Total hip arthroplasty relieves chronic pain and improves movement in millions of patients in the advanced stages of osteoarthritis or arthritis. There arthroplasty is successful because of the availability of biocompatible materials (Solis Astillero, 2012). At present, the prescription of antiinflammatory and bone-resorbing suppressing agents is reported to inhibit osteolysis caused by bone cement. Among the former are cyclooxygenase 2 (cox-2) inhibitors, which plays an important role in wear debris-induced osteolysis. Studies conducted by Zhang et al. (2001) show the effectiveness of the drug as applied in mice.

14.5 UHMWPE Versus Other Biomaterials

Other researchers have reported that bisphosphonates, a class of molecules which inhibit bone resorption, showed an inhibitory effect on osteolysis-induced particles in vitro and in animal models (Trevisan et al., 2013). The debris mechanism is due to the fact of wear in implants, prostheses, joint, or bearing of medical grade UHMWPE uses in ortophaedia. The friction between the metal and the polyethylene causes wear and generates the formation of microparticles. Debris wear is a serious clinical problem because it generates aseptic loosening of prostheses and implants and activates the osteolysis mechanism. UHMWPE debris particles produced in hip implant wear simulation tests are classified as round debris, flake-like debris, and stick debris, which are closely related to the primary mechanisms of abrasive wear, adhesive wear, and fatigue wear (Wang, 2013). Many investigations have reported on the behavior of UHMWPE in living tissue (Kurtz, 2015). Macdonald et al. (2013) studied the damage mechanisms and oxidative stability of remelted UHMWPE. Remelted, highly crosslinked polyethylene (HXLPE) has been introduced into total knee replacements (TKR) since 2001 to reduce wear and particle-induced lysis. They observed that remelted HXLPE inserts had lower oxidation indices compared to conventional inserts. They were able to detect slight regional differences within the HXLPE cohort, specifically at the bearing surface. In conclusion, remelted HXLPE was effective at reducing oxidation in comparison to gamma inert sterilized controls. Moreover, long-term HXLPE retrievals are necessary to ascertain the long term in vivo stability of these materials in TKRs.

14.5 UHMWPE VERSUS OTHER BIOMATERIALS Medical grade UHMWPE has qualities that have led it to be selected as the material of choice for use in orthopedics in relation to other biomaterials, such as: ceramics, mainly hydroxyapatite and oxidized zirconium oxide, metals such as cobalt, chrome, titanium, and even polymers. Many investigations have been performed by several authors. Studies show the high effectiveness of both conventional and highly crosslinked UHMWPE in orthopedic components, such as bearings, joints, and implants. First-generation HXLPE tibial inserts became commercially available for TKA in 2001 (Kurtz, 2009). Since polyethylene wear is a major cause of osteolysis and related complications, it was assumed that HXLPE inserts would reduce wear, as has been observed in THA patients. The success of highly crosslinked HXLPE inserts in THA surgery was shown relatively soon after its introduction onto the market in the late 1990s. Several clinical studies of THA inserts showed the benefit of this HXLPE, namely the reduced incidence of osteolysis and reduced wear compared to conventional polyethylene (CPE). Studies performed by Berry et al. (2014) show that wear and corrosion are the most important and major causes of failure in joint arthroplasty. From the point

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of view of its current clinical importance, they mention the main four categories of wear and tribocorrosion: wear of polyethylene, wear of ceramic ceramic (CoC) bearings, wear of metal-to-metal (MoM) bearings, and tapered tribocorrosion. Problems with knee wear have become less prominent as they have many problems with hip PE bearings that result from the success of crosslinked UHMWPE. However, MoM and conical tribocorrosion joints have been associated with soft tissue inflammatory responses and, as a result, have become prominent clinical concerns. An investigation by Inacio et al. (2013) compared the short-term revision risk in alternative surface bearing knees (oxidized zirconium (OZ) femoral implants and HXLPE inserts) with that of traditional bearings (cobalt chromium (CoCr) on CPE). They analyzed 62,177 primary TKA cases registered in a total joint replacement registry from April, 2001 to December, 2010. The final steps for the analysis were all-cause revisions, septic revisions, or aseptic revisions. Bearing surfaces were categorized as OZ-CPE, CoCr-HXLPE, or CoCr-CPE. HXLPE inserts were stratified according to brand name. The results showed that in all review processes, both aseptic and nonaseptic, the risk of early revision does not show statistically significant damage. Another relevant outcome is that no specific brand of HXLPE insert was associated with a higher risk of all-cause, aseptic, or septic revision compared to CoCr-CPE. They concluded that their study did not show any evidence of damaging effects from the use of alternative bearings for TKA on short-term outcomes. Longer-term followups will be required to determine whether the potential benefits of wear reduction justify continued use of these bearings (Inacio et al., 2013).

14.6 BACKGROUND ON BIOPOLYMERS IN LIVING TISSUE The use of biopolymers is growing more and more, and its characteristics, such as biocompatibility, biodegradability, and easy organic absorption, have allowed it to grow as a substitute for other materials in the biomedical area and in the food industry. The main biopolymers used are starch, polylactic acid, chitosan, poly (glycolic acid), and polycaprolactone. Its main applications are in the area of tissue engineering, drug release, and packaging. Due to the increasing demand in the application of these biomaterials, many research projects have been developed. The use of biopolymers in the specific field of tissue engineering requires properties such as biodegradability, biocompatibility, bioadhesivity, hemocompatibility, nontoxicity, and stretchability (compatibility with the mechanical properties of the tissue where it is to be implanted). Biopolymers derived from polysaccharides and proteins possess these characteristics, but have poor mechanical properties (Zhang et al., 2013). They are nontoxic, have the ability to interact with living cells, and have low costs (Cascone et al., 2001).

14.7 Present and Future of Biopolymers, Bioplastics

Some of the most commonly used are collagen, chitin/chitosan, alginate, keratin, fibrin, hyaluronic acid, albumin, starch, cellulose, and pectin (Mano et al., 2007; Sionkowska, 2011; Zeeshan et al., 2015).

14.7 PRESENT AND FUTURE OF BIOPOLYMERS, BIOPLASTICS, AND NANOBIOMATERIALS The use of polymers throughout history has been linked with the knowledge about its structure and even synthesis. However, it is a well-known fact that synthetic polymers from the petrochemical industry represent a problem from an environmental point of view because of their long periods of degradation that exceed 100 years. Polymers that are biodegradable or made from renewable resources also represent an alternative possibility. These are newer and less well-known materials that promise a greater sustainability of plastics in the future. Currently, typical polymers, such as starch, cellulose, wool (which are biopolymers), have long been used by society. Paradoxically, the first polymer of industrial use that created a high social impact was a biopolymer called rubber. This is obtained from the bleeding of the bark of trees, from which a very viscous white substance called latex is extracted, with which a variety of plastic products can be made. After this event, in 1839 a scientist named Charles Goodyear modified the structure of rubber by applying a process known as vulcanization, with which it was possible to obtain a highly resistant elastomeric type structure. Another fact of great interest was the modification of cellulose, forming synthetic fibers called rayon. These developments paved the way for the production of polymers from biopolymers, as with what happened with vulcanized rubber and cellulose. Currently, the applications of biopolymers are steadily growing mainly in the medical, food, and tissue engineering sectors; as surgical equipment and instruments, implants inside the body, edible containers, pharmaceuticals, creams for external use, etc. As previous topics developed in the chapter maintained; biopolymers are of great interest for their wide variety of applications, ranging from drugs to cardiovascular implants, including orthopedics, ophthalmology, spine, ear, skin substitutes, biodegradable polymers, etc. The main examples of these applications are bioplastics, drugs, medicines (drugs, smart sensors, and nanomaterials), edible packaging, and tissue engineering. As explained in previous sections, bioplastics are natural polymers based on cellulose, soy, starch, molasses, vegetable oil, and natural rubber, and have the advantage of being biodegradable. Bioplastics can also be defined as a form of plastic derived from renewable biomass sources, such as vegetable oil, corn starch, pea starch, or microbiota, rather than plastic from fossil fuels, which are derived from petroleum.

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At present, polymers derived from natural resources are divided into three large groups depending on their origin: (1) Polymers from biomass (polysaccharides and proteins), such as starch, cellulose, casein, and gluten. (2) Polymers from chemical synthesis using natural monomers, such as biopolyester and PLA (Pacheco et al., 2014). (3) Polymers obtained from microorganisms, such as PHA and PHB. These bioplastics contribute to sustainable development because of their origin (renewable resources), but they are still not up to standard plastic, due to the limited and high production cost of bioplastics compared to plastic produced from petroleum (Bolufer, 2009). Other interesting and important biomaterials to be applied in the medical area are nanobiomaterials, since they have demonstrable properties, such as biocompatibility, hardness, elasticity, mechanical strength, biodegradability, and stability within living tissue. These properties have been a determinant for the growth in the use of nanomaterials in different areas, such as food, pharmaceutical, electronic, industrial, and biomedical industries. Nanobiomaterials, together with bioplastics and other biopolymers, have now become materials of great demand for their range of applications and because they are environmentally sustainable. In addition, biomaterials have been used for the treatment of diseases, such as spinal cord lesions. Moreover, nanocompounds, mainly nanopolymers, are being used for the treatment of rheumatoid arthritis; as biosensors for the control of biomolecules, such as glucose; or in biological synthesis nanoparticles, such as quantum dots. This represents a potential future alternative to the use of organometallic and aqueous synthesis (Zhang et al., 2013; Cai et al., 2007; Cai and Hong, 2012). However, the high prices of some bioplastics and the lack of knowledge regarding technology, as recent as that of nanomaterials, do not provide a sufficiently clear picture for its future use. Due to the importance of nanomaterials, many studies involving these novel compounds have been carried out, such as nanorafeno, nanochitosan, polylactic acid, nanoparticles, etc. Graphene is a newly discovered nanomaterial composed only of carbon atoms, forming an overlapping structure with sp2 type hybridization, arranged in a Bravais crystal network in hexagonal form, and forming a structure similar to that of honeycomb. The main properties of graphene are: high thermal conductivity (5000 m21 k21), electric conductivity, high elasticity, hardness, high surface area (2360 m2 g21), light weight, transparency, easy synthesis, high mechanical strength (200-times higher than that of steel), and highly porous structure. These characteristics, unlike other biomaterials, allow it to be used as an extracellular matrix in the regeneration of tissues and variants, including graphene oxide (Veliz, 2016a,b; Felli et al., 2015). This structure is advantageous since it is physiologically stable, compatible, and capable of transporting drugs and biomolecules; in turn permitting it to have great application in the biomedical area (Veliz, 2016a,b; Shen et al., 2012). In addition, graphene is used as a

14.7 Present and Future of Biopolymers, Bioplastics

reinforcement for other biomaterials because it provides a heterogeneous structure with mechanical, chemical, and electrical properties that it did not previously possess. For example, the mechanical strength and elasticity of PVA and PMMA was increased by reinforcement with graphene oxide (Zhang et al., 2011; Ramanathan et al., 2008). The most important applications of graphene include its use in biosensors of biomolecules, in photothermal and gene therapy, in the study of bioimagenes, in tissue engineering, and as drug release systems (Astillero, 2012). In addition to its use in biomedicine, graphene has been explored for use in other areas, such as electronics where graphene transistors are used; it has also been used in engines to improve efficiency as well as in renewable energy (Gamero Este´vez, 2010). Due to the findings of a study on nanomaterials, the diversity of uses and accelerated growth that nanocomposites have every day in key areas of biomedicine, such as tissue engineering, biochemistry, genetics, biomedical engineering, among others, mean that the applications of these materials increase daily. Investigations carried out by Felli et al. (2015) give a review of the different types of graphene used in tissue engineering, and in this review they intend to give a sample of the family of graphene as well as to give an idea of the progress made to date in this field of research. They analyzed the production methods of graphene; conventional and green, and found that their excellent physical and chemical properties as well as their biocompatibility with living tissue are of interest to be used with a high probability of success in engineering. However, research into the toxicity of the graphene family of nanoparticles is still in its blooming stages and it is difficult to conclude the potential health risks associated with their use and whether they should be modified chemically; in some cases, to reduce their toxic risk (Syama and Mohanan, 2016). Similarly, the biodegradation process is not entirely clear and much remains to be studied. However, the enzymatic degradation of different types of graphene oxide dispersions with the human enzyme, myeloperoxidase, has already been demonstrated, which makes this material very promising for applications in tissue engineering (Kurapati et al., 2015). Other encouraging materials for tissue engineering are biomaterials of marine origin, such as alginate and chitosan (Ratner et al., 2013). The mechanical properties necessary to produce scaffolds with great porosity could be improved on by the addition of graphene and its derivatives. Sodium alginate can be considered as a biodegradable biomaterial that has been used in various fields, such as controlled drug release, tissue engineering, and biological studies. Guedes et al. (2015) in their book in the chapter on polyethylene blends, composites, and nanocomposites, investigated the behavior of UHMWPE as a joint in total arthroplasty as well as THK. The analysis did so by studying the behavior in the presence of antioxidants, such as vitamin E and vitamin C, under the action of gamma rays. Subsequently they did a review and discussion on biocompatibility, manufacturing processes, tribological behavior, aging by

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oxidation, and the observed changes in mechanical properties. Finally, they analyze the viscoelastic behavior of UHMWPE and its implications on the long-term survival of TJA. The first conclusion from the results was to highlight the impressive rise that medical grade UHMWPE has had in the past five decades. Another relevant aspect is the complexity of UHMWPE, which complicates the lifetime analysis and mechanical behavior prediction of UHMWPE as a load bearing component in TJA and TKA. Additionally, they infer that wear is dependent on many events occurring simultaneously and sequentially. This means that the amorphous and crystalline UHMWPE structure constantly undergoes interactions, which cause changes both in the mechanical properties and in the contact area of the amorphous crystalline phase; but they affirm that the interaction process is not completely understood. In addition, they emphasize that the oxidation processes that occur are due to the oxygen being fed by the fluids surrounding the UHMWPE component, which deteriorate the mechanical properties. Finally, they conclude that the viscoelastic properties of UHMWPE increase the creep of the material, a situation that should be corrected in later investigations.

14.8 CONCLUSIONS This chapter describes the characteristics of different biomaterials, mainly biopolymers, both natural and synthetic, as well as bioplastics that are of organic, vegetable, and mixed origin, such as biopolyethylene, and finally, nanomaterials such as graphene and its derivatives. The graphene specifically dates from little time of discovery, presents exceptional qualities to be used in the biomedical area. Its nanomolecule sensing properties, biocompatibility, coupled with its excellent thermal and electrical properties, make it one of the main biomaterials used in tissue engineering. The main biopolymers used are starch, PLA, chitosan, PGA, and polycaprolactone. The most used blends are between polysaccharides and aliphatic or aromatic polyesters. These biopolymers have important properties, such as biocompatibility and biodegradability, but lower mechanical properties and require the forming of “hybrid” biopolymers with synthetic polymers, such as PP or polyurethane, to be effectively applied in tissue engineering. Nanobiopolymers, such as nanoparticles of chitosan and PLA, are also presented as possible drugs in the near future. Studies show the importance of these new technologies at the level of living tissue, as well as their importance at an environmental level, for their easy biodegradability compared to other materials. In addition, reference is made to the future use of these materials. However, here the results obtained can not be considered completely satisfactory at the level of the organism as is the case of nanomaterials due to the presence of a synthetic polymer in the structure of the bioplastic. In other words, although they present an environmental advantage due to their degradation through microorganisms, the presence of synthetic

14.8 Conclusions

polymers in the structure of the mixed bioplastics undermines their sustainability from an economic point of view. Several investigations carried out, mainly on the use of nanocomposites and nanoparticles, have led to the conclusion that more tests, studies, investigations, and analyses are required, since their safety or side effects in drugs or food have not been completely verified. The behavior of medical grade UHMWPE is extensively detailed because of its great importance as a component used in orthopedics, both inside and outside of living tissue. From the different analyses, it can be seen that medical grade UHMWPE varies its behavior according to the type of atmosphere of ionizing irradiation. The results show resistance to oxidation in the absence of oxygen, with the predominance of crosslinking. On the contrary, in air it is observed that oxidation processes predominate as well as the formation of oxidation cascades, and as a consequence of which, the reactions and mechanisms of chain rupture prevail. The successful results observed in the behavior of medical grade UHMWPE, when subjected to combined external factors, mainly ionizing irradiation, annealing and melting heat treatments, and storage in natural substances, such as vitamin E and aloe vera are also mentioned. In this case, it is observed that the properties of the material, mainly its resistance to wear, improve as a result of crosslinking; Davidson (2012) showed an improvement of nearly 20% with the use of vitamin E. In this case, improvements around 17% in the value of this property were observed (see Fig. 14.4A and B). The appreciated behavior result innovative and important in the use of this substances, due to allows reducing operating costs of storage and transport that favor the properties of medical grade UHMWPE in vitro. On the contrary, other investigations with components such as SBF and bovine serum on UHMWPE samples were shown to lead to oxidation cascades. Section 14.4 describes in detail the mechanisms of debris and osteolysis as the main cause of the aseptic failure of orthopedic prostheses in tribological systems involving UHMWPE as an orthopedic component of the prostheses in hip and knee arthroplasty. Finally, different investigations comparing UHMWPE crosslinked with other tribological components used in orthopedics, such as conventional PE with zirconium oxide, and cobalt-chromium with PE, show that tribological systems involving crosslinked UHMWPE (HXLPE) show higher wear resistance than the rest. This means that the presence of HXLPE decreases the harmful mechanisms of osteolysis and debris, but cannot prevent them. These analyses are highly useful since they are fundamental to optimizing the properties of the material. This means that by understanding the in vitro behavior of UHMWPE, it is possible to reduce the wear that represents the leading cause of failure in prostheses and UHMWPE components when used in living tissue, and of decreasing the longevity of prostheses and orthopedic components. This knowledge in turn is of great interest for the design and manufacture of these components; however, tests of the components on living tissue are required.

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Mano, J.F., Silva, G.A., Azevedo, H.S., Malafaya, P.B., Sousa, R.A., Silva, S.S., et al., 2007. Natural origin biodegradable systems in tissue engineering and regenerative medicine: present status and some moving trends. J. R. Soc. Interface 4 (17), 999 1030. Mata, A., Kim, E.J., Boehm, C.A., Fleischman, A.J., Muschler, G.F., Roy, S.A., 2009. Three-dimensional scaffold with precise micro-architecture and surface micro-textures. Biomaterials 30, 4610 4617. McKeen, Lawrence, W., 2014. Handbook of Polymer Applications in Medicine and Medical Devices. Published by Elsevier Inc. Mitragotri, S., Lahann, J., 2009. Physical approaches to biomaterial design. Nat. Mater. 15, 15 23. Mutlu Hatice, Meier Michael, A.R., 2010. Castor oil as a renewable resources for the chemical industry, vol. 112, issue 1, no. 1, pp. 10 30. Nich, C., Takakubo, Y., Pajarinen, M.J., Salem, A.A., Sillat, T.R., Allison, J., et al., 2014. Macrophages key cells in the response to wear debris from joint replacements. J. Biomed. Mater. Res. A 101 (10), 3033 3045. Oldini, C., 2014. El camino del titanio como material de reemplazo o´seo. Revista facultad de ciencias exactas, fı´sicas y naturales, vol. 1, no. 1. Oral, E., Keith, K., Wannomae, N.M., Harris, W.H., Muratoglu, O.K., 2006. Radiation crosslinking in ultra-high molecular weight ‘polyethylene for orthopaedic applications, Nucl. Inst. Methods Phys. Res. B, 265. pp. 18 22. Pacheco, G., Flores, N.C., Romina, R.S., 2014. Biopla´sticos. BioTecnologı´a, An˜o 18 (2). Palakurthi, S., Yellepeddi, V., Kumar, A., 2013. Nanocarriers for cytosolic drug and gene delivery in cancer therapy. In: Olsztynska, S. (Ed.), Biomedical Engineering, Trends, Research and Technologies. INTECH, pp. 245 272. Phillips, F.M., Garfin, S.R., 2005. Cervical Disc Replacement Spine, 30 (17S) Supplement: 1, pp. S27 S33. Plaza Torres, M., Aperador, W., 2015. Nuevos materiales para mejorar los niveles de corrosio´n. Rev Cubana de Investigaciones Biome´dicas 34 (3). Puertolas, J.A., Martinez-Morales, M.J., Mariscal, M.D., Medel, F.J., 2010. Thermal and dynamic mechanical properties of vitamin E infused and blended ultra-high molecular weight polyethylenes. J. Appl. Polym. Sci. (JAPS) . Available from: https://doi.org/ 10.1002/app.33454. Ramakrishna, S., Mayer, J., Wintermantel, E., Leong, K.W., 2001. Biomedical applications of polymer-composite materials: a review, Composites Science and Technology, vol. 61. View at Publisher View at Google Scholar View at Scopus, pp. 1189 1224, no. 9. Ramanathan, T., Abdala, A.A., Stankovich, S., Dikin, D.A., Herrera-Alonso, M., Piner, R. D., et al., 2008. Functionalized graphene sheets for polymer nanocomposites. Nat. Nanotechnol. 3, 327 331. Ratner, B.D., Hoffman, A.S., Schoen, F.J., Lemons, J.E., 2013. Biomaterials science, An Introduction to Materials in Medicine, third ed. Academic Press. Rı´os, R., Pue´rtolas, J.A., Martı´nez-Nogue´s, V., Martı´nez-Morlanes, M.J., Pascual, F.J., Cegon˜ino, J., et al., 2013. Mechanical behavior, microstructure and thermooxidation 5 properties of sequentially crosslinked ultra-high molecular 6 weight polyethylenes. J. Appl. Polym. Sci. 1 30. Available from: https://doi.org/10.1002/app.38956. Schwarz, E., 2016. Wear Debris-Induced Osteolysis and Aseptic Loosening of Total Joint Replacements Schwarz Lab. Medical Center, University Rochester. Science in School, Polymer in Medicine, Issue 21, winter, 2011. Shen, H., Liming, Z., Min, L., Zhijun, Z., 2012. Biomedical applications of graphene. Theranostics 2 (3), 283 294.

Further Reading

Singhal, R., Salim, J., Walker, P., 2005. Idiopathic multicentric osteolysis: a case report and literature review. Acta Orthop. Belg. 71, 328 333. Sionkowska, A., 2011. Current research on the blends of natural and synthetic polymers as new biomaterials: Review. Prog. Polym. Sci. 36 (9), 1254 1276. Stephens, C., Benson, R., Martı´nez-Pardo, M.B., Naaker, E., Walter, J., Stephens, T., 2005. The effect of dose rate on the crystalline lamellar thickness distribution in gammairradiation of UHMWPE. Nucl. Inst. Methods Phys. Res. Sect. B: Beam Interact. Mater. Atoms 236, 540 545. Syama, S., Mohanan, P.V., 2016. Safety and biocompatibility of graphene: a new generation nanomaterial for biomedical application. Int. J. Biol. Macromol. 86, 546 555. Trevisan, C., Nava, V., Mattavelli, M., Garcia Parra, C., 2013. Bisphosphonate treatment for osteolysis in total hip arthroplasty. A report of four cases. Clin. Cases Miner. Bone Metab. 10 (1), 61 64. Valero, M., Ortegon, Y., Uscategui, Y., 2013. Biopolimeros: avances y perspectivas. Dyna 181, 171 180. Van Dijk, M., Smit, T.H., Arnoe, M., et al., 2003. The use of poly-L-lactic acid in lumbar interbody cages: design and biomechanical evaluation in vitro. Eur. Spine 28, 1802 1808. Veliz, O., Miguel, J., 2016a. El grafeno y sus derivados en la ingenierı´a tisular, revista nereis, vol. 8, pp. 71 81. Veliz, O., Miguel, J., 2016b. Biomedical Applications of Graphene, MoleQla. Wang, X., 2013. Overview on biocompatibilities of implantable biomaterials. Rosario Pignatello. Advances in Biomaterials Science and Biomedical Applications. INTECH, pp. 111 155. Wong, J., Leach, J., Brown, X., 2004. Balance of chemistry, topography, and mechanics at the cell-biomaterial interface: issues and challenges for assessing the role of substrate mechanic on cell response. Surf. Sci. 570, 119 133. Zeeshan, S., Najeeb, S., Khurshid, Z., Verma, V., Rashid, H., Glogauer, M., 2015. Biodegradable materials for bone repair and tissue engineering applications. Materials 8 (9), 5744 5794. Zhang, X., Morham, S.G., Langenbach, R., Young, D.A., Xing, L., Boyce, B.F., et al., 2001. Evidence for a direct role of cyclo-oxygenase 2 in implant wear debris-induced osteolysis. J. Bone Miner. Res. 16 (4), 660 670. Zhang, L.Z., Wang, Z., Xu, C., Li, Y., Gao, J., Wang, W., et al., 2011. High strength graphene oxide/polyvinyl alcohol composite hydrogels. J. Mater. Chem. 21, 10399 10406. Zhang, Y., Chan, H.F., Leong, K.W., 2013. Advanced materials and processing for drug delivery: the past and the future. Adv. Drug Deliv. Rev. 65 (1), 104 120.

FURTHER READING Brandi, F., Sommer, F., Goepferich, A., 2007. Rational design of hydrogels for tissue engineering: impact of physical factors on cell behavior. Biomaterials 28, 134 146. Dri Dietary Reference, 2005. Intakes for Energy, Carbohydrate, Fiber, Fat, Fatty Acids, Cholesterol, Protein, and Amino Acids. Institute of medicins.

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Fishman, M., Coffin, D., Onwulata, C., Willett, J., 2006. Two stage extrusion of plasticized pectin/poly(vinyl alcohol) blend. Carbohydr. Polym. 65 (2006), 421 429. Li, M., Mondrinos, M.J., Chen, X., Gandhi, M.R., Ko, F.K., Lelkes, P.I., 2006. Coelectrospun poly(lactide-co-glycolide), gelatin, and elastin blends for tissue engineering scaffolds. J. Biomed. Mater. Res. Part A 79A (4), 963 973. Maitz, M.F, 2015. Applications of synthetic polymers in clinical medicine. Biosurf. Biotribol. 1 (3), 161 176. Mingyong, G., Lu, P., Bednark, B., Lynam, D., Conner, J.M., Sakamoto, J., et al., 2013. Templated agarose scaffolds for the support of motor axon regeneration into sites of complete spinal cord transection. Biomaterials 34, 1529 1536. Parra-Cid, C., Tiscaren˜o Pe´rez, A., Go´mez Garcı´a, R., 2014. Investigacion en discapacidad. 3(1):27. Www.medigraphic.org.mx. Rodrı´guez, S., Joana, L., Alzate, O., Eduardo, C., 2016. Aplicaciones de mezclas de biopolı´meros y polı´meros sinte´ticos. Revisio´n bibliogra´fica 2, 25. Rogers, M.J., Gordon, S., Benford, H.L., Coxon, F.P., Luckman, S.P., Monkkonen, J., 2000. Cellular and molecular mechanisms of action of Bisphosphonates. Cancer 88 (Suppl. 12), 2961 2978. Shi Rong, G., Liu, H.T., Huang, X.L., 2010. Behaviour and wear debris characterization of UHMWPE on alumina ceramic, stainless steel, CoCrMo and Ti6Al4V hip prostheses in a hip joint simulator. J. Biomimet. Biomater. Tissue Eng. 7, 7 25. Solis-Arrieta, L., Leo´n-Herna´ndez, S.R., Villegas-Castrejo´n, H., 2012. Ana´lisis de partı´culas de desgaste en tejido periprote´sico de cadera y rodilla con microscopia electro´nica de barrido. Cir. Cir. 80, 239 246. Vigara Astillero, G., 2012. Grafeno, el material del futuro. posibilidad real o pura fantası´a?, Revista MoleQla, no8. Sevilla: Universidad Pablo de Olavide, Diciembre 51, % pp. 62 65. Yuqi Yang, Y., Asin Abdullah, M., Zhiwen, T., Yuehe, D.D.L., 2013. Graphene based materials for biomedical applications. Mater. Today. 16(10):365 373. Zuluaga, C., Fabio, H., 2013. Alguna aplicaciones del a´cido poli-l-la´ctico: Revista de la Academia Colombiana de ciencias exactas, fı´sicas y naturales, ISSN 0370-3908. 37 (142):125 143. ?

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Application of polymethylmethacrylate, acrylic, and silicone in ophthalmology

15

Hossein Aghamollaei1, Shiva Pirhadi2, Soodabeh Shafiee3, Mohammad Sehri4, Vahabodin Goodarzi5 and Khosrow Jadidi4 1

Chemical Injuries Research Center, Systems biology and Poisonings Institute, Baqiyatallah University of Medical Sciences, Tehran, Iran 2Department of Biomedical Engineering, Science and Research Branch, Islamic Azad University, Tehran, Iran 3Department of Biochemistry, Faculty of Biological Sciences, Tarbiat Modares University, Tehran, Iran 4Vision Health Research Center, Semnan University of Medical sciences, Semnan, Iran 5Applied Biotechnology Research Center, Baqiyatallah University of Medical Sciences, Tehran, Iran

15.1 INTRODUCTION The human eye is a very complex system with an important role in everyday life. Despite its importance, this organ is the target of a wide variety of disorders and, hence, has the center of attention for implants and accessory biomedical devices. A wide range of biomaterials are used to fabricate ocular devices to correct functional deficiencies caused by disease, age, and ocular trauma. Due to their unique properties, polymethylmethacrylate (PMMA) and silicone are used in many ocular implants.

15.1.1 SILICONE Silicone has been used in medical devices for more than 50 years. Properties of silicone include chemical stability, dynamic mechanical properties, and biocompatibility (Baino, 2010). Moreover, silicone chemistry is suitable for optimization, for instance, attaining a specific viscosity of this material and varying degrees of optical clarity and permeability could be achieved relatively easily, making it a promising material for ocular implants (Lloyd et al., 2001). The relatively large bond angles of the repeating helical silicon oxygen (Si O) structure of the polymer backbone is one of the main reasons for the wide range of applications of silicone. Variability of the substituent, groups that attach to the open valences of the silicon atoms, is another reason for its multifunctional properties. The bond angles yield large amounts of free volume, leaving space for Materials for Biomedical Engineering: Thermoset and Thermoplastic Polymers. DOI: https://doi.org/10.1016/B978-0-12-816874-5.00015-3 © 2019 Elsevier Inc. All rights reserved.

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FIGURE 15.1 Structure of polysiloxane chain.

design or, more specifically, for managing the amount and type of substituent group and filler, such as resin, that are often assimilated into a silicone system and further increase the variability of its products (Fig. 15.1) (Mark, 2004; Pierscionek et al., 2015; Baino et al., 2014). As an example, incorporation of phenyl groups into the siloxane polymer backbone of silicone molecules is capable of influencing silicone’s permeability and refractive index (RI). RI is defined as the ratio of the velocity of light in a vacuum to the velocity of light in a material and is of vital importance in the case of ocular implants. However, refraction is not the property of every light that strikes a surface or interface. In some cases, light reflects off, instead of transmitting through, the material. The more light a material refracts and the less it reflects, the more transparent it will appear. The more slowly light passes through the silicone, the larger is the RI. Silicones which are more optically clear are much more responsive to light, meaning that far more light waves are transmitted through the material than reflected back out. One of the reasons that make silicone a suitable choice in ophthalmic applications is the opportunity that it affords for RI adjustment according to different requirements for repairing or enhancing sight. In this way, silicone is accommodating to the function of intraocular and contact lenses (Mark, 2004). Assuming a natural rather than intrusive feel is another task of medical devices, and silicone rises to the challenge to achieve this. In ophthalmic applications, silicone allows the benefits and not the burden of a device to meet the eye.

15.1.2 POLYMETHYLMETHACRYLATE Poly(methyl 2-methylpropenoate) (Fleming et al., 1989) is a polymer of methyl methacrylate, with the chemical formula (C5H8O2)n (Fig. 15.2). Physically, it is a clear, colorless polymer. It is available in both pellet and sheet form under the commercial names of Plexiglas, Acrylite, Perspex, Plazcryl, Acrylplast, Altuglas, Lucite, etc. It is also called acrylic glass or acrylic (Kumar et al., 2016). Structurally, PMMA is a member of a huge family of methacrylate esters in which the group attached can be any alkyl group or even aryl group (Ali et al., 2015). Both forms of attached groups can be replaced by different reactive or dormant groups which gives great flexibility over structural design of PMMA. As an example, attachment of an alcohol group to the ester unit produces so-called HEMA. This attachment results in properties like water solubility and also provides a space for the attachment of other additional groups to the alcohol

15.2 Application of Biomaterials in Intraocular Lenses

FIGURE 15.2 Chemical structure of poly(methyl methacrylate).

Table 15.1 Physical Properties of Polymethylmethacrylate (PMMA) Density

1.15 1.19 g/cm3

Water absorption Moisture absorption at equilibrium Linear mold shrinkage Melt flow

0.3% 2% 0.3% 0.33% 0.003 0.0065 cm/cm 0.9 27 g/10 min

component of poly-HEMA. Acrylates share this structural flexibility which makes these two families of monomers and their polymers some of the most-widely explored and used of all those available currently (Ali et al., 2015).

15.1.2.1 Properties and advantages of polymethylmethacrylate PMMA has many desirable properties including high mechanical strength and low elongation at break. Therefore, it is one of the hardest thermoplastics and is also highly scratch-resistant. PMMA has low moisture and water absorbing ability which cause its products to have good dimensional stability. Both of these characteristics increase as the temperature rises (Ali et al., 2015). The main physical characteristics of PMMA are shown in Table 15.1. PMMA has perfect optical properties, for instance, in comparison to glass it transmits more light (up to 93% of visible light). This feature of PMMA along with its biocompatibility makes it a good choice for replacement of intraocular lenses (IOLs) and for contact lenses. Unlike glass, PMMA does not filter ultraviolet light. It transmits UV light down to 300 nm and allows infrared light of up to 2800 nm to pass (Van Krevelen and Te Nijenhuis, 2009). Table 15.2 summarizes some of optical properties of PMMA.

15.2 APPLICATION OF BIOMATERIALS IN INTRAOCULAR LENSES IOLs are a type of lens that is permanently implanted inside the eye. IOLs are sometimes implanted to replace the natural lens in the eye (usually after cataract

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Table 15.2 Optical Properties of Polymethylmethacrylate (PMAA) Optical Properties

Value

Haze Transmission, visible Refractive index

1% 96% 80% 93% 1.49 1.498

surgery) and are sometimes implanted into the eye over the natural lens (to correct refractive errors in nearsighted or farsighted patients or those with astigmatism), in which case they are referred to as phakic IOLs. The first IOL was implanted into an eye by Harold Ridley on November 29, 1949 (Apple and Sims, 1996). This ophthalmologist had observed that the shattered pieces of cockpit (made from PMMA) that had entered pilots’ eyes following antiaircraft bullets in World War II had caused them no complications. This observation led to the invention of IOLs. Before the implantation of these lenses, patients had to use thick spectacles or special contact lenses.

15.2.1 LENSES USED IN CATARACT SURGERY These lenses are implanted during cataract surgery to replace the patient’s natural lens. The lens is a transparent biconvex body located behind the cornea and iris that provides one-third of the refractive power of the eye. The lens is fixed behind the iris by the zonulae originating from the ciliary body. With its inherent flexibility, the lens is able to contract or expand its spherical shape depending on the tension exerted by the zonulae and, thus, change the refractive power of the eye to see near and far objects. This process is called accommodation. The human lens is an epithelial tissue that is enclosed by an external membrane (the capsule). The anatomy of the lens involves the capsule, epithelium, membrane, and nucleus (Roberts, 2011). The lens cells are composed of substantial amounts of proteins called crystalline along with electrolyte and water. The proteins are delicately placed for the purpose of lens transparency and to allow light to pass through the lens’ layers. The uniform and transparent structure of the lens is lost and it becomes opaque with age, ongoing epithelial lens cell proliferation, compaction of the cells, and other risk factors such as smoking and exposure to UV radiation. The opacity of the lens can be caused by inflammation of the eye, diabetes, surgical manipulations, trauma to the eye, the use of certain medications by one’s mother during the embryonic period, infectious diseases during pregnancy, or metabolic and genetic factors. The loss of transparency and the resultant opacity of the human lens is referred to as cataract. There are no effective treatment options available for this disease and surgery is the only definite treatment. In cataract surgery, the natural lens that is opaque is replaced by a transparent artificial lens. Common

15.2 Application of Biomaterials in Intraocular Lenses

lens extraction techniques currently used include Extra Capsular Cataract Extraction (ECCE) and Phaco. Both techniques extract the lens via an incision (ECCE 5 12 14 mm and phaco 5 1.8 3.2 mm) and leave the lens capsular bag intact (Princz et al., 2016). The smaller the lens replacement incision, the better will be the surgery outcome. To achieve this objective, it is necessary to extract the opaque lens with the smallest incision possible (Baino, 2010) and implant the IOL through the same incision (Lloyd et al., 2001). The former can be accomplished by taking advantage of the innovations made in phaco machines, and the latter with the use of foldable IOLs. Any failure in the latter renders the benefits gained from accomplishing the former ineffective (Agarwal et al., 2002). In phacoemulsification, the lens is broken down with ultrasound waves using a fine device inserted into the eye through an incision; given that this incision is very fine, no sutures are required and the wound will heal by itself. The extraction of the opaque lens leaves the posterior capsule untouched for the artificial lens to be inserted over it. A transparent artificial lens is folded and inserted into the eye through the same fine incision. The folded lens opens inside the eye and fixes itself properly in the eye.

15.2.2 PHAKIC LENSES Lenses designed for intraocular implantation and insertion over the natural lens are referred to as phakic lenses. These lenses are most widely used to correct refractive errors, especially high nearsightedness that cannot be corrected with laser refractive surgery on the cornea. Phakic lenses are either the anterior chamber type or the posterior chamber type. Implantable collamer lens (ICL) is the most well-known posterior chamber type and artisan PIOL is the most wellknown anterior chamber type. ICL is implanted right over the natural lens of the eye, while artisan lens is fixed onto the iris (Iris-Claw PIOL); artisan lens has another variety called aphakic artisan, signifying that aphakic eyes are also implantable. Aphakic eyes lack either natural or artificial IOLs. Foldable artisan lenses are called artiflex lenses. Unlike artisan lenses that are implanted in the eye through a 5 6 mm incision (in optical size), this lens is inserted into the eye through a 3 mm incision.

15.2.3 INTRAOCULAR LENS STRUCTURE IOLs are made from optic and haptic parts. The optic part is in charge of the refractive function of the lens and the haptic is responsible for attaching the lens to the surrounding intraocular structures. Lenses with optic and haptic parts made of the same material and with uniform structures are called one-piece or singlepiece, and lenses with haptic parts made from a different material and separately attached to the optic part are called multi-piece or three-piece. The optic diameter and length of modern lenses are 5.5 7 mm and 12 14 mm, respectively.

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15.2.4 IMPLANT POSITIONS FOR INTRAOCULAR LENSES There are two general groups of IOLs: 1. Anterior Chamber IOLs The anterior chamber type of lenses can take two positions: Angle-supported: In this type, the haptic parts are positioned in the angle of the eye (where the iris meets the cornea). Iris-fixated: In this type, the pikes of the lenses are fixed on the iris (like the artisan type) 2. Posterior Chamber IOLs The posterior chamber type of lenses can take three positions: a. Right behind the iris: In this type, the pikes are placed in the ciliary sulcus (where the ciliary body meets the iris). This implant position is referred to as sulcus fixation. b. In the natural lens (in the space inside the capsule): Following the drainage of the patient’s lens, the sulcus diameter is about 12 mm and the capsule diameter about 10 mm. Sulcus-fixated posterior chamber lenses should preferably be of the three-piece or the large-diameter one-piece type. c. Scleral fixation: In this type, the lens pikes are sutured to the adjacent sclera from the inside.

15.2.5 THE PROPERTIES OF INTRAOCULAR LENS MATERIALS IOL materials should have specific properties (Agarwal et al., 2002), including: Biocompatibility: These materials should be chemically neutral and not have inflammatory, infections, or immune reactions with the ocular tissues and should not be carcinogenic. Optical Compatibility: The materials should possess all the optical properties of natural lenses (crystalline lenses), should be perfectly transparent, have high optical resolution, and be able to filter out UV radiation. An ideal IOL is able to focus on different distances. Mechanical Compatibility: The materials should be resistant against the mechanical pressures exerted during manufacturing, but, at the same time, be flexible enough to be implanted in the eye with confidence and with no risk of breaking. In addition to these properties, what differentiates foldable lenses from relatively rigid and inflexible lenses is not merely their foldability and ductility for implantation through a fine incision, but their rapid return to their initial shape and dimensions. These lenses are able to spring back to shape quickly or slowly and, therefore, have superior mechanical compatibility. All the noted properties depend on the chemical composition of the materials with which the lenses are made. The key criteria for selecting materials for IOLs, especially foldable IOLs, are concerned with the glass transition temperature,

15.2 Application of Biomaterials in Intraocular Lenses

elongation, and tackiness. The glass transition temperature of monomers should not exceed 37 C, which is the normal body temperature, since these polymers possess foldability only at temperatures higher than 37 C. It is therefore better to use polymers with a glass transition temperature below the normal body temperature and not higher than normal room temperature, so that the lenses can easily fold at room temperature. Another property is mechanical strength. The lenses should be strong enough to fold without breaking or disintegration. Lenses with 150% 200% elongation are preferred. Tackiness is expressed in Tack Quotient, which is defined as the ratio of the material’s actual tackiness to its standard tackiness. Materials with a Tack Quotient of 1 1.5 are suitable for IOLs (Tripti et al., 2009).

15.2.5.1 Intraocular lens materials Once Ridley implanted the very first IOL made from PMMA in the eyes of a patient, IOLs began being manufactured from different materials and were massmarketed. Conventional materials currently used in manufacturing IOLs include PMMA, silicone, hydrophobic acrylic, hydrophilic acrylic, and collamer.

Acrylic Acrylic is a polymer derived from acrylic acid or meth acrylic acid esters and can be rigid or flexible (Agarwal et al., 2002).

Poly methyl methacrylate PMMA, also called Perspex or Plexiglas, is a rigid, nonfoldable and hydrophobic acrylic (water content , 1) with a RI of 1.49. PMMAIOLs have an optical diameter of 5 7 mm and are usually single-piece with fragile haptics, unless compression molding is used in their production. These lenses are usually thin, so that the material’s rigidity can balance out the low RI (Bellucci, 2013). Although PMMA has been the material of choice for IOLs over the past 40 years, it is still associated with extensive endothelial damage immediately after implantation, the adhesion of inflammatory cells after surgery, and potentially serious postoperative complications. Moreover, PMMA lenses cannot be folded at room temperature and, therefore, need an incision as large as the lens itself during surgery, which prolongs the recovery of the eye and leads to astigmatism in the eye as an undesirable postoperative complication (Bozukova et al., 2010). As a result, they are rarely used nowadays, although they still have a major role in countries where the ECCE technique is used (Colvard, 2009).

Foldable hydrophobic acrylic Foldable acrylic is a copolymer of phenylethylacrylate and phenylethyl methacrylate. This crosslinking aims to generate foldability and provides acrylic with viscoelasticity and good three-dimensional stability (Agarwal et al., 2002). Hydrophobia is a physical property of molecules that avoid water. IOL materials are divided into a hydrophobic or a hydrophilic group depending on the angle

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that a drop of water makes with the surface of the material. Water forms a bigger contact angle on hydrophobic surfaces. The normal contact angle of these materials with water is 73 degrees. The RI of these lenses varies from 1.44 to 1.55. These lenses are available in one-piece and three-piece designs. The optical diameter of the lenses varies in the range of 5.5 7 mm and their length varies around 12 13 mm. They are easily implanted in the eye and require a minimum incision of 2.2 mm. They have little tendency to self-center and their correct positioning during implantation requires utmost care (Bellucci, 2013).

Foldable hydrophilic acrylic (hydrogel) Hydrophilic acrylic materials are composed of Hydroxyethyl methacrylate (polyHEMA) and acrylic monomer (Bellucci, 2013). These materials make up a completely heterogeneous composition that contains large amounts of water. Their contact angle with water is less than 50 degrees. The lenses made from these materials contain different amounts of water, usually between 18% and 38% (Bellucci, 2013; Bozukova et al., 2010; Colvard, 2009). These lenses are cut in a dehydrated mode and are then hydrated and kept in a solution. Unlike hydrophobic lenses, hydrophilic lenses should be packaged submerged in normal saline (Bellucci, 2013). The RI of these materials varies in the range of 1.46 1.48 (Bellucci, 2013). The lenses are mostly single-piece and can pass through incisions less than 2 mm; as a result, they are suitable for microincision cataract surgery (MICS). Small incisions are better for the recovery of the eye.

Silicone Silicone is a polyorganosiloxane polymer used as an elastomer (polydimethylsiloxane) that has biomedical applications (Agarwal et al., 2002). The first foldable IOL was made from silicone and was implanted in the eye in 1989 (Bozukova et al., 2010). Silicone is hydrophobic and has a contact angle of 99 degrees with water, which is greater than the angle in hydrophobic acrylic materials. The RI of silicone varies from 1.41 to 1.46. The optical diameter of these lenses is between 5.5 and 6.5 mm. Common silicone lens models are three-piece and have an optic part made from PMMA, polyvinyl difluoride (PVDF), or polyamide materials (Bellucci, 2013). The elasticity and malleability of silicone make it suitable for use in accommodative lenses, since these lenses should be able to change shape millions of times during the patient’s life. The sudden opening of silicone IOLs in the anterior chamber remains a problem for surgeons (Bellucci, 2013). The intraocular opening of acrylic lenses is much gentler and more controllable than that of silicone lenses. Greater care should be taken in the use of silicone lenses since they are spontaneously covered with a liquid hydrophobic film whenever in contact with silicone oil. Silicone oil is used to repair retinal detachment. These lenses are, therefore, not the right choice for patients with silicone oil in their eyes or for highly nearsighted patients

15.2 Application of Biomaterials in Intraocular Lenses

or diabetic patients who may require vitreoretinal surgery (Bellucci, 2013; Bozukova et al., 2010)

Collamer There is another type of silicone IOL that contains collagen and is known as collamer. The term “collamer” is made by combining the terms collagen and polymer (Bozukova et al., 2010). This material is made from a proprietary hydrophilic collagen polymer (a copolymer of 63% hydroxyethyl-methyl-acrylate, 0.2% porcine collagen, and 3.4% benzofenone for UV absorption) with a 34% water content. The RI of this material is 1.45 at the temperature of 35 C (Argal, 2013). This material is used exclusively in the manufacturing of phakic and aphakic lenses by STAAR Co., including Visian ICL. These lenses are highly compatible with ocular tissue and easily implanted in the eye. The high water content of these materials makes them very soft, such that they are also suitable aphakic IOLs. IOLs made from this material need to be packaged wet (Koch, 2005).

15.2.6 THE EFFECT OF DIFFERENT INTRAOCULAR LENS MATERIALS ON POSTOPERATIVE COMPLICATIONS 15.2.6.1 Posterior capsular opacification Intraocular lens implant following cataract surgery arouses the external body’s reaction to IOL and the response of the lens epithelial cells (LECs). Both these responses may produce different patterns and are generally considered indicators of the biocompatibility of IOL. Inflammatory cell reaction to IOLs involves the adhesion of macrophage derived cells to the anterior surface of IOLs. The reaction of LECs is explained by an early response such as anterior capsular opacification (ACO) and membrane growth from the rhexis edge to the IOL surface, which is composed of a sheet of proliferated LECs (Fig. 15.3). The proliferation and migration of LECs from the anterior capsular edge and the equatorial region toward the posterior capsule may lead to future posterior capsular opacification (PCO) and the emergence of Elschnig pearls (Colvard, 2009; Tognetto et al., 2003). PCO is the most common complication in cataract surgery that leads to the loss of sight (Sinha et al., 2013). In cataract surgery, the posterior capsule is normally left intact. This capsule frequently suffers various degrees of opacity after surgery, severe cases of which are referred to as After Cataract. A meta-analysis of 69 articles examined the incidence of this complication. According to a general estimate, the prevalence of PCO after surgery with PMMA lenses was 11.8% in the first year, 20.7% in the third, and 28.4% in the fifth year (Schaumberg et al., 1998). Nonetheless, the incidence of this complication is declining due to the advances in surgical techniques and the design and manufacturing of new lenses. PCO is treatable by capsulotomy with Nd:YAG (neodymium) laser. This laser is used to remove the central part of the posterior capsule of the lens to make the

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FIGURE 15.3 (A) The obstruction of the migration of lens epithelial cells in the round-edge posterior optic edge compared to (B) the sharp intraocular lens (Colvard, 2009).

vision path transparent and thus correct vision. This treatment is associated with serious ocular risks in 5% of cases, such as damage to the IOL, cystoid macular edema (CME), retinal detachment, and increased intraocular pressure (IOP). All these complications are treatable, but not always successfully. Furthermore, this surgery is not economical and is the second-most expensive surgery in the US Medicare system after cataract surgery. Avoiding PCO is, thus, ethically, medically, and financially justified (Bozukova et al., 2010). PCO is a multifactorial complication whose occurrence can be affected by the patient’s age or disease, the surgical technique used and the IOL design and material (Huang et al., 2014). The activity and growth of the epithelial cells of the lens decrease in older age, and the incidence of this complication therefore becomes less common with age. In young people, if the epithelial cells remain in the capsular bag, PCO occurs even with a suitable lens implant and irrespective of the lens type. In pediatric cataract surgery, lensectomy and anterior vitrectomy are performed, so that this medium is removed and the proliferation of these cells is impeded. In surgical terms, the likelihood of this complication is increasingly reduced with the more careful cleaning and removal of the lens residue. In addition, in cataract surgery, the surgeon forms a circular vent in the anterior capsule of the lens; this process is called capsulorhexis. The larger is capsulorhexis, the less likely is the incidence of PCO, since the LECs become more concentrated on the anterior capsule and around the equatorial region of the lens. The further removal of the anterior capsule further reduces the likelihood of developing this complication. In terms of design, IOLs with rectangular optic edge reduce the incidence of PCO due to their optic tackiness and by reducing the migration of

15.2 Application of Biomaterials in Intraocular Lenses

the LECs. Nonetheless, sharp optic edges may also have flaws. In some cases, the implant of lenses with rectangular edges combined with a high RI have led to the persistent edge glare phenomenon (Holladay et al., 1999; Erie et al., 2001). A sharp-edged design makes the light rays refracted through the peripheral IOL more severely affect the peripheral retina. Round edges scatter the light rays onto a larger area of the retina, thereby, causing less glare (Colvard, 2009). A study conducted on 90 eyes implanted with lenses in the bag showed that the material with which IOLs are made significantly affects the migration of the LECs on the posterior capsule (Hollick et al., 1998). The incidence and onset time of PCO was investigated in 147 eyes with silicone lenses and 585 eyes with PMMA lenses. The incidence of PCO was reported as 27.9% in the silicone group, but only 7% in the PMMA group. Over a 4-month period, 65.9% of the participants in the silicone group and 28.6% of those in the PMMA group required capsulotomy (Milauskas, 1987). The vision output, percentage of PCO and rate of laser capsulotomy were assessed 1 and 2 years after the surgery in the groups with PMMA, silicone, and hydrogel IOLs. This specific type of hydrogel lens was associated with a higher incidence of PCO and laser capsulotomy compared to the PMMA and silicone types (Hollick et al., 2000). The incidence of PCO was compared between a group with hydrophilic hydrogel and a group with hydrophobic acrylic IOLs and was found to be significantly higher in the hydrophilic hydrogel group (Hayashi and Hayashi, 2004). In a meta-analysis of 29 articles, despite the differences in the techniques used to determine the incidence of PCO, none of the articles showed a higher incidence of PCO with silicone lenses than with acrylic lenses. Silicone and acrylic IOLs have shown better PCO resistance compared to hydrogel lenses (Hayashi and Hayashi, 2007). The reason for the higher incidence of PCO in hydrophilic lenses appears to be their lack of tackiness (since water is basically not an adhesive material) which makes them not easily adhere to the capsule (Argal, 2013). These materials have a lower capsular biocompatibility and, therefore, a higher incidence of PCO (Huang et al., 2014).

15.2.6.2 Glistenings Glistenings are fluid-filled microvacuoles that appear inside the intraocular lens optic when the lens is placed in an aqueous medium (Werner, 2010). The hypothesis about the appearance of glistening is that water penetrates the micro-channels of IOL and minute accumulations of water form inside the micro-voids of the lens material (Rønbeck et al., 2013). The difference between the refractive indices of water and the lens material makes glistening detectable with a slit lamp. Water has a RI of 1.33 and its significant difference from the RI of the IOL material causes refraction and glare at the water polymer interface, which leads to the sparkling appearance of the fluid-filled vacuoles (Werner, 2010). The diameter of clinically observed glistenings varies from 1 to 10 microns; however, larger diameters have been reported by in vitro studies with dramatic temperature fluctuations (Rønbeck et al., 2013)

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Several clinical studies have been conducted to examine the effect of glistenings in IOLs on contrast sensitivity (CS) and visual acuity (VA), and some have stated that glistenings have significantly negative effects on high spatial frequency CS, while others have been nonconclusive. A few studies have shown that the increased severity of glistening reduces VA, but most studies have argued that VA is not affected by glistening (van der Mooren et al., 2013). The development of IOL glistenings may be affected by various factors such as IOL optic material, manufacturing and packaging, temperature fluctuations, glaucoma, breakdown of the blood-aqueous barrier, and the use of antiinflammatory and antiglaucoma eye drops. Previous studies are indicative of the appearance of glistenings in different optic materials, including PMMA, silicone, and hydrophilic and hydrophobic acrylics; however, this complication has been reported most in hydrophobic acrylics, especially AcrySof IOL (Rønbeck et al., 2013). In a 12-year follow-up, the degree of glistenings was compared in three different IOL materials, that is, PMMA, silicone, and hydrophobic acrylic. The results showed significantly higher glistening in hydrophobic acrylic lenses compared to lenses of other materials and the near lack of the complication in PMMA lenses (Rønbeck et al., 2013). In another study, CS and high-order aberrations were compared in PMMA, silicone, and acrylic lenses; the 10-year follow-up of the subjects after surgery showed that, in acrylic lenses, glistenings were significantly independent of visual function and optical aberrations.

15.2.6.3 Calcification Calcification has long been known as a characteristic of different implants in the human body. Given the presence of the blood eye barrier and the lower calcium concentration, the intraocular medium is considered a privileged medium, but is still not immune from calcification (Schoen et al., 1988). The IOL calcification occurring with lens optic opacification caused by calcium sediments has potentially serious complications that are rarely reported. Although hydrogel IOL calcification was first reported in 1987 (Buchen et al., 2001) many of its aspects, including etiology and pathogenesis, remain unclear to date. This complication is often associated with the loss of vision, and replacing IOL is the only way to improve the vision problems developed (Yu and Shek, 2001; Javadi et al., 2007), although the replacement of the lens may also be associated with complications such as tears in the posterior capsule and the detachment of the zonulae. The results show that calcification is initially a surface phenomenon, but calcium penetrates deeper into the lens over time. Calcification is, therefore, a multifactorial phenomenon with the lens material and the host’s intraocular medium as its components (Buchen et al., 2001). Cases of calcium sedimentation in silicone IOLs have been reported in patients with asteroid hyalosis (Wackernagel et al., 2004); however, the majority of the reports are associated with opacification and calcification after hydrophilic acrylic lens implants (Sinha et al., 2013; Bucher et al., 1995; Murray, 2000; Apple et al., 2000; Fernando and Crayford, 2000; Werner et al., 2000) Most lenses associated with this complication in the United

15.2 Application of Biomaterials in Intraocular Lenses

States include Hydroview (Bausch & Lomb), Memory (Ciba Vision), the SC60BOUV (Medical Development Research Inc.) and Aqua-Sense (Ophthalmic Innovations International Inc.) (Werner, 2010). Studies have shown that a high percentage of the patients experiencing hydrogel IOL calcification have an underlying disease such as glaucoma or diabetes (Yu et al., 2001; Pandey et al., 2002). The effect of the forceps remaining on the optic surface of the lens during its folding also affects lens calcification. In a study, the effect of forceps had remained on the optic surface of 96% of the lenses extracted from the eye (Yu et al., 2001). A study conducted by the manufacturer of Hydroview (Bausch & Lomb) blamed the presence of silicone in the lenses’ packaging for the tendency of calcium to deposit onto the surface of the lenses; silicone is gradually absorbed to the lens surface and, thus, acts as a catalyst in binding calcium and lipid to the lens surface (Rosenberg, 2004).

15.2.7 THE EFFECT OF DIFFERENT INTRAOCULAR LENS MATERIALS ON THE QUALITY OF VISION IN PSEUDOPHAKIC EYES The performances of PMMA, silicone, and foldable acrylic IOLs were compared in pseudophakic eyes. IOLs were implanted in 55 eyes with cataract using phacoemulsification. PMMA, silicone, and acrylic lenses were implanted in, respectively, 19, 2, and 16 eyes. Pseudophakic eyes with comorbid eye diseases were excluded from the study. VA, glare sensitivity, CS, and mesopicacuity were assessed 6 weeks after the surgery. The PMMA and acrylic groups showed better visual functioning than the silicone group. No statistically significant differences were observed between the acrylic and PMMA lenses in terms of the variables assessed (Kohnen et al., 1996). In another study, visual functioning was assessed and compared between PMMA and acrylic lenses. Pseudophakic eyes, including 50 eyes with foldable acrylic lens implants and 41 eyes with PMMA lens implants, were compared with 45 phakic eyes as the controls. The results obtained showed that, in terms of CS and disability glare, the quality of vision (QoV) is not as good in pseudophakic eyes as in phakic eyes; yet, foldable acrylic lenses performed better than PMMA lenses (Gozum et al., 2003). In another study, the QoV was compared in pseudophakic eyes with hydrophobic and hydrophilic acrylic lens implants; 28 eyes with hydrophilic acrylic lenses and 43 with hydrophobic acrylic lenses with clear posterior capsules were assessed at least 3 months after surgery in terms of Corrected Distance Visual Acuity (CDVA), total eye aberrations, and PSF values. The CDVA and PSF values were worse in the hydrophilic acrylic group. No differences were found between the two groups in the spherical equivalent; no differences were observed in terms of higher-order aberrations either, except in Z (2.0). According to the results, optic quality was better in the hydrophobic acrylic group (Nanavaty et al., 2011).

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15.2.8 DIFFERENT INTRAOCULAR LENSES MATERIALS IN CONGENITAL CATARACT SURGERY IN CHILDREN A number of studies have also been conducted on pediatric cataract surgery in order to determine the most suitable IOL material for children. Cataract is the most common treatable vision problem in children. Currently, lensectomy and posterior capsule removal and anterior vitrectomy are considered standard techniques for the treatment of congenital and developmental pediatric cataract (Faramarzi and Javadi, 2009). Because of the differences between children and adults in the size of the eyes, the more severe inflammatory response in children, the longer life ahead of children and the longer period for which IOLs have to remain in children’s eyes, IOL implantation in children became accepted as a safe method of aphakic correction for children many years after its use in adults; however, there is still a lack of consensus about the minimum age for initial implantation, the most suitable IOL power and the best lens material and design for children. Although some studies have reported the implantation of PMMA, silicone, and acrylic lenses in children, acrylic is currently the material of choice for pediatric IOLs (Bhusal et al., 2010). According to a study conducted by members of the American Association for Pediatric Ophthalmology and Strabismus (AAPOS), the majority of surgeons use hydrophobic acrylic IOLs for aphakic correction in children (Wilson and Trivedi, 2007). A recent study compared one-piece and three-piece hydrophobic acrylic lenses with PMMA lenses following lensectomy with posterior capsulotomy and anterior vitrectomy in the treatment of congenital and developmental pediatric cataract and found no significant differences between the three groups in terms of Best Corrected Visual Acuity; astigmatism a month after surgery was higher in the PMMA group compared to the others, but by the end of the follow-up period, no significant differences were observed between the three groups in this regard. Moreover, no significant differences were observed between the three groups in terms of postoperative uveitis, visual axis reopacification, and the need for laser capsulotomy. The researchers argued that, although in-the-bag foldable IOL implantation is the lens of choice for aphakic correction in children following lensectomy along with posterior capsulotomy and anterior vitrectomy, the implantation of three-piece hydrophobic and PMMA lenses is a safe and effective method when sulcus fixation is required (Doozandeh et al.).

15.2.9 ELIMINATION OF UV AND BLUE RAYS FROM INTRAOCULAR LENSES The energy of ultraviolet radiation (10 400 nm) is greater than the energy of visible light (400 700 nm) and infrared light (700 1200 nm) and, thus, has greater potential for biological damage. All radiations with a wavelength of less than 200 nm are absorbed by the air and most lights with wavelengths of 200 290 nm

15.2 Application of Biomaterials in Intraocular Lenses

are absorbed by the stratosphere. The main protection should, therefore, be against lights with wavelengths of 290 400 nm (Bozukova et al., 2010). The cornea absorbs wavelengths under 300 nm, while crystalline lenses absorb lights below 400 nm, and the retina and uvea absorb lights between 400 and 1400 nm. The light transmittance of the lens in human eye changes with age, so that the penetration of light reduces over time. In infancy, the human lens freely transmits near-UV light and visible lights with a wavelength above 300 nm, but yellow pigments and fluorogens develop in the lens as a result of exposure to UV radiation. At age 54, the lens can no longer transmit wavelengths under 400 nm and the transmission of lights between 400 and 500 nm is dramatically reduced. As the lens ages, its ability to filter near-UV light and blue light increases; therefore, the removal of the opaque lens in cataract surgery eliminates the natural protection provided by the aged lens. Concerns, therefore, arise about the degree of absorption of the implanted artificial lens and whether harmful UV radiations can reach the cornea or not. Exposure to intense UV radiation has been shown to reduce photoreceptor sensitivity with CME and age-related macular degeneration (AMD). To protect the cornea, it is recommended to implant IOLs with high UV light absorption (Bozukova et al., 2010). Since 1986, UV-blocking chromophores have been incorporated into all commercially-available IOLs (Mainster, 1986). Examples of UV-blocking agents are provided in the study by Bozukova et al. (2010). Visible violet (400 440 nm) and blue (440 500 nm) lights are increasingly absorbed by crystalline lenses with age. The relatively high energy of blue and violet lights that can harm the cornea were neglected for a long period. Through the implantation of UV-blocking IOLs, the cornea is exposed to greater amounts of blue light than before surgery (Bozukova et al., 2010). In 1991, the first blue light-filtering yellow IOL was introduced to the market in Japan by HOYA Crop in Tokyo. These lenses then became widespread and are currently identified as blue light-filtering (blocking) IOLs for their ability to absorb violet and blue lights. The study by Bozukova et al. discusses common blue light-blocking agents added to IOL materials (Bozukova et al., 2010). The light transmission range of blue light and UV-blocking IOLs may vary according to their base material and the filters used in them. IOLs with different filters are available in the market and all of them attempt to prevent blue light from reaching the cornea by creating cut-offs at different wavelengths in order to protect the cornea (Dı´ez-Ajenjo et al., 2014). Blue light-filtering lenses use yellow or orange-colored chromophores to achieve this goal (Brockmann et al., 2008; Mentak, 2010). The advantages and disadvantages of these lenses in terms of the QoV have been the subject of much debate. Most of the debates have been concerned with mesopic and scotopic vision, both of which are associated with vision in low light conditions. A reduction in scotopic vision is expected with the full elimination of blue light, since the dominant wavelength in these conditions is around 507 nm. One study showed that scotopic sensitivity is slightly reduced with blue light-filtering lenses, although this difference has been insignificant in

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terms of vision and the patients have likely gone without noticing (Werner, 2005). Yellow and orange-colored IOLs were compared in terms of color differences and contrast visibility in photopic and mesopic vision conditions; the results obtained showed no significant differences between the two lenses (Mun˜oz et al., 2011). Another study assessed the effect of lens color on optic quality and VA and determined optic quality using transfer function modulation and spectral transmission; the results showed that optic quality and VA were the same with both lenses, which suggests that lens color has no effect on the QoV in patients (Dı´ez-Ajenjo et al., 2014). In summary, it can be said that all the types of IOLs have advantages and disadvantages and that ophthalmologists must select the most suitable lens according to the needs of the patient. Table 15.3 shows the advantages and disadvantages of different materials IOLs. In summery it can be said that PMMA was the first material that IOLs were made from, but it is a rigid material and the inability of folding limits its use because it cannot pass through small incisions. The first material that foldable lenses were made of it, was silicone. This material is unfolded rapidly and is not the right choice for patients with silicone oil in their eyes or patients who may Table 15.3 Advantages and Disadvantages of Different Material Intraocular Lenses (IOLs) Poly (MMA) Advantages

Long-term experience Low cost

Disadvantages

Rigid, needs large incisions

Silicone Foldable—small incision Very low incidence of PCO Second generation of silicone with higher refractive index, thus thinner IOLs Rapid unfolding in the eye Cannot be used with silicone oil Slippery when wet Low refractive index of the first generation of silicone, thus thicker IOLs

Hydrophobic Acrylic

Hydrophilic Acrylic, Hydrogel

Foldable—small incision High refractive index-controlled unfolding Very low incidence of PCO

Foldable— small incisioncontrolled unfolding

High incidence of glistening High refractive index would give a shiny reflex from the pupil at night

High incidence of PCO Possible calcification

Adapted from Bozukova, D., Pagnoulle, C., Jérôme, R., Jérôme, C., 2010. Polymers in modern ophthalmic implants—historical background and recent advances. Mater. Sci. Eng. R: Rep. 69 (6), 63 83.

15.3 Artificial Cornea

require vitreoretinal surgery. Acrylic lenses are more manageable than silicon lenses and are used in microincision cataract surgery. High incidences of glistenings and dysphotopsia have been reported in association with hydrophobic acrylic foldable lenses. Some types of hydrophilic acrylic lenses are also associated with high rates of PCO and calcification. Hence, research into discovering the ideal material of IOLs must continue.

15.3 ARTIFICIAL CORNEA According to estimates by the World Health Organization (WHO) in 2010, 285 million people are visually disabled worldwide of whom 39 million are blind, and 12% of these blind (9.4 million) people are suffering from bilateral corneal blindness (Oliveira et al., 2011; Pascolini and Mariotti, 2010; Stevens et al., 2013). Diseases that affect the cornea are significant causes of blindness around the world (Al Arfaj, 2015). Globally, bilateral corneal opacity is the fourth leading cause of blindness after cataract, glaucoma, and AMD, which affects roughly 4 8 million people, of whom 90% live in developing countries (Akpek et al., 2014). Currently, the only effective treatment for advanced corneal disease is allogeneic corneal transplant to restore sight (Riau et al., 2015). However, the approach for using donor cornea often encounters several clinical, ethical, and logistical limitations such as global shortage of high-quality corneas, failure of allograft transplant due to immunological rejection, or endothelial decompensation as well as cases in which the transplantation of donor cornea is contraindicated (Griffith and Harkin, 2014). Besides, in many countries necessary infrastructures for transplantation are not available due to lack of storage facilities and religious or cultural issues (Riau et al., 2015). In addition, health care and financial constraints can be considered a major obstacle: For example, given the need for microbial testing, administration issues, and transportation, a donor cornea from an eye bank in the United States can cost around US$3000. As a result, a durable, safe, and cheap alternative, especially for developing countries, is required (Cruzat et al., 2013). Another limiting factor is the development of LASIK surgery procedures that compromises the biomechanics of stroma and, thus, makes the donor tissue undesirable for transplantation (Ruberti et al., 2011). These limitations, along with frequent unsuccessful transplants, which greatly increase the chance of rejection of the next transplants, have promoted research on finding a replacement for all, or part of, the cornea that is able to restore the structure and function of corneal tissue and also be compatible with the human eye (Chuo et al., 2011). There are generally two types of corneal substitutes: (1) keratoprosthesis; and (2) tissue engineered-corneal analogs (Duan et al., 2006). The long history of attempts to replace the human cornea using alloplastic materials dates back to at least 230 years ago. The term artificial cornea or keratoprosthesis refers to corneal implants made of plastic polymers or other synthetic

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materials (Akpek et al., 2014). However, recently, this term has been widely used for any implanted material that causes the restoration of the cornea structure and function (Griffith and Harkin, 2014).

15.3.1 HISTORY AND DEVELOPMENT OF KERATOPROSTHESIS The basic idea of taking advantage of the artificial cornea to replace opaque or damaged cornea was first published by a French doctor named Guillaume Pellier de Quengsy and coincided with the French Revolution (Salvador-Culla and Kolovou, 2016). His design consisted of a convex glass plate surrounded by a silver border (Avadhanam et al., 2015). The first era of using polymers in ophthalmic applications coincided with the observation of minor stimulation of the immune system by placing PMMA pieces into the cornea (Ruberti et al., 2011). After the discovery of the biocompatibility of PMMA during World War II, keratoprosthesis underwent remarkable development (Al Arfaj, 2015). Gradually, the two-section structure of core-skirt keratoprosthesis was developed by Choyce and Stone separately and, later, Dohlman and colleagues introduced the button-collar model made of PMMA (Avadhanam et al., 2015). Subsequently, several groups continued to improve, change, or replace keratoprosthesis with polymers and components of the core and skirt, but none of them showed promising results in the long term. The first generation of keratoprosthesis were unsuccessful due to its failure in stable biointegration into the host tissue caused by the impenetrability of PMMA applied in their design, causing subsequent complications such as extrusion, stromal melting, aqueous humor leakage, infection, retroprosthetic membrane (RPM) formation, retinal rupture, and glaucoma (Duan et al., 2006; Avadhanam et al., 2015). Since this generation of alternatives could not achieve the necessary conditions for integration, the second generation of engineered implants were developed including solid polymer corneal biomaterials and porous polymeric corneal biomaterials. Solid polymeric corneal biomaterials contain at least three solid polymer optical keratoprosthesis with basic clinical applications: Cardona keratoprosthesis, Boston keratoprosthesis (B-KPro), and Osteo-odontokeratoprosthesis (OOKP) (Ruberti et al., 2011). Considerable progress in keratoprosthesis in terms of design, materials, and reduced postoperative complications in order to increase optical quality, infection control, increased compatibility, and longevity has been achieved over the past 40 years. The material used to build the cylindrical optic support is a decisive factor of difference in keratoprosthesis designs (Sevgi et al., 2016). A wide range of materials such as metals, ceramics, glass, plastic polymers, and biological tissues have been tested to create skirts, plates, and haptic of artificial corneas; however, medical grade PMMA has only been used in the construction of the central optical core for some 50 years (Baba et al., 2015; Sarode et al., 2012). The most common synthetic materials that are used to build keratoprosthesis backplate contain PMMA, titanium, hydroxyapatite, PTFE, or hydrogel (Sevgi et al., 2016).

15.3 Artificial Cornea

15.3.2 BOSTON KERATOPROSTHESIS The most commonly used artificial cornea worldwide is Boston keratoprosthesis, formerly known as Dohlman-Doane keratoprosthesis. This keratoprosthesis was first proposed by Dr. Claes Dohlman from Massachusetts Eye and Ear Infirmary. Although B-KPro was designed in the 1960s, it only received FDA approval in 1992 and received certification marking in 2014 (Ruberti et al., 2011; SalvadorCulla and Kolovou, 2016). There are currently two main layouts of Boston-KPro. Both devices have the same collar and button shape, are among keratoprosthesis that perforate the cornea (Kadakia et al., 2008), and are designed as a doubleplated structure with a central hard PMMA optics which requires an allograft or autograft-carrier corneal tissue as the skirt. Because of the stiffness of PMMA used in these keratoprostheses, they are called hard KPro. Front plates comprising optical stem (hard central core) are made of PMMA and form the optic component. The backplate is a disc-shaped piece with a diameter of 5.8 mm, which has a large central aperture and 8 or 16 peripheral holes with small diameter (Avadhanam et al., 2015). Stem and front plates must necessarily be made of transparent material to allow the revision of the central carrier cornea, but as the most surface of the backplate is in front of the front compartment, this feature (being transparent) is not necessary (Todani et al., 2011). A titanium ring locks the backplate to the central stem. A donor cornea that plays the role of a carrier in the tool is in the form of a sandwich between frontal and the backplate, and the resulting complex is sutured to the patient’s eye (Ruberti et al., 2011; SalvadorCulla and Kolovou, 2016). There are a range of variable models of keratoprosthesis from fully synthetic to fully bioengineered models. Although several KPro tools have been developed and tested in small series over the past decade, in practice only a handful of them have had successful long-term results and still have routine clinical use. Among developed keratoprosthesis only four types are used commercially, including: B-KPro, OOKP, AlphaCor, and KeraKlear (SalvadorCulla and Kolovou, 2016). Here the improvements in the design of two types of keratoprosthesis which are widely used worldwide, that is, B-KPro and OOKP, will be mainly discussed because the preliminary results of the planting B-KPro and OOKP despite the short-term monitoring which has been performed on them—were promising and they have remained as the reference forms (Moussa et al., 2016).

15.3.2.1 Improvements over time Since the mid-1900s, signs of progress in the design of B-KPro and management of postoperative complications have played a significant role in obtaining the improved results. These include: (1) The use of broad-spectrum antibiotic prophylaxis in topical and daily forms as well as corticosteroids to reduce the amount of infectious keratitis and endophthalmitis. (2) The replacement of the solid backplate with a perforated plate (inclusion of 16 round holes with a diameter of each equivalent to 1.17 mm) on adult B-KPro backplate with a size of 8.5 mm and 8

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holes (each with a diameter of 1.3 mm) on the children B-KPro backplate (with a size of 7 mm) for better nutrition of stroma and keratinocytes of the donor cornea through aqueous humor, which is an essential factor in maintaining the integrity of the carrier cornea and reduction of tissue melting (sterile keratolysis). In addition, these holes are thought to be able to provide the liquid that has evaporated from the surface of the cornea and as a result allow the cornea to remain hydrated and prevent dryness that can lead to shrinking and subsequent leakage (Al Arfaj, 2015). (3) In 2005, replacement of PMMA with titanium as the main component of backplates was studied and seems to cause less postoperative reactivity than PMMA (Magalha˜es et al., 2012) and, thus, improves the biocompatibility and retention of the tool. (4) The use of a larger backplate to reduce the amount of RPM formation and subsequent melting of the corneal stroma (Akpek et al., 2014; Ruberti et al., 2011; Sivaraman et al., 2013). (5) In 2004, the idea of a titanium locking c-ring was suggested to prevent the separation of pieces inside the eye due to insufficient manual twisting. However, this system still presents several defects because twisting pages manually needs the back plate rotation that would cause extensive damage to the back transplant layers. (6) In order to prevent such damage and also facilitate the use of tools, a new design with a threadless stem (snug fit) was introduced in 2007 (Akpek et al., 2014). Change in the design of the instrument from threaded to threadless has made the assembly process easier and causes less damage to the endothelium of the donor cornea (Presanthila, 2015). Recent research has focused on the search for alternative materials in order to improve the results of the B-KPro application. Although PMMA is a transparent and biologically inert material with a long history of safe use in the eye, a few postoperative complications have been attributed to thick PMMA backplates. Todani et al. evaluated clinical results from a combination of stem and front plate made of PMMA with a backplate made of titanium (this model does not require a locking ring) and reported a decrease in the formation of RPM from 46.1% in eyes with threaded PMMA to 31.2% and threadless PMMA to 13% in the threadless titanium backplate. The latest design, which was approved by the FDA in 2013 for use in both type 1 and 2, has a modified backplate that eliminates the need for an integrated locking ring and can be made of PMMA or titanium (Akpek et al., 2014; Avadhanam et al., 2015). In another multicenter study in a 6month follow-up period after implanting B-Kpro, Todani et al. examined the effect of materials, which are utilized in back plate combination, on the level of RPM formation, known as a measure of inflammation. They found that 41.8% of eyes with B-Kpro and PMMA backplates, as well as 13% of eyes with B-Kpro with titanium backplates, showed the development of RPM. It seems that when titanium is used as a material for B-KPro backplates the RPM formation is less than PMMA (Todani et al., 2011). Restoration of vision in eyes with severe disease and deep vascularization of the cornea, limbal stem cell deficiency (LSCD), autoimmune diseases, and severe chemical burns which are prone to rejection is possible with the help of B-KPro (Al Arfaj, 2015). With the improvement of

15.3 Artificial Cornea

design and management of postoperative complications and by taking advantage of the mentioned strategies, B-KPro gained enormous popularity so that up to December 2015 more than 11,000 type I and about 200 type II implants were performed by 598 surgeons in 66 countries around the world (Salvador-Culla and Kolovou, 2016; Lee et al., 2016).

15.3.2.2 Outcomes of boston type-1 KPro In 1974, the first series of patients underwent type 1 B-Kpro surgery (Al Arfaj, 2015). Type 1B-KPro has a retention rate of 90% in nonimmunological disorders and 50% of those with autoimmune diseases associated with the mucous membranes (Lansini and Kwitko, 2015). The first and largest multicenter study on the type 1B-KPro was published in 2006 by Zerbe et al. In this study, a total of 141 type I B-KPro implants by 39 surgeons at 17 different locations was conducted from January 2003 to September 2005. Most patients had a significant improvement in VA such that this parameter in 57% of cases improved to 20/200 after surgery. Transplant survival rate after an 8.5-month monitoring period was 95%. In those with chemical injuries, VA of 6/60 had finally been maintained in 94% of eyes (Zerbe et al., 2006). Other recent studies suggest 83% 100% transplant retention rate, with visual results of at least 20/200 in 77% 83% of patients (Chuo et al., 2011). In one cross-sectional study, 33 eyes from 26 patients underwent type IB-KPro surgery. BCVA in 12 people had improved to at least 1. Keratoprosthesis retention rate was 85.71%. They verified the significant impact of their VA improvement after planting B-KPro on improving their quality of life (Lansini and Kwitko, 2015). Ciolino et al., in a multicenter study, investigated retention rate of type I and effective factors in prosthesis extrusion in 300 eyes. The probability of retaining the prosthesis after one and two years was 94% and 89%, respectively. The reasons for prosthesis rejection included keratolysis, fungal infections, bacterial endophthalmitis, and dense retroprosthesis membrane (Sinha et al., 2014). In another multicenter study, Rudnisky et al. reported the results from an investigation of logarithms of the minimal angle of resolution (logMAR) after type 1 implant in 300 patients from 2003 to 2008. After 17 months, VA improved significantly to 20/150. There were 19 eyes with light sensation, although the absence of light sensation was developed in 9 eyes (Rudnisky et al., 2016). Magalhaes et al., in a prospective study, reported the results of type I implanting in the treatment of injuries caused by eye burns. After a follow-up period 25.7 6 10.8 months after surgery, VA improvement in 90% of the cases was better than 20/200 and in 60% of the cases was better than 20/60. The overall rate of prosthesis survival was 90%. A significant breakthrough has been represented by Nallasamy et al. via planting KPro in children with amblyopiogenic conditions (Nallasamy and Colby, 2010). Keratoprostheses-related complications such as infection, glaucoma, chronic inflammation, retinal detachment, necrosis, and tissue melting are normally frequent and severe, and often lead to the loss of the eye (Dohlman et al., 2014). In an effort to reduce and eliminate these complications, Salvador-Culla et al. (2016)

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designed a miniature model of B-KPro called m-KPro to be implanted in rodents which enables the study of host immune response including the level of proinflammatory markers such as cytokines and collagenase after implanting the tool and comparing results with proinflammatory markers after penetrating keratoplasty. The results emphasize the use of strategies with the aim of reducing the allogenicity of carrier tissues, such as the use of syngeneic or acellular corneal tissue (Crnej et al., 2015). The researchers also assessed three ways for reducing the progression rate of glaucoma after surgery: (1) creating a shunt glaucoma using ferrofluids and magnetism; (2) using a contact lens for delivering drugs to treat glaucoma and reduce its complications, which can release a fixed amount of Latanoprost for at least 28 days in the anterior chamber of eye; and (3) eliminating the limitations of measuring IOP as the prognosis for glaucoma with the introduction of a novel wireless IOP transducer (Salvador-Culla and Kolovou, 2016). In order to solve the problem of tissue necrosis and melting, the Salvador-Culla team used new materials such as hydroxyapatite and titanium to produce stronger attachment between the cornea and tools. On the premise that changing the stem surface of B-KPro optics can increase the biological integration between the BKPro and carrier donor cornea and, therefore, reduce the risk of keratosis and endophthalmitis, the use of sandblasted titanium sleeves significantly increased the attachment of B-KPro to rabbit cornea. It is expected that titanium-based materials, unlike PMMA, have surface roughness and are able to improve biointegration of donor cornea in humans (Salvador-Culla et al., 2016).

15.3.2.3 B-KPro type II Type II is a revised design of the first type and is similar to it, except that type II has a longer optical stem that is the result of a 2 mm nub in the anterior plate that has been designed to penetrate between the eyelids or along the upper eyelid using a surgery method with closed eyelid (Dohlman et al., 2006). This makes BKPro type II beneficial for patients with severe ocular surface disease, extreme dry eye, imperfect blinking, chemical burns, ectropion eye pemphigoid, autoimmune, and inflammatory diseases such as SJS that do not have the necessary space to support type I (Ruberti et al., 2011; Sevgi et al., 2016). Lee et al., in the largest (to date) single-center study with the longest average follow-up period (70.2 months) on type II B-KPro performed on 48 eyes in Massachusetts Eye and Ear Infirmary between 1999 and 2015, showing that after a short monitoring period of 5.9 months, the improvement in VA of 91.7% of eyes was at least 20/ 200 and in 75% at least 20/50. In the following, longer follow-up period of 70.2 months, VA in 37.5% of people improved to 20/200 and 33.3% progressed to 20/ 100. The prosthesis retention rate was 50% (Lee et al., 2016). Results of the study by Pujari et al., which was conducted on 29 eyes with mucous membrane pemphigoid, SJS, and other ocular surface diseases, showed that VA in 23 eyes that underwent type II KPro surgery improved to 20/200 and in 10 eyes improved to 20/30. After a 5-year monitoring period, 6 eyes of 13 eyes gained VA of 20/200 (Pujari et al., 2011).

15.3 Artificial Cornea

15.3.3 OSTEO-ODONTO-KERATOPROSTHESIS OOKP methodology was originally founded by an Italian surgeon Strampelli in 1963 with the primarily aim of vision retrieval in people with impaired cornea (Chuo et al., 2011). However, the weaknesses of this approach have been mitigated by Falcinelli over the past 40 years delivering an outstanding performance, known as modified OOKP (M-OOKP) (Akpek et al., 2014). The method underwent more reforms which are referred to as the Rome-Vienna Protocol (Liu and Grabner). OOKP is an autograft in which the tooth tissue covered with autologous oral mucosal cells is used as the interface, hence, it is also known as “a tooth for an eye” or “tooth in eye surgery” (Griffith and Harkin, 2014; Sarode et al., 2012). OOKP surgery is a complicated combination of oral ocular processes in which a lamina propria of tooth root and buccal mucosa graft are used to keep a PMMA optical cylinder in the anterior part of the eye (Shetty et al., 2014). Biological coverage is vital for the survival of the alveodental layer because it provides blood flow for the bone and acts as a natural barrier against the entry of germs and environmental damage. Buccal mucosal membrane (BMM) is an ideal tissue for this purpose (Avadhanam et al., 2014). This tool is a nonpenetrating and semibiologic keratoprosthesis (a combination of a synthetic optics with the biological sensory part) in which a biocompatible host graft interface is used (Moussa et al., 2016). Hard tissues with slow metabolism such as cartilage, root canal dentin of teeth, and tibia or alveolar bones are widely used to produce the sensory part. The strong bond of above-mentioned part with PMMA optical cylinder leads to preservation against extrusion and fistula (Liu et al., 2005). To increase retention of the tool in the eye the tissue is obtained from the patient’s own body due to its similarity with the cornea, thus facilitating integration and better biocompatibility (unlike synthetic polymers). The main difference between the Boston-KPro and OOKP approach is the use of autograft material and a preserved bacterium between the optical center/hard tissue and blade-holder layer. In addition, the blood supply originated from the supporting tissue will help the optic support to remain healthy (Ruberti et al., 2011). The main advantage of OOKP compared to other types of keratoprosthesis is that it can withstand dry and keratinized eye surfaces (Liu and Grabner). The severe conditions such as opacity of the cornea, ocular surface inflammation such as SJS, chemical burns, trachoma, and dry eye could be only addressed through OOKP, the most effective treatment (Baba et al., 2015). The main design of OOKP consists of a PMMA optical stem and a singlerooted tooth with alveolar bone from the patient to be used as the biological skirt. The optical component is implanted in the cornea that is supported by the tooth (Weisshuhn et al., 2014). Nails also can be used instead of teeth (onycho-keratoprosthesis), but the drawback of this variant is that hangnail is capable of growth inside the eye (Baba et al., 2015). Chondro-keratoprosthesis (skirt made of cartilage) is not currently used, and temprano-keratoprosthesis (skirt made of the tibia) in edentulous patients is used as a suitable alternative for OOKP (Avadhanam et al., 2015). Although autologous tissue is preferred, in certain instances allograft

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OOKP may be obtained from canines donated from a living relative who have similar human leukocyte antigen (HLA) with the patient. In these cases, despite matched HLA, due to a higher risk of attracting the lamina, the likelihood of rejection is more than other conditions (Gulati et al., 2015). The results of multiple studies endorse the superiority of OOKP to other keratoprostheses due to long-term sustainability along with the degree of vision recovery (Baba et al., 2015). The strength of bond between the stem and tooth plays a central role in the maintenance of OOKP. To the best of our knowledge, PALACOS® is the only material used for attaching optical stem to dentin (Weisshuhn et al., 2014). Weisshuhn et al. investigated the effect of three different adhesives to attaching cow tooth to PMMA optical cylinder. The result showed that universal resin cement establishes a stronger attachment compared to glass-ionomer cement and PMMA bone cement and, therefore, can be a potential alternative to acrylic bone glue in the preparation of OOKP (Weisshuhn et al., 2014). The attachment of the haptic portion to the corneal surface and the optic can confine retroprosthesis membrane formation. Tan et al. (2012) systematically evaluated eight case series including all OOKP surgeries performed from 2004 to 2011 at the researchers’ center. The sample size consisted of 4 181 eyes. In all studies, the anatomical survival of the OOKP was 87.8% (range: 67% 100%) after 5 years and in half of the people, improvement in VA was more than 6/18; the results of three series of studies suggest the retention rate was 81% (range: 65% 98%) in a 20-year period. Additionally, in 60% of eyes subjected to OOKP surgery, VA recovered to more than 6/18 (range: 46% 72%). In 10 years after surgery, there was no significant difference between the retention rate of OOKP and temprano-keratoprosthesis. Although OOKP showed better retention than temprano-keratoprosthesis (bone resorption is faster than teeth, so in tempranokeratoprosthesis extrusion occurs faster as well). A significant complication during surgery included vitreous hemorrhage and blinding in the long-term, glaucoma, and endophthalmitis in the range of 2% 8% (Tan et al., 2012). Endophthalmitis is a common complication after planting keratoprosthesis. The bioinert material, titanium, has a high biocompatibility and is one of the most suitable options for the synthetic skirt of OOKP and can withstand adverse conditions in the eye caused by bacterial infection. Several studies have confirmed the efficiency of stabilized titanium by antimicrobial agents in vitro. Antimicrobial peptides have broader spectra of activity and less microbial resistance than antibiotics. Tan et al. evaluated the capability of OOKP Ti-AMP implanted in rabbit eyes with corneal keratitis for preventing infection during surgery. AMP bactericidal capabilities for Staphylococcus aureus were comparable to antibiotic prophylaxis after surgery. The effectiveness of the method against Pseudomonas aeruginosa was weaker than S. aureus. This approach can be effective in improving the biocompatibility and antibacterial role of artificial cornea in patients with severe corneal disease (Tan et al., 2014). Overgrowth of BMM is one of the OOKP postoperative complications. In a group of 60 people, about 12 people showed this complication. Avadhanam et al. reported the successful prevention of

15.3 Artificial Cornea

recurrence of BMM overgrowth with the local application of mitomycin C (MMC), an antimetabolite antibiotic derived from Streptomyces caespitosus in four people after modification. In all cases, an overgrowth of BMM, after onetime MMC consumption was stopped and there was no need for further consumption (Avadhanam et al., 2014). Regarding the field of OOKP, the current studies aimed at not only creating synthetic analogs so as to replace dental lamina propria but also taking precise measurements of IOP in patients post transplantation. (Salvador-Culla and Kolovou, 2016; Baba et al., 2015). Huhtinen et al., evaluated three bioglasses with different biological activity for use as a substitute of the dental lamina propria for OOKP porous skirts. Cell culture results indicated the potential use of bioartificial porous glasses as OOKP skirts due to limited potential in stimulating inflammatory responses, the ability for induction of endogenous growth, and tissue support (Huhtinen et al., 2013). In another study by Tan et al., bioneutralsynthetic materials were compared with hydroxyapatite as a reference for use in keratoprosthesis. The study material included: titanium oxide, aluminum oxide, and yttria-stabilized zirconia. They found that the bacterial binding to these materials was less than hydroxyapatite. Induction of keratocyte proliferation and also cell binding by TiO2 was considerably more in comparison to other materials. Therefore, TiO2 is the most suitable substitute for use as keratoprosthesis skirts because it improves cell functionality and reduces bacterial binding. This leads to increased tissue integration and reduces the extrusion of the tool (Tan et al., 2011). One of the most challenging applications of corneal biomaterials are when, for instance, the cornea does not grow due to damage or disease. There are several requirements in order to design an alternative for cornea that can be derived from its functional roles: protection, transparency, formation of a nearly perfect optical interface, biocompatibility, and integration (Ruberti et al., 2011). There are two main merits why the core-skirt is more successful than other designs. Firstly, it is able to resist against biodegradability owing to the PMMA, and the second capability is to allow the surrounding tissue of the cornea to be fused with host tissue. However, according to the surface properties of PMMA which are undesirable for tissue adhesion, it has become a challenge to improve the integration of skirtPMMA. This, combined with the constant movement of the eyes, blinking, and detachment of surrounding tissue from the surface of PMMA, can provide the conditions for the low growth of the epithelium, bacterial infection, aqueous humor outflow, and possibly failure of the tool. Changes in the surface of PMMA can be very effective to create a surface that is more compatible and ready to engage with a corneal tissue. Good adhesion between the two different materials could eventually lead to reduced long-term complications associated with KPro (Riau et al., 2015). Riau et al., in their study, modified the surface of PMMA using four different combinations, using hydrogels made of collagen type I that mimic human corneal tissue, to evaluate the connection between hydrogel and modified surfaces of PMMA as KPro models. They show that d-CaP (dopamine

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followed by calcium phosphate) in addition to increasing the adhesion of PMMA to hydrogel, has longer durability than other models and can promote biological integration (Riau et al., 2015). Moreover, Wang et al. studied the effectiveness of this model in increasing integration of the corneal tissue by coating the surface of PMMA with hydroxyapatite (HAp) and demonstrated that coating implants with hydroxyapatite is effective in reducing the inflammation around the PMMA cylinder and can potentially be used in combination with keratoprosthesis (Wang et al., 2011).

15.3.4 CARDONA KERATOPROSTHESIS Considering the pioneering work of Cardona there was an apparent change in the approach to the second generation of biomaterials, despite the fact that the major optical materials like methyl methacrylate was similar to the materials that were utilized in the mid-1950s. Cardona, with the beginning of a series of successful studies in rabbit models that were reported in 1962, designed penetrating corneal implants made of pure methyl methacrylate. There are two designs for the Cardona system with very similar features to the design of B-KPro. Although the use of Cardona keratoprosthesis has been stopped, in its design it has several common features with B-KPro (Ruberti et al., 2011).

15.3.5 PINTUCCI BIOINTEGRABLE KERATOPROSTHESIS A synthetic biointegrable keratoprosthesis with a design similar to OOKP and nonpenetrating properties has a central PMMA cylinder surrounded by a skirt which is made of Dacron (Tan et al., 2012). In 1979, Pintucci et al. presented initial reports of the use of their keratoprosthesis in the human eye. In their reports, 20 eyes received this tool from 1987 to 1991 and were monitored over an average period of 58 months after surgery. The relative improvement was observed in the sight of all people, among them, 13 people maintained the improvements for over two years (Chuo et al., 2011; Gomaa et al., 2010). In 2006, Maskati conducted a study on 31 eyes with an average monitoring period of 2.3 years. The vision of 24 eyes was improved to more than counting fingers from the distance of 1.5 , so that they could operate independently. The vision of four eyes (13%) improved to 20/200 or better. But, the functional outcome was poor as only 6.4% of people obtained vision of 6/12. In addition, significant complication was observed in 12 eyes, and 7 eyes lost their sight (Maskati and Maskati, 2006).

15.3.6 KERAKLEAR (KERAMED) Another corneal prosthesis is a nonperforating foldable tool that is implanted into a pocket created from part of the corneal thickness using femtosecond laser through a hole with a diameter of 3.5 mm (Studeny et al., 2015). In 2009 it was CE marked, but currently has not been approved by the FDA. Shiuey et al.

15.3 Artificial Cornea

studied 19 patients who underwent KeraKlear surgery. Results showed that after 4 years the vision from 20/400 before surgery improved to 20/40 after planting KeraKlear. No cases of keratoprosthesis membrane, glaucoma, or endophthalmitis were observed. After this period, 89% of the prostheses were preserved. Currently, the use of this prosthesis is limited to research in America.

15.3.7 MOSCOW EYE MICROSURGERY COMPLEX IN RUSSIA Due to medium- and long-term complications this product has been removed from the production cycle (Salvador-Culla and Kolovou, 2016).

15.3.8 WHAT NEXT? Research to complete the design, reduce complications, and repeat the essential features of the cornea is progressing (Sevgi et al., 2016). Future plans follow in the wake of the inclusion of new materials in the combination of artificial corneas to finally achieve an ideal tool that is able to take advantage of the natural cornea by having the following features: excellent optical quality; reduced deviations and determined power (strength); high biointegration with eye tissue; to provide resistance to infection and durability for a long time in the patient’s eye; to have some functional features of the cornea like drug penetration; and make measurement of IOP possible (Liu et al., 2005). The current trend of studies of B-Kpro has been focused on finding tools for better evaluating the intraocular to diagnose and manage glaucoma, as well as alternative materials to improve the biocompatibility and maintenance of the prosthesis. Recent advances in OOKP mainly focus on improving biological skirts to prevent or reduce keratolysis and the absorption of lamina (Sevgi et al., 2016). Full keratoprosthesis has not yet been discovered, although every day we get closer to this goal.

15.3.9 RECENT TRENDS Multiple third-generation tissue engineering techniques have been developed to replace the function of the cornea. The initial corneal tissue engineering technology started with the works of Jim Zieske and colleagues (1994) and then progressed to phase I clinical trial and its current reports by the May Griffith group (Ruberti et al., 2011). These researchers used biosynthetic analogs of cornea extracellular matrix to replace the pathologic anterior cornea of 10 patients who had considerable loss of eyesight, to facilitate endogenous tissue regeneration without the use of donor tissue (Fagerholm et al., 2010). In line with the progress of synthetic cornea, the trend currently is toward the use of natural materials such as acellular cornea tissue and tissue substitutes made of corneal cells in vitro. Biologically inspired materials with a view that they contribute to the growth of endogenous tissue are used for implants on their own (Griffith and Harkin, 2014). Recently, acellular extracellular matrix of different target organs has been

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presented as a scaffold to regenerate the cornea. Some research groups have used cornea cells to build their extracellular matrix in laboratory conditions as an alternative to solid organ transplant. In addition, some other groups utilized macromolecules of the extracellular matrix or polymers that mimic the extracellular matrix function to create artificial cornea that stimulates endogenous cells to migrate from the host to the implants to promote corneal healing. Several artificial corneas inspired by nature and natural extracellular matrix macromolecules have been developed, including artificial corneas fabricated from biological materials, corneas from self-assembly, collagen-based constructs, silk fibroin-based constructs, and recombinant human collagen-based corneal implants (Griffith and Harkin, 2014). In conclusion, complete keratoprosthesis has not yet been developed, although every day we become closer to this goal. Significant progress in keratoprosthesis has been accomplished over the past 40 years. A wide range of materials such as metals, ceramics, glass, plastic polymers, and biological tissues have been examined to create the skirt, plate, and haptic of artificial corneas, but only during the past 50 years has medical grade PMMA been used in the construction of the central optical core. The most common synthetic materials that are used to build keratoprosthesis backplates contain PMMA, titanium, hydroxyapatite, PTFE, or hydrogel. In parallel with the improvement in synthetic cornea, the trend currently is toward the use of biologically inspired materials that can dramatically promote regeneration of cornea.

15.4 GLAUCOMA DRAINAGE DEVICES Eye diseases leading to elevated IOP and finally reduced visual field totally are known as high-tension glaucoma. Increased IOP (normal IOP is usually between 10 and 21 mmHg) can cause damage to the optic nerve and lead to permanent blindness. According to the US National Institute of Health (NIH), approximately 120,000 are blind due to glaucoma, which accounts for 9% 12% of all cases of blindness in the United States (Maleki et al., 2012). Glaucoma is a very common disorder of the eye affecting an estimate of over 60 million people around the world by 2010, and is considered the second-leading cause of blindness (Van Krevelen and Te Nijenhuis, 2009). Glaucoma is an optic neuropathy characterized by progressive degeneration of retinal ganglion cells (Samuelson et al., 2011; Tavares, 2014). This group of diseases known by changes in retinal nerve fiber layer and optic nerve head resulting from increased IOP reduce the visual field sensitivity (Kersey et al., 2013). Treatment of refractory glaucoma is possible only by reducing IOP with the help of medical and surgical therapies as well as cyclodestructive laser procedures. When treatment with medication or laser is unable to slow the pace of nerve damage caused by failure to provide the desired amount of IOP, surgery is suggested (Patel and Pasquale, 2010). In general, the

15.4 Glaucoma Drainage Devices

primary goal of glaucoma treatment is to reduce IOP between 20% and 50% of the initial pressure in which the damage has occurred in the nerve (Samuelson et al., 2011). Since glaucoma is treatable and visual impairment in this disease is irreversible, early diagnosis is crucial.

15.4.1 HISTORICAL PERSPECTIVE The most common glaucoma surgery is trabeculectomy (guarded filtering procedure). However, the complications associated with the procedure are one of the essential reasons for the search for alternative methods (Patel and Pasquale, 2010). More than a century has passed since it was found that the accumulation of fluid in the eye can be a cause of blindness. The hypothesis of redirecting aqueous humor flow from the anterior chamber to subtenons or supraciliary space goes back to that time. According to the documents, the first step toward this was with the use of a seton from horse hair to drain fluid from the eye and was reported by Rollet in 1907. Since then, several other materials were used such as silk, gold, platinum, tantalum, glass rod, and polythene tube, none of which were successful (Ayyala et al., 2014). In 1966, Molteno reported his pioneering glaucoma drainage devices (GDDs). All existing GDDs have been designed based on the concept of the basic idea of Molteno implant (Giovingo, 2014). After that, two other major changes were made to improve the GDDs, one of which was the use of the valve mechanism to create resistance to flow and, thereby, reduce the incidence of hypotony (Ayyala et al., 2014). In this regard, Theodore Krupin, in 1976, introduced a pressure-sensitive unidirectional valve. In 1993, Marteen Ahmed described a pressure-sensitive valve mechanism designed to open when IOP is 8 mmHg. The next major modification in the original design by Molteno was the increase in the surface area of the endplate to 270 mm2 using a doubleplate valve. Baerveldt, in 1992, introduced nonvalved silicone implants with a surface area larger than the previous version. Other changes to up 2004 were made in Molteno implant valves that are available currently with two sizes of 175 and 230 mm2. The material used in its construction was changed from polypropylene to silicon. The double-plate Ahmed implant also became available. Changes in surface area and materials used in endplates are effective solutions to increase the efficiency of the tool in reducing IOP (Patel and Pasquale, 2010). GDDs are means to reduce IOP in patients with glaucoma through the creation of an alternate pathway for physical filtration of aqueous humor into subtenon’s space or supraciliary space (draining of aqueous humor from inside the eye into a bleb behind the eyelid). This approach is effective in reducing the pressure on the optic nerve and so preventing further damage or blindness due to glaucoma (Lloyd et al., 2001). These devices are known by other names, such as tube implants, glaucoma tube shunts, glaucoma drainage implants, and seton. The American National Standards Institute (ANSI) prefers the term aqueous shunts for this group of devices (Minckler et al., 2006).

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15.4.2 FUNDAMENTAL PRINCIPLES OF GLAUCOMA DRAINAGE DEVICES All GDDs carry out aqueous humor drainage through small silicon tubes (with less than 1 mm in diameter) attached to a nonbiodegradable silicon or polypropylene explant plate (Mark, 2004). A flexible rubber tube is inserted into the anterior or posterior chamber and is connected to a plastic or silicon plate with a large surface that is secured to the posterior sclera and covered by the conjunctiva. The plate as a physical barrier prevents injury of the conjunctiva and provides a large surface bleb to the sclera (Van Krevelen and Te Nijenhuis, 2009). The main differences among the GDDs are the surface area of the endplate, shape, thickness, and material as well as the presence or absence of flow-restricting components such as a valve mechanism (Mark, 2004). Although, trabeculectomy is the frontline of glaucoma treatment, the use of GDD is on the rise in early glaucoma treatment. These devices have been aimed to reduce the complications of trabeculectomy and are used in cases such as neovascular or uveitic glaucoma, primary open-angle or secondary angle-closure glaucoma that have not been treated with other methods. The success rate of these devices ranges from 30% to 90% depending on the duration of follow-up and studied specific diagnosis (Chen et al., 2015). These devices are available in two basic designs, that is, nonvalved and valved GDDs. The nonvalved devices (restrictive types) include Molteno and Baerveldt with a tube without active resistance to regulate the aqueous humor flow. The valved devices (nonrestrictive types) include Krupin slit valve and Ahmed with a tap at the end of the tube to create a unilateral flow restriction to the plate. The nonvalved implants may be better in the long term, but they are essentially at higher risk of early postoperative hypotony (Maleki et al., 2012). Therefore, the advantage of the valved types is dependent on pressure and, thus, reduced risk of postoperative hypotony. They all have an extraocular reservoir. Single- or double-plate Ahmed and Molteno implants are available in the market (Van Krevelen and Te Nijenhuis, 2009). Despite the different types of these devices that have been developed, only four are listed—the Ahmed glaucoma valve (AGV), Baerveldt glaucoma implant, Molteno implant, and Krupin—and are used routinely by surgeons (Giovingo et al., 2014; Mark, 2004). All these devices have similar performance, but there are important differences in their effect on ocular pressure in the first weeks after surgery and long term, as well as healing of the eye around the shunt (Lloyd et al., 2001). Generally, treatment approaches approved by the FDA for glaucoma are Trabectome (NeoMedix Corp.), Aqueous Shunts, including: Ahmed (New World Medical), Baerveldt (Advanced Medical Optics), Krupin (Eagle Vision), Molteno (Molteno Ophthalmic), and Micro-Stents, involving: Ex-PRESS Micro Shunt (Alco), iStent Micro-Bypass (Van Krevelen and Te Nijenhuis, 2009) and other approaches which have not yet been confirmed.

15.4 Glaucoma Drainage Devices

Tube versus trabeculectomy (TVT) is a multicenter randomized clinical trial comparing the safety and efficiency of tube shunt surgery to trabeculectomy with MMC in patients requiring cataract or glaucoma surgery (unsuccessful trabeculectomy) (Gedde et al., 2012b). IOP Reduction: The results of this 5-year study showed that both methods could cause a significant and continuous reduction in IOP, so that the IOP level was similar during this period in both groups (4.14 mmHg in tube shunt group and 6.12 mmHg in trabeculectomy group). During the period of 5 years, a similar percentage of patients in both groups had IOP of 14 mmHg. Medical therapy: A significant decrease was observed on the need to medication in both groups. The need for mean number of antiglaucoma medications after surgery was similar between the two groups (4.1 in the tube shunt group and 2.1 in the trabeculectomy group). During the first 2 years after surgery, the tube shunt group had used considerably more pharmaceutical supplements compared with the trabeculectomy group. Treatment Outcomes: The tube shunt group leads to significantly less likely to fail compared with the trabeculectomy group and has a higher overall success rate. However, the overall success rate was similar between the two groups (Gedde et al., 2012b; Gedde et al., 2009). Early postoperative complications were higher in the trabeculectomy group, but late complications or cataract progression and blindness rate were identical in both groups (Gedde and Group TVTS, 2009; Gedde et al., 2012a). The results related to the incidence and outcomes of reoperation for glaucoma in both groups showed that the need for reoperation in the trabeculectomy group was higher than those in the tube shunt group. Thus, the reoperation was performed in order to manage the postoperative complications on one patient in the tube shunt group and five patients in the trabeculectomy group (Saheb et al., 2014). The findings of the TVT resulted in the growing use of tube shunts and, in many cases, was the preferred surgery. According to the reports of Medicare, 43% reduction in the trabeculectomy and 184% increase in tube shunt implantation occurred between 1995 and 2004. There is much debate over the efficiency of trabeculectomy compared to the shunt on the one hand, and the types of the shunt on the other hand (Giovingo, 2014; Gedde and Group TVTS, 2009). The PTVT (primary tube versus trabeculectomy) is a multicenter randomized clinical trial comparing the safety and efficiency of tube shunt implantation with 350 mm2 Baerveldt implant to trabeculectomy with 4.0 mg/mL MMC for 4 minutes. The tube shunt group consisted of 117 people and the trabeculectomy group, with 108 patients, were analyzed during a 1-year follow-up period. The failure rate between the two groups showed significant differences: 20% for the tube shunt group and 8% for the trabeculectomy group. Hong et al. conducted a literature review from 1969 to 2003 and compared the efficiency of five different types of GDDs including single- and double-plate Molteno, Baerveldt, Krupin, and Ahmed. They demonstrated that all five types of devices equally (even for varieties of Molteno) resulted in a significant reduction in the IOP. There was no significant difference among these devices in terms of follow-up period, the incidence of hypotony, reducing the number of medications,

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decreased postoperative blindness, and overall success rate; except for doubleplate Molteno that had a better success rate and a follow-up period less than the rest (about 12 months) (Ayyala et al., 2014).

15.4.3 TYPES OF GLAUCOMA DRAINAGE DEVICES Studies associated with the GDDs often have been focused on the two types of AGV and Baerveldt glaucoma implant; therefore, we next discuss mainly the efficiency, success rate, and safety of these two implants.

15.4.3.1 Ahmed glaucoma valve The AGV is the most commonly used GDD. The AGV is a silicone tube connected to a silicone sheet valve held in a polypropylene body (Ayyala et al., 2014). The devices are now available with a lumened silicone rubber tube and in both models of single- and double-plate as rigid polypropylene and flexible silicone rubber (Minckler et al., 2006). The valve mechanism consists of thin silicone elastomer membranes designed to be opened when the IOP is 8 10 mmHg. The success rate of AGV has been reported between 60% and 825 after 2 years and 49% after 5 years of follow-up with a failure rate of 10% per year (Kaya et al., 2012). The results of AGV made of silicon and polypropylene in children younger than 10 years with pediatrics glaucoma and after 2 years of follow-up revealed that silicon AGV works better compared to polypropylene AGV in controlling IOP and causes a greater reduction in IOP. The retention rate of silicon AGV is also longer and associated with fewer antiglaucoma drops, but the rate of complications between the two groups showed no significant difference (El Sayed and Awadein, 2013). Similar studies in adults have reported better or equal results. Silicone creates less inflammation and fibrosis compared with polypropylene. It is thought that less encapsulation around the explant will cause lower resistance to fluid flow, resulting in better control of IOP and prolonged success of devices (El Sayed and Awadein, 2013). The GDDs, such as AGV, have a longterm and moderate success rate in children with glaucoma. In pediatric patients, the first GDD leads to success in 46% 70% of patients after 5 years of taking drugs, the second GDD is successful in 37% 75% of patients within 5 years after the surgery (Chen et al., 2015). The devices in children’s uveitic glaucoma treatment also have acted relatively safely, and their success rate can rise with longterm control of inflammation (Papadaki). Two multicenter randomized clinical trials compared the relative efficiency and failure rate of valved Ahmed Model FP7 with nonvalved 350 mm2 Baerveldt in treating refractory glaucoma. In the first study, or Ahmed Baerveldt Comparison (ABC), the comparison of Ahmed FP7 Glaucoma Valve (AGV) with Baerveldt 101-350 Glaucoma Implant (BGI) in 143 patients in the AGV group and 133 patients in the BGI group after 1-year follow-up period revealed that the mean IOP was slightly higher in the AGV devices. On the other hand, early postoperative complications were lower in the AGV group compared to the BGI group (Budenz et al., 2011). The results of this

15.4 Glaucoma Drainage Devices

study after 3-year follow-up showed that further adjuvant drugs are needed significantly after the AGV implantation to achieve the same success rate with BGI; thus, this leads to a higher relative risk associated with reoperation for glaucoma. In the BGI group, more patients experienced serious postoperative complications (Barton et al., 2014). Moreover, the success rate was similar for both implants during 5-years of follow-up. The BGI implantation was associated with a greater reduction in IOP as well as a lower level of need for reoperation for glaucoma compared with the AGV implantation, while the BGI implantation was associated with a higher failure rate caused by safety issues such as persistent hypotony and loss of light perception (Budenz et al., 2015). Therefore, this study highlighted the improved safety of AGV. Three strategies to prevent early hypotony include: (1) connecting the tube with biodegradable sutures and creating pores in the tubes during the procedure; (2) string suture inside the tube to provide a barrier; and (3) the use of nonvalved implants (Patel and Pasquale, 2010). The second study or Ahmed Versus Baerveldt (AVB), an international multicenter randomized trial to compare the efficiency of Ahmed FP7 and Baerveldt 350 mm2 in treating refractory glaucoma showed that both devices have been equally effective in reducing ocular pressure. However, the BGI group had a lower failure rate and were needed to fewer antiglaucoma medications after a 3year period. However, the BGI group was associated with higher serious complications, similar to the study results of ABC. The AGV reduces IOP more quickly and has a higher level of prognosis as well as being associated with fewer postoperative complications (Giovingo, 2014). Abe et al. evaluated the AGV efficiency in 108 patients with refractory glaucoma who had undergone the AGV surgery between 2000 and 2012. The results indicated AGV success rate of 50% after 5 years of implantation (Abe et al., 2015). HaiBo et al., during systematic review and meta-analysis, compared the efficiency and safety of AGV with trabeculectomy. Six controlled clinical trials were evaluated in this meta-analysis. There was no significant difference between the two approaches in reducing IOP, the number of drugs, antiglaucoma, the success rate, and the rate of the most common complications, although the AGV was considerably associated with a lower frequency of general complications (HaiBo et al., 2015).

15.4.3.2 Baerveldt glaucoma implants The Baerveldt glaucoma implants (BGI) is a flexible and silicone tube attached to a soft barium-impregnated silicone endplate with a surface area of 250 or 350 mm2 and with fenestrations in the plate (Ayyala et al., 2014). The BGI can be a safe and effective therapeutic approach to manage the failure to response of pediatric glaucoma to treatment. Rolim de Moura et al. used the BGI for the management of pediatric glaucoma. The results showed that the mean IOP was reduced to 4.16 mmHg after surgery from 2.31 mmHg before surgery. The cumulative probability of surgical success rate was reached to 58% after 48 months follow-up from 95% after 6-months follow-up (De Moura et al., 2005).

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Iverson et al. compared the results of trabeculectomy and BGI in the eyes with chronic inflammatory glaucoma caused by uveitis and uncontrolled IOP to determine the approach for long-term control of IOP in uveitis resistant to medical therapy. The cumulative probability of failure rate after 5-years of follow-up was higher significantly in the eyes with trabeculectomy (62%) compared to the eyes with BGI (25%). The IOP reduction and pharmaceutical therapy were similar in both groups. The surgery with nonvalved GDD will be more likely effective compared with trabeculectomy in maintaining IOP control and avoiding the need for reoperation in the eyes with chronic inflammatory glaucoma (Iverson et al., 2015). Ortiz et al. introduced a modified method of BGI implantation called “Ortiz’ partial titrated ligature” technique for effective reduction of IOP in the first postoperative days (Arismendi et al., 2013). Ahn et al. have developed two novel membrane-tubes (MT) type glaucoma shunt device to minimize the risk of hypotony and other complications associated with the tube. These devices are called Finetube MT and MicroMT (with a microtube) containing a membrane reservoir made of expanded polytetrafluoroethylene (e-PTFE) and a silicone tube or an intraluminal stent (Ahn et al., 2016; Han et al., 2016). Currently, there is insufficient evidence to draw conclusions concerning significant differences between clinical outcomes of trabeculectomy or aqueous shunts in similar patients as well as the preference for a particular type of aqueous shunt compared with other types (Minckler et al., 2006).

15.4.3.3 Molteno Single-plate Molteno is a silicone tube attached to a polypropylene endplate of 135 mm2 (Ayyala et al., 2014). Molteno3 is a nonvalved type made of polypropylene instead of medical grade silicone used in Ahmed and Baerveldt.

15.4.3.4 Krupin slit valve Theodore Krupin introduced the first type of valved GDD in 1976, but the original design was without a reservoir. The device consists of a silicone tube with a slit valve attached to a silicone oval endplate and with a surface area of 180 mm2 (Ayyala et al., 2014). Krupin has a unilateral pressure-sensitive valve to prevent the drainage of fluid. The slit valve has been designed to open at a pressure of 11 mmHg and close at a pressure of 9 mmHg. The device has been set with the characteristics of resistance, and the required pressure is quite varied for opening and closing the valve. In some devices, the opening pressure of the valve is 11 14 mmHg and the closing pressure of valve is 2 mmHg (Ayyala et al., 2014).

15.4.3.5 Ex-PRESS mini glaucoma shunt: Ex-PRESS glaucoma filtration device This small device made of stainless steel is a GDD without extraocular reservoirs and is placed in the anterior chamber under the scleral flap and, thereby, rotates

15.4 Glaucoma Drainage Devices

the trabecular meshwork, and by conducting the aqueous humor forms a bleb covered by the perilimbal subconjunctival space (Salim, 2013). In general, little change has been created on the principles of trabeculectomy and GDD. Most recent advances have been limited to a set of procedures as minimally invasive glaucoma surgery (MIGS). Due to the high rate of complications and failure of current methods of glaucoma treatment (e.g., trabeculectomy and tube shunt implantation), there is a continuous search for a safer and more effective approach. Using modern methods of nonpenetrating and independent of the bleb, MIGS has been popular with the aim of providing an alternative to reduce IOP with fewer complications and shorter recovery time compared to conventional glaucoma surgery. These methods based on surgical procedure can be divided into two groups, that is, AB interno and AB externo. AB interno methods include trabeculotomy (Trabectome, excimer laser), trabecular micro-bypass (iStent), suprachoroidal stent (Cypass) intracanalicular scaffold (Hydrus), and subconjunctival implant (Aquesys). Externo AB methods include canaloplasty, Stegmann Canal Expander, and Suprachoroidal Gold Micro Shunt. Currently, the two methods of trabeculotomy and insertion of iStent more than others are used to reduce IOP in Great Britain (Resende et al., 2016; Azarbod et al., 2015), which are not within the scope of this chapter (Resende et al., 2016). Certain devices such as the Shocket, OptiMed, White shunt pump, and Joseph implants all lack substantial peer-reviewed literature and will not be discussed in this review. In upcoming studies regarding the management of glaucoma, there are challenges that should be evaluated by using GDDs such as material and size of the implant and approaches to reduce fibrosis. The next step in improving these devices is reduction in the failure rate. Therefore, understanding the reasons for failure can lead to an insight of the things that can be changed in order to increase the success rate. In order to hinder encapsulation, fibrosis and therefore the failure of devices, the focus would be on tool design, biomaterials, and other changes in plates whether anterior or posterior (Patel and Pasquale, 2010). Materials used in the construction of the plate should be bioinert. Silicone and polypropylene cause stimulation of the inflammatory response. It is necessary to search for substances that have a great impact on the success rate of the devices. The biomaterials such as Vivathane and PMMA have been tested, but they also failed to cause less inflammatory response than existing materials. Ex-PRESS creates high biocompatibility with the minimal inflammatory response. Titanium is one of the promising materials which can be used in these devices because of its low reactivity. Another option is to cover the current valves using compounds that reduce inflammation or bind collagen and white blood cells to the surface of the plate. One of the recent applications of GDDs is the conversion of their endplate to the antifibrotic slow-release drug delivery system. In order to test the effectiveness of this approach, Ayyala et al. investigated an approach using two different polymeric implants attached to the AGV endplate containing two drugs of MMC and 5-fluorouracil (5FU) on the eyes of rabbit models. One of these implants made of a nonbiodegradable polymer based on polyhydroxyethyl methacrylate (PHEMA)

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releases MMC for about 3 weeks. The second implant made of biodegradable polymer of PLGA (polylactic-co-glycolic acid) releases the MMC as a 3-day continuous and explosive initial period and then releases the 5FU in a slow-release period of 30 days. Three months after surgery, the assessment of bleb wall thickness around the plate represents lower wall thickness in the group with the second implant; no infection and cystic bleb formation were observed (glaucoma: 2014 research report). To avoid postoperative fibrosis caused by the formation of scar, a drug delivery system of CsA-PLGA (CsA: cyclosporine A) based on the GDDs has been designed to release the CsA as sustained-release (Dai et al., 2016).

15.5 INTRACORNEAL RINGS Intracorneal rings (Intacs) are made from biomaterials and implanted to stroma using various methods. The effectiveness of intracorneal rings for reducing corneal steepening and improvement of refractive errors are reported in some corneal disorders including keratoconus, post LASIK ectasia, and pellucid marginal degeneration. In general, five types of poly (methyl methacrylate) Intacs have been used for correction of myopia and also reducing the spherocylindrical error and cornea irregularities in corneal ectatic diseases; Ferrara intracorneal ring (Mediphacos Inc.), Bisantis segments, which comprises of intrastromal segmented perioptic implants (Opticon 2000 SpA and SolekoSpA), Intacs that currently is distributed by Addition Technology, KeraRing (ICR; Mediphacos, Minas Gerais, Brazil) which are similar to Ferrara ring segments and the only nonsegmented full-ring system, MyoRing (DIOPTEX) (Ertan and Colin, 2007; Pinero and Alio, 2010).

15.5.1 INTACS SEGMENTS Intacs segments consist of two semicircular PMMA pieces. Typically, each segment has an external diameter of 8.10 mm and an internal diameter of 6.77 mm. Each of these pieces has a circumference arc length of 150 degrees, a hexagonal transverse shape, and a conical longitudinal section. The refractive effect is modulated by the thickness (0.25 0.45 mm increments), and current designs have a predicted range of myopic correction from 1.00 to 4.10 diopters (D) (Barraquer, 1966). In recent years some developments have taken place in the design of Intacs, for instance, a new Intacs segment has been designed by Addition Technology which has an inner diameter of 6.0 mm and an oval cross-section shape. This Intacs type has two thicknesses available, 400 µm (for steep K-readings of 57.0 62.0 D and cylinder , 5.0 D) and 450 µm (for steep K-readings .62.0 D

15.5 Intracorneal Rings

and cylinder . 5.0 D). To the best of our knowledge, there are no published reports of these new segments. The first report on using Intacs for treatment of keratoconus was by Colin et al. (2000). Results of this study showed that Intacs technology could reduce corneal steepening and astigmatism associated with keratoconus, and was proposed as an additive surgical procedure for keratoconus management. This surgical option provided an interesting alternative approach aiming to delay if not to avoid corneal grafting in ectatic corneal disease. There are various reports of using Intacs for treatment of ecstatic corneal disorders which puts emphasis on visual and refractive properties of this type of PMMA-based implant (Pinero and Alio, 2010; Colin et al., 2000; Pinero et al., 2009).

15.5.2 FERRARA RING SEGMENTS The first attempt of using modified PMMA rings was conducted on experimental animals (rabbit) by Ferrara in 1986. This study was expanded for medical use as in corneal stromal tunnel construction for implanting the rings. In 1996, he replaced the single ring with two segments, each having 160 degrees of arc, and obtained improved results for high myopia (Ferrara dC, 1995). These rings are made of computer-lathed PMMA Perspex CQ acrylic segments. The Ferrara ring segments are being produced in two diameters: 6.0 mm for myopia up to7.00 D and 5.0 mm for higher degrees of myopia. Thickness of the segments differs from 150 to 350 mm. The internal and external diameters are 4.4 and 5.4 mm, respectively, for the 5.0 mm optical zone and 5.4 and 6.4 mm, respectively, for the 6.0 mm optical zone. The segment cross-section is triangular, and the base for every thickness and diameter is a constant 600 mm. The segments have 160 degrees of arc. In the case of Ferrara ring segments, the ring diameter and addition of tissue together determine the degree in which the cornea will be flattened. This result is due to the fact that in this technique, the addition of tissue to the cornea periphery results in its flattening. Therefore, increasing ring thickness by adding more tissue, and a smaller ring diameter, would result in better correction of myopia (Ferrara dC, 1995). In a retrospective study, Abu Ameerh et al. used Ferrara segments to treat 50 patients, 79 eyes, diagnosed with progressive keratoconus. Results of this study showed that an overall significant postoperative improvement in both uncorrected visual acuity (UCVA) and best spectacle corrected visual acuity (BSCVA) throughout follow-up visits. Moreover, results illustrated a significant decrease in spherical equivalent (SE) and keratometric readings (lower, higher, and the average) (Ameerh et al., 2012). In another study, Miranda et al. used intrastromal Ferrara ring segments in patients with severe keratoconus to evaluate safety and efficacy of this procedure. No patient had a loss of VA. Uncorrected VA improved in 28 eyes (77.78%), and best spectacle-corrected VA improved in 29 eyes (80.56%). In their conclusion,

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the authors stated that Ferrara intrastromal ring segments could be an alternative for the treatment of severe keratoconus (Miranda et al., 2003). Despite similarities between Ferrara and Intacs segments, they differ in some physical properties including arc length (150 degrees in Intacs vs 160 degrees in Ferrara) and cross section (hexagonal in Intacs vs triangular in Ferrara). But the most important difference between these two could be referred as the diameter of inserts. This difference may affect clinical results because an insert nearer the center could theoretically increase the applanation effect.

15.5.3 BISANTIS INTRASTROMAL SEGMENTED PERIOPTIC IMPLANTS In the first generation of segmented perioptic implants, the differentiating element was the number of implants. In the first phase, three procedures were tested: implantation of three segments of 120 degrees, four segments of 90 degrees, or six segments of 60 degrees, all having a circular section and 150 mm of thickness, placed in a 3.5 mm optical zone diameter; and, at present, implantation of four segments of 80 degrees, each with an oval cross-section, vertical diameter of 250 mm, and horizontal diameter of 200 mm. The only variable parameter is the amount of curvature of the inserts to obtain optical zone parameters of 3.5, 4.0, and 4.5 mm (Ertan and Colin, 2007; Kubaloglu et al., 2011).

15.5.4 MYORING In addition to “segmented” ring systems, there is also another full-ring system available. MyoRing ring is a 360 degrees continuous full-ring implant to be implanted into a corneal pocket for the treatment of myopia and keratoconus. The internationally patented device combines two a priori contradictory qualities: rigidity for the modeling and stabilization of the corneal shape after implantation and flexibility (shape memory) for the implantation via a small pocket entry to preserve the corneal biomechanics (Daxer, 2008; Daxer, 2010; Mahmood et al., 2011). Our previous study showed that the depth and visual outcomes of MyoRing implantation using mechanical dissection via PocketMaker microkeratome as againt Melles hook are comparable (Pirhadi et al., 2018). In a long follow-up study, Janani et al. evaluated the long-term outcome on implantation of MyoRing for management of keratoconus (Janani et al., 2016). Three years after surgery, significant improvement was observed in UDVA, CDVA, SE, and K-readings, and manifest sphere and cylinder results of this study showed that 81% of patients were moderately to highly satisfied 3 years after surgery. In a retrospective, comparative, cohort study we compare MyoRing implantation alone versus MyoRing implantation with previously corneal collagen crosslinking (CXL) in moderate and severe keratoconus patients (NOBARI et al.,

15.5 Intracorneal Rings

2016) After 1-year follow-up, results showed that there was no significant difference in UDVA, refractive astigmatism, spherical equivalent error, and keratometric values between two groups. Mean CDVA was better in MyoRing group than MyoRing CXL one. This study showed that MyoRing implantation alone is a safe and efficient method for keratoconus treatment. In another study including 54 eyes from moderate and severe keratoconus patients, we showed that MyoRing implantation using pocket maker microkeratome is a safe and effective procedure for the treatment of keratoconus (Mojaled Nobari et al., 2014)

15.5.5 KERARING KeraRing (Mediphacos, Belo Horizonte, Brazil) KeraRings were developed specifically for keratoconus management. In terms of design, configuration and width, there is no difference between KeraRings and Ferrara rings. The main option of this ring is a various length of the arc (90, 120, 160, and 210 degrees) which provides more astigmatic control for the surgeon. Geometrically each KeraRing segment has an internal diameter of 4.40 mm and an external diameter of 5.60 mm. One of the best properties of this type of ICRSs is its ability to modulate their characteristics in order to reach a personal ring for each individual. For example, the size of KeraRing is comparable with an exclusive structural building block with specific properties of action on the corneal dome. Improving the thickness will result in increasing the effect of the segments and altering the arc length can modify the sphere and cylinder adjustments. Changes in stages of the disease, pupil diameter, and corneal pachymetry can be compensated with different optical zones (Vega-Estrada and Alio, 2016). Coskunseven et al. reported the results of KeraRing implantation using a femtosecond laser (IntraLase Corp, Irvine, California, USA) in keratoconic patients. In their final report, the authors stated that KeraRing implantation led to UCVA and BSVA improvement in patients with keratoconus. Coskunseven et al. (2008). In a previous study, we approved the efficacy and safety of KeraRing 355 degrees implantation for the correction of keratoconus in 15 eyes of 15 patients. Six months after surgery significant improvement in UCVA and BSCVA was observed, and other parameters such as cylinder, sphere, SE, and keratometry readings were significantly reduced, we showed that KeraRing 355 degrees intrastromal corneal ring segment is very efficient and cost-effective for improving VA in nipple-type keratoconic corneas (Jadidi et al., 2015). The slit lamp photograph of KeraRing implantation in one of the patient is shown in Fig. 15.4. Generally, Intracorneal rings showed good outcomes in the treatment of keratoconus, PMD, and post LASIK ectasia. Two main advantages of these rings are: (1) there is no need for the removal of corneal tissue; and (2) the reversibility of corneal biomechanical parameters when the ring is explanted. Using an advanced method such as the femtosecond laser-assisted surgical method may increase the rate of treatment by intracorneal rings. It seems that cooperation

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FIGURE 15.4 A photo slit image from an eye with KeraRing 355 degrees implantation.

between researchers and engineers in biomaterials, chemistry, nanotechnology, biotechnology, and tissue engineering will lead to more efficient and costeffective implants for vision improvement in patients.

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Index Note: Page numbers followed by “f” and “t” refer to figures and tables, respectively.

A 2-Acetamido-N-chloroacetyl-2-deoxy-β-Dglucopyranosylamine, 322 Acetone, 357 358 Acetonitrile, 357 358 Acrylate, 57, 65f Acrylate-methacrylate, 65f Acrylate polymers, 254 255 Acrylic, 513 Acrylic monomer, 514 Acrylic pimaric acid (APA), 165 Activators and polymerization initiators, 221 Active targeting, 373 Agaricus bisporus (ABA), 311 Agarose, 485 Ahmed Baerveldt Comparison (ABC), 538 539 Ahmed glaucoma valve (AGV), 538 539 Ahmed Versus Baerveldt (AVB), 539 Allyl alcohol lactic acid (ALA), 117 Aluminum, 274 275 N-(2-Aminoethyl) methacrylamide (AEMA), 311 313 Amphiphilic triblock copolymers, 311 313 Annulus fibrosus, 191 192 Antimicrobial applications, 205 207 Apatite wollastonite (A W), 192 193 Applications in biomedical engineering, 134 137 Aqueous shunts, 535, 540 Aramid fibers, 297 Artificial cornea. See Keratoprosthesis Artificial implants, 147 150 Atrisorb, 392 Autoclaving method, 8 Azide-alkyne, 88 89

B Baerveldt glaucoma implants (BGI), 538 540 Bakelite, 109, 275 Barium titanate (BT) nanoparticles, 132 133 Base metal alloys, 276 Benzoin ethers, 74 75 Benzoyl peroxide (BPO), 222f, 277 Bilateral corneal opacity, 523 524 Biocompatibility, 30, 479 480 defined, 353 impact of roughness and wettability on, 30 32 Biocompatible polymers, 3 4

Bio-Dot device, 325, 335 336 Bioepoxy resins, research works based on, 161 163 Biomaterial, 217 Biomedical engineering, 177 Biomedical polymers, 2 5 Biomedical sensors, 58 Biomedical thermoplastic and thermosetting polymers, 7 17 Bioplastics, 485 486, 497 500 Biopolyethylene, 487 Biopolymer applications, polyacrylates for, 81 88 Biopolymers, 58 59, 89 90 in living tissue, 496 500 Bisantis intrastromal segmented perioptic implants, 544 Bis-GMA/TEGDMA resin, 135 Bismuth subcarbonate, 10 Bismuth trioxide, 10 Bis-N-phenylbenzoxazine derivatives, 114 Bisphenol-A (BPA), 161 163 Bisphenol-A-glycidyl-dimethacrylate, 66 67 Bisphenol A glycidyl methacrylate, 219 Bisphenol E cyanate ester (BECy), 129 130 Blue light photoinitiator, 78 79 Blue rays elimination from intraocular lenses, 520 523 Bone cement, 201 204 Bone grafting thermoplastic and thermoset polymers in, 136 137 Bone repair thermoset polymers and their composites in, 136 Boron nitride nanotubes (BNNTs), 419, 424 428 properties, 427 428 structure, 424 synthesis, 424 427 Boston keratoprosthesis, 525 528 B-KPro type II, 528 improvements over time, 525 527 outcomes of Boston type-1 KPro, 527 528 Bovine bone hydroxyapatite (BHA) reinforced UHMWPE composites, 195 196 Breast cancer cells, 398 Buccal mucosal membrane (BMM), 529 531 Bulk composites, 153 158, 155t Buserelin acetate, 369t 1,4-Butylene glycol dimethacrylate (BGDMA), 286

555

556

Index

C Calcification, 518 519 Calcitonin, 463 Calcium phosphate (CaP) nanoparticles, 136 137 Camphorquinone (CQ), 135, 222f activation of, 222f with amine photoinitiator system, 78 79 Carbohydrate-polymer amide bond formation, 324 Carbohydrate-polymer ester bond formation, 324 Carbon fiber (CF), 134, 155t, 156 157, 296 Carbon fiber fillers, 160 161 Carbon nanofibers (CNFs), 136 137, 203 204 Carbon nanotube (CNT), 163, 192 193, 196 197, 418 Carboxyl-terminated butadiene acrylonitrile (CTBN) rubber, 129 130 Cardona keratoprosthesis, 532 Casein, 486 Cassie-Baxter model, 31f, 32, 47 48 Cataract surgery, lenses used in, 510 511 Cationic polymerization, 80 CCK-8 assay, 163 165 Celluloid, 275 Cellulose, 486 Cellulose nanofibers (CNF), 153 156, 155t CF/flax/epoxy plates, 136 Chain-growth polymerization, 109, 111 Chemical ageing, 244 by hydrolysis, 257 259 by radiolysis, 259 Chemical grafting of thermoplastics, 34 Chemically cured PMMA, 280 281 Chemical vapor deposition (CVD), 34, 424 425 Chitin, 158 159, 159f Chitosan (CS), 15, 86 87, 155t, 158 159, 159f, 401, 485 graft copolymerization of, 86 87 Chondro-keratoprosthesis, 529 531 Clapeyron’s law, 246 248 Clodronate (CL), 394 Coating polymers, 484 Cohesive energy density (CED), 239 Collagen, 392 393, 486, 515 Collamer, 515 Colloidal neoglycoconjugates and glyconanoparticles, 328 334 Composite coatings, 155t, 158 160 Composite preparation techniques for applications in implants and fixation plates, 152t Compression molding technique, 181 182, 279 Congenital cataract surgery in children different intraocular lenses materials in, 520

Contact lenses based on PVA, 14 Continuous-wave (CW) lasers, 38 Controlled release systems, 317 318, 349 Conventional polyethylene (CPE), 495 Conventional thermopolymerization, 226 227 Copper-catalyzed alkyne azide coupling (CuAAC), 311 313 Corneal collagen crosslinking (CXL), 545 Corneal substitutes, 523 524 CORTOSS, 136 Coupling agents, 221 CPTi, 148 150 Critical stress intensity factor. See Fracture toughness (FT) Crosslinked glycoconjugates, synthesis of, 324 Crosslinked lectin sorbents, 341 343 Crosslinked neoglycoconjugates (CLGCs), 325 326, 341 342 Cryptococcus neoformans, 317 318 C-type lectin-like domain (CTLD), 310 311 C-type lectin receptors (CLRs), 310 311 Curing, 105 106, 111 114 Cyclosporine A (CsA), 372 373 Cynate ester (CE) composites, 134 N,N-Cynomethyl methylaniline (CEMA), 135 Cytotoxic T cells (CTLs), 368

D Dacron, 532 Dark reaction, 80 Debris, 495 Dental composites, 78 79, 222 227 particle size and distribution of fillers, 223 225 polymerization mode, 225 227 viscosity, 225 Dental engineering, 372 373 Dental resin composites, classification of, 224f Dentistry epoxy composite systems in, 160 161 thermoset metal polymer composites in, 134 135 Denture base material (DBM), 274, 289 290, 299 Deuterated chloroform (CDCl3), 454 455 Deuterated dimethyl sulfoxide (DMSO), 454 455 Dextran, 83 84 D-glass fibers, 297 298 Diabetes management, 13 Diacrylate, 65f Dialysis method, 358 Diamino dimethyl sulfone (DSS), 129 130 Diblock and triblock copolymers of polylactide and polyglycolide synthesis of, 452 454 Dichloromethane, 163 165, 358 359

Index

Diethyl toluene diamine, 113 Differential scanning calorimetry (DSC), 119 120, 492 Diglycidylether of APA (DGAPA), 165 Diglycidyl ether of bisphenol A (DGEBA), 113, 117, 129 130, 165 Diglycidylether of terephthalic acid (DGT), 165 DiMarzio’s equation, 237, 254 Dimethacrylate of polycaprolactone, 65f, 82 7,7-Dimethyl-2,3-dioxobicyclo[2 2 1] heptane-1carboxylic acid (CQCOOH), 78 79 Dimethylacetamide, 357 358 Dimethylformamide, 163 165, 357 358 Dimethylsulfoxide (DMSO), 357 358 3-(4,5-Dimethylthiazol-2-yl)-2,5diphenyltetrazolium bromide (MTT) assay, 167 168, 458 Dispersed filler composite blocks, 227 Divinyl ether (DIVEMA), 317 318 Docetaxel, 472 Dohlman-Doane keratoprosthesis, 525. See also Boston keratoprosthesis Dot-blotting, 325 Doxorubicin (DOX)-loaded PLGA NPs, 372 Dynamic light scattering (DLS) experiments, 321 322 Dynamic mechanical tests, 132 Dynamic mechanical thermal analysis (DMTA/ DMA), 122 124, 133

E E-glass fibers, 136, 297 299 Einstein’s formula, 225 Electrospinning, 183 184 Emulsion-based methods, 356 357 emulsification reverse salting out, 357 emulsification solvent diffusion, 356 357 emulsification solvent evaporation, 356 Epoxidized hemp oil (EHO), 117 118 Epoxy/CNF composites, 153 154 Epoxy composites, 145 artificial implants, 147 150 dental applications, 160 161 fixation plates, screws, and intramedullary nails, 150 153 general biomedical applications, 166 168 green composites, tribological characterization of, 153 160 bulk composites, 153 158 composite coatings, 158 160 research works in, for biomedical applications, 168, 169t shape memory polymers (SMPs), 163 166

Epoxy/PCL electrospinning process, 164f Epoxy resins, 127, 292 Ethylene, chemical structures of, 186f Ethylene-based polyolefins, 175 176 Ethylene butene copolymers, 198 Ethylene glycol dimethacrylate (EGDMA), 67 68, 286 Ethylene oxide (ETO) exposure, 8 Ethyl methacrylate (MMA), 135 European Society for Medical Oncology (ESMO), 485 Exenatide, 369t Exendin-4 (exenatide, EXT), 461 462 Expanded polytetrafluoroethylene (e-PTFE), 540 Ex-PRESS mini glaucoma shunt, 540 542 Extrusion, 182 183, 183f Eyring’s theory, 241

F Fabrication, 180 Fac-[Ir(ppy)3], photoredox catalysts, 90 Femtosecond (Fs) lasers, utilization of, 47 48 Ferrara ring segments, 543 544 Fick’s law, 250 252 Fictive temperature, 244 245 Filament winding, 184 Filler ratio, 228f, 240 Fillers, 127 128, 131 Finetube MT, 540 Finger joint implants, 200 Fixation plates, screws, and intramedullary nails, 150 153 Flexion and elastic modulus tests, 160 161 Flexural strength (FS), 131 132, 280, 282 285 Flory approach, 242 243 5-Fluorouracil (5-FU), 373, 541 542 Foldable hydrophilic acrylic (hydrogel), 514 Foldable hydrophobic acrylic, 513 514 Fourier transform infrared spectroscopy (FTIR), 113, 117 118, 454 Fourier Transform Infrared Spectroscopy technique, 158 159 Fox-Flory’s equation, 236 Fracture toughness (FT), 285 286 Free radical and cationic photopolymerizations, 80t “Free volume” theory, 253

G Gelatin, 84 86, 359 Gelatin methacrylate (Gel MA), 84 86 Gelatin methacrylate-based hydrogel formulations, 68 69

557

558

Index

Gel permeation chromatography (GPC), 455 Gel point, 107 109 Gene transfection and tissue engineering, 372 Genexol-PM, 463 465, 471 472 Glass-ceramic apatite wollastonite (A W), 192 193 Glass fiber reinforced epoxy composites, 134 Glass fibers, 297 300 properties of glass fiber reinforced denture base resins, 298 300 types and composition of glass fibers, 297 298 Glass transition, 122 123 Glass transition temperature, 5 6, 236 238 Glaucoma, 534 535 Glaucoma drainage devices (GDDs), 534 542 fundamental principles of, 536 538 historical perspective, 535 types, 538 542 Ahmed glaucoma valve (AGV), 538 539 Baerveldt glaucoma implants (BGI), 539 540 Ex-PRESS mini glaucoma shunt, 540 542 Krupin slit valve, 540 Molteno, 540 Glistenings, 517 518 Glycoconjugates (GC), 318 N-(2-Glyconamidoethyl)-methacrylamides, 311 313 Glyconanoparticles (GNPs), 314 Gold, 273 Graft copolymerization, 386 Graft photopolymerization, 63 64 Graphene, 419, 498 499 applications of, 499 properties of, 498 499 Graphene oxide (GO), 419, 421 424 properties, 423 424 structure, 421 synthesis, 422 423 Green composites, tribological characterization of, 153 160 Griffith’s equation, 242 Growth hormone, 369t Growth vapor trapping (GVT) approach, 424 425

H Halpin-Tsai equation, 240 Healthcare-acquired infections (HAIs), 205 207 Heat cured PMMA, 277 279 doughy stage, 278 initiation and activation, 277 polymerization stages of heat cured PMMA, 277 propagation, 277 278 rubbery stage, 278

sandy stage, 278 stiff stage, 278 279 stringy stage, 278 termination, 278 Henry’s law, 246 248 Hexagonal boron nitride (hBN), 203 Hexamethylene diisocyanate, 59 60 High-density polyethylene (HDPE), 175 176, 187 193 Highly crosslinked polyethylene (HXLPE), 495 High molecular weight polyethylene, 176 Hill coefficient, 327 Hip prostheses prototype, 488f Human leukocyte antigen (HLA), 529 531 Human mesenchymal stem cells (hMSCs), 372 Hummers method, 422 Hyaluronic acid (HA), 87 88, 361 364 Hybrid composites, 223 Hydrogels based on PVA, 14 chitosan-based, 86 87 Hydrophilic macromers, 58 Hydrostatic extrusion process, 182 183, 183f Hydroxyapatite, 148 149, 531 532 Hydroxyapatite nanorods (HANRs), 190 191, 203 204 1-[4-(2-Hydroxyethoxy)-phenyl]-2-hydroxy-2methyl-1-propanone (Irgacure 2959), 75 N-(2-Hydroxyethyl) acrylamide (HEAA), 311 313 Hydroxyethyl methacrylate, 514 2-Hydroxyethyl methacrylate, 66 68 Hydroxyethyl starch (HES), 12 Hyper branched epoxy (HBE) system, 133 Hyperbranched polyesters (HBP), 60 62 Hyperbranched polyurethane acrylate (HBPUA), 63 Hyperbranched-urethane acrylates (HB-UAs), 60 62

I Implantable collamer lens (ICL), 511 Infrared (IR), 117 Injection molding technique, 279 280 polymerization cycles, 280 In situ method for particle formation, 358 Insulin, 13 Intervertebral discs (IVDs), 191 192 Intracorneal rings (Intacs), 542 546 Bisantis intrastromal segmented perioptic implants, 544 Ferrara ring segments, 543 544 Intacs segments, 542 543 KeraRing, 545 546 MyoRing, 544 545

Index

Intraocular lenses, 509 523 cataract surgery, lenses used in, 510 511 in congenital cataract surgery in children, 520 elimination of UV and blue rays from, 520 523 implant positions for, 512 materials, 513 515 acrylic, 513 collamer, 515 foldable hydrophilic acrylic (hydrogel), 514 foldable hydrophobic acrylic, 513 514 poly methyl methacrylate, 513 silicone, 514 515 phakic lenses, 511 on postoperative complications, 515 519 calcification, 518 519 glistenings, 517 518 posterior capsular opacification, 515 517 on the quality of vision in pseudophakic eyes, 519 structure, 511 Intraocular pressure (IOP), 527 528 reduction, 537, 540 Irgacure 2959, 75 Irgacures (CIBA), 74 75 Isophorone diamine (IPDA), 113, 120 121 Isophorone diisocyanate (IPDI), 66 67, 135, 452

K Kambour’s theory, 255 KeraKlear (Keramed), 532 533 KeraRing, 545 546 Keratoprosthesis, 523 524 Boston keratoprosthesis, 525 528 B-KPro type II, 528 improvements over time, 525 527 outcomes of Boston type-1 KPro, 527 528 Cardona keratoprosthesis, 532 future, 533 history and development of, 524 KeraKlear (Keramed), 532 533 Moscow eye microsurgery complex in Russia, 533 osteo-odonto-keratoprosthesis (OOKP), 529 532 Pintucci biointegrable keratoprosthesis, 532 recent trends, 533 534 Ketopinic acid, 78 79 Korsmeyer-Peppas equation, 318 Kovacs Aklonis Hutchinson Ramos (KAHR), 245 246 Krupin slit valve, 540

L Lactic acid, 384 385 condensation and coupling of, 385 Lactide, 385 D-Lactide, 385 chemical structure of, 384f L-Lactide, chemical structure of, 384f Lactosylamine, 323 Langmuir isotherm, 249 Langmuir-type diffusion, 252 253 Lanreotide, 369t Laser surface texturing (LST), 34 35, 37f challenges and future trends, 48 49 components of LST setup, 36 process fundamentals, 35 36 processing parameters, 36 39 beam mode, 38 pulse energy, 39 pulse length, 38 39 wavelength of the laser radiation, 36 38 of thermoplastic polymers, 39 48 poly(etheretherketone) (PEEK), 39 40 poly(ethylene terephthalate) (PET), 45 46 poly(methyl methacrylate) (PMMA), 47 48 polycarbonate (PC), 40 42 polyethylene (PE), 44 45 polypropylene (PP), 42 44 ultra-high-molecular-weight polyethylene (UHMWPE), 46 47 LASIK surgery procedures, 523 524 Latex, 497 Lectin carbohydrate interactions, 311 313 Lectin-glycans (LGI), 310 311 Lectinology, 309 Lectins binding assays, 325 326 crosslinked neoglycoconjugate sorbents, binding properties of, 325 326 dot-blotting, 325 UV-visible absorbance measurements, 325 Lectin sensors, development of, 315f, 334 338 Lectin sorbent preparation, 316f Lens epithelial cells (LECs), 515 Leuprolide acetate, 369t Ligand-based targeting, 373 Light cured PMMA, 281 Light-mediated thermoset polymers, 57 light-sensitive vinyl monomers, 69 73 vinyl carbonate, 70 71 vinyl esters, 72 73 N-vinyl pyrrolidone, 69 70 mechanisms of light sensitization, 79 80 photoinitiators, 73 79

559

560

Index

Light-mediated thermoset polymers (Continued) 2,4,6-trimethyl benzoyl-diphenyl phosphine oxide, 75 78 camphorquinone with amine photoinitiator system, 78 79 Irgacure 2959, 75 polyacrylates for biopolymer applications, 81 88 chitosan, 86 87 dextran, 83 84 gelatin, 84 86 hyaluronic acid (HA), 87 88 polycaprolactone (PCL), 82 starch, 82 poly-acrylates/methacrylates, 65 69 as dental materials, 66 67 as hydrogel biomaterials, 67 69 polyurethanes, 59 64 radiation curing in surface modification, 63 64 shape memory polymers (SMPs), 64 synthetic routes to photocrosslinkable polyurethane polymers, 60 63 recent advancements and trends in lightmediated polymerizations, 88 90 Light sensitization, mechanisms of, 79 80 Linear low-density polyethylene (LLDPE), 176 Linear polymer, 111 112, 112f Liquid composite molding, 153 154 Lithium phenyl-2,4,6-trimethylbenzoylphosphinate (LAP), 77 78 Lithography, 33 Low-density polyethylene (LDPE), 175 176, 187 Lower critical solution temperature (LCST), 456 Luteinizing hormone-releasing hormone (LHRH) analogs, 364

M Macrofill conventional composites, 223 Macrosized topographies, 30 31 Maleated high-density polyethylene (mHDPE), 189 Maleic anhydride copolymers, 309 colloidal neoglycoconjugates and glyconanoparticles, characterization of, 328 334 crosslinked lectin sorbents, 341 343 instrumentation, 321 322 materials, 321 methods, 322 326 N-glycyl-β-glycopyranosylamines, synthesis of, 322 324 lectins binding assays, 325 326

neoglycoconjugates and glyconanoparticles, synthesis of, 324 325 neoglycoconjugates and metal-labeled glyconanoparticles, synthesis of, 326 328 silver (or gold)-labeled neoglycoconjugate, 334 341 lectin sensors, development of, 334 338 UV-visible absorbance spectroscopy, 338 341 Matrix, composition of, 220 221 activators and polymerization initiators, 221 coupling agents, 221 monomers, 220 polymerization inhibitors, 221 Means cluster size (MCS), 248 Medical devices, thermoset metal polymer composites in, 137 Medical grade UHMWPE in living tissue, behavior of, 492 495 Medical imaging techniques, 89 90 Melamine formaldehyde (MF) resins, 125 126 Melt electrospinning, 183 184, 184f Membrane-tubes (MT) type glaucoma shunt device, 540 Meso-lactide, chemical structure of, 384f Methacrylate, 65f, 220f Methacrylate networks, polymerization of, 233 234 mechanistic aspects, 233 polymerization kinetics, 234 Methacrylates for dental applications, 219 220 Methacrylates polymers, 262 Methacrylic anhydride (MLA), 121 1,6-bis(Methacryloxy-2-ethoxycarbonylamino)2,4,4-trimethylhexane, 220 2-Methoxyestradiol (2ME2), 371 372 Methoxypoly(ethylene glycol)-b-poly(lactide) (mPEG-PLA-SA), 461 2-Methyl ethylene, 199 4,4’-Methylene biscyclohexanamine (PACM), 113 Methylene diphenyl diisocyanate, 59 60 Methyl methacrylate, 218 Micelles, 402 Microfill composites, 223 Microfluidic technique, 359 360 Microincision cataract surgery (MICS), 514 Micromilling, 33 MicroMT, 540 Microparticles (MPs), 129 130, 403 Microwave curing PMMA, 281 282 Mid-infrared laser radiation (MIR), 37 38 Minicircle DNA (mcDNA), 402 Minifill composites, 223

Index

Minimally invasive glaucoma surgery (MIGS), 541 Minocycline, 369t m-KPro, 527 528 Modulus of rupture. See Flexural strength (FS) Molding, 181 182, 182f Molecular functionality, defined, 111 112 Molecular imprinting, 371 372 Molteno, 540 Mono-acyl phosphine oxide (MAPO), 227 Monomers, 218 221 matrix, composition of, 220 221 activators and polymerization initiators, 221 coupling agents, 221 monomers, 220 polymerization inhibitors, 221 methacrylates for dental applications, 219 220 methyl methacrylate, 218 Moscow eye microsurgery complex in Russia, 533 MP5510, 137 μ-TiO2 particles, 149 150 “Multicopy-multivalent” nanoscale glycoconjugates, 311 313 Multiple methyl methacrylate (MMA) molecules, 277 Multiwalled carbon nanotubes (MWCNT), 136, 147 148 MyoRing, 544 545

N NAD(P)H-dependent oxidoreductace, 151 Naltrexone, 369t Nanobiomaterials, 497 500 Nanoclay, 130 Nanofilled composites, 223 224 Nanofillers, 130 131 Nanohydroxyapatite (nHA), 195 196, 203 Nanoplate fillers, 127 128 Nanoprecipitation method, 357 358 Natural biopolymers, 484 Naturally derived biodegradable polymers, 3 4 Natural polymers, 58 59 N-carboxyanhydride (NCA), 311 313 Neoglycoconjugates and glyconanoparticles, synthesis of, 324 325 carbohydrate-polymer amide bond formation, 324 carbohydrate-polymer ester bond formation, 324 crosslinked glycoconjugates, synthesis of, 324 silver/gold glyconanoparticles, synthesis of, 325 Neoglycoconjugates and metal-labeled glyconanoparticles, 326 328 N-glycyl-2-actamido-2-deoxy-β-Dglucopyranosylamine (NGly-GlcNAc), 322

N-glycyl-4-O-β-D-galactopyranosyl-β-Dglucopyranosylamine (N-Gly-lactose), 322 324 N-glycyl-β-glycopyranosylamines, synthesis of, 322 324 N-methyl-2-pyrrolidone (NMP), 357 358 5-Norbornene-2,3-dicarboxylic anhydride, 118 119 5-Norbornene-2-yl(ethyl) chlorodimethylsilane, 118 119 Normal human dermal fibroblasts (NHDF), 440 Nuclear magnetic resonance (NMR), 118 119, 454 455 Nucleotides, 485 Nylon 6,6, 48 Nylons. See Polyamides

O Octreotide, 369t OncoGel, 465 Organomineral fillers, 223 Organoorganic fillers, 223 Osseointegration, 148 149 Osteolysis, 493 494 Osteo-odonto-keratoprosthesis (OOKP), 529 532 1,4-bis[(2-Oxiranylmethoxy)methyl]-benzene (BOB), 113 2,5-bis[(2-Oxiranylmethoxy)methyl]-furan (BOF), 113

P Paclitaxel (PTX), 361 364, 463 465, 471 472 PALACOS®, 529 531 PAN-based fibers, 161 Particle formation using template/mold, 359 PEG-grafted boron nitride nanotubes (PEG-gBNNTs), 419 PEGMAs, 63 64 PEGylation, 12 13 Pentaerythritol (PTOL), 453 454 Pentaerythritol functionalized with methacrylic anhydride (PMLA), 121, 122f Peri-implantitis, 148 149 Periodontitis, 392 Perspex, 513 Phakic lenses, 511 Phenol formaldehyde, 275 Phenol formaldehyde resin, 126 Pheroid system, 373 Phosphate buffered saline (PBS), 433 Photoclick reactions, 88 89 Photocrosslinkable modified polymers, 83f

561

562

Index

Photocrosslinkable poly(ethylene glycol) acrylate, 68 69 Photocrosslinkable polyurethane polymers, synthetic routes to, 60 63 Photocurable bioadhesives, 57 58 Photocurable endoprosthesis systems, 58 Photocuring, 73 74, 80 Photocuring urethane acrylate, 61f Photofabrication techniques, 89 90 Photoinitiators, 59, 66, 73 80, 91 camphorquinone with amine photoinitiator system, 78 79 Irgacure 2959, 75 2,4,6-trimethyl benzoyl-diphenyl phosphine oxide, 75 78 Photolithography, 33, 58 Photopolymerization, 226 227 Photopolymerized thiol-ene hydrogels, 88 89 Physical ageing, 244 Pintucci biointegrable keratoprosthesis, 532 Plasma- and chemical-based etching, 33 Plastic-based biomedical materials and devices, 3, 9 10 Plastic polymers, 9 10, 109 Plexiglas, 481, 513 Pluronics, 459 Poisson’s ratio, 241 Poly-1 methylpentene (PMP), 176 Poly (2-ethyl-2-oxazoline)-poly(L-lactide), 402 Poly(2-hydroxyethylmethacrylate) (PHEMA), 67 68, 191 192 Poly (ADP-ribose) polymerase (PARP), 485 Poly(caprolactone) (PCL), 191 192 Poly(D,L-lactic-co-glycolic acid) (PLGA), 449 450, 541 542 PLGA-E100 nanoparticles, 469 470 PLGA-PEG-PLGA triblock copolymers, 453 Poly(ε-caprolactone) (PCL), 417 418 Poly(etheretherketone) (PEEK) laser surface texturing (LST) of, 39 40 Poly(ethylene-alt-maleic acid) (EM), 321 Poly(ethylene-alt-maleic anhydride) (EMA), 321 Poly(ethylene glycol)-block-poly(propylene glycol)-block-poly(ethylene glycol), 399 Poly(ethylene glycol) diacrylate, 68 69 Poly(ethylene glycol) dimethacrylate, 86 Poly(ethylene glycol) methacrylate, 68 69 Poly(ethylene terephthalate) (PET), 191 192 laser surface texturing (LST) of, 45 46 Poly(glycolic acid) (PGA), 3 4, 417 418, 484 Poly (hydroxyethyl methacrylate), 67 68 Poly(lactic acid) (PLA), 349, 484 advancements, 368 373 active targeting, 373

cellular interaction, 371 372 dental engineering, 372 373 gene transfection and tissue engineering, 372 pheroid system, 373 super paramagnetic iron oxide nanoparticles (SPIONS), 371 vaccination, 368 370 biocompatibility and safety of, 353 354 biological cycle of, 351f challenges with particulate system, 360 clinically used products of, 369t copolymer particles of, 366t future perspectives, 374 preparation methods of PLA micro-and nanoparticles, 355 360, 355f dialysis, 358 emulsion-based methods, 356 357 in situ method for particle formation, 358 microfluidic technique, 359 360 nanoprecipitation method, 357 358 particle formation using template/mold, 359 spray drying, 358 supercritical fluid (SCF) technique, 358 359 production of, 350 352 products under clinical use, 368 products under preclinical and clinical trial, 360 367 unique properties of, 352 353 Poly(lactic-co-glycolic acid) (PLGA), 349, 356, 361, 372 Poly(L-lactic acid), 417 418 Poly(L-lactide), 16 Poly(methyl 2-methylpropenoate), 508 Poly(methyl methacrylate) (PMMA), 68 69 chemical modification of, 293 chemical structure of, 509f contemporary denture base materials and modifications of, 291 292 epoxy resins, 292 polyamides, 291 292 polycarbonates, 292 laser surface texturing (LST) of, 47 48 properties of, 282 290 biocompatibility and cytotoxicity, 290 color stability, 289 crosslinking, 286 flexural strength (FS), 282 285 fracture toughness (FT), 285 286 impact strength, 286 radiopacity, 289 290 residual monomer, 288 sorption and solubility, 286 288 thermal conductivity, 288 radical polymerization reaction of, 228 233

Index

kinetic aspects, 230 233 mechanistic aspects, 228 230 Poly(methyl methacrylate) (PMMA) denture base materials, reinforcement of, 293 300 different types of fibers used in dentistry, 296 300 aramid fibers, 297 carbon fibers (CFs), 296 glass fibers, 297 300 polyethylene (UHMWPE) fibers, 297 fiber reinforcement, 294 296 effect of fiber orientation, 295 effect of silane treatment on properties of PMMA denture base resins, 296 effects of fiber length on properties of fiber reinforced denture base resins, 294 295 effects of resin impregnation on PMMA resin-based materials, 295 296 metal wires/mesh, reinforcement with, 293 294 Poly(methyl methacrylate) (PMMA) resins, classification of, 276 282 according to ISO standards, 276 according to method of polymerization, 277 282 Poly(N-isopropylacrylamide) (poly-NIPAAM), 459 Poly(N-vinyl-pyrrolidone-alt-maleic acid), 321 Poly(propylene fumarate) (PPF), 417 418 applications, 421 properties, 420 421 synthesis, 419 420 Poly(propylene fumarate) (PPF)-based bionanocomposites characterization of, 428 442 antibacterial properties, 437 440 biodegradability, 431 434 cytotoxicity, 440 441 hydrophilicity, 431 434 mechanical properties, 434 437 morphology and structure, 428 431 protein adsorption, 431 434 thermal properties, 434 tribological properties, 441 442 future perspectives, 442 443 preparation of, 428 Poly(styrene-alt-maleic anhydride), 318, 321 Poly(styrene-cobleck-4-vinylbenzocyclobutene)polyisobutylenepoly(styrene-coblock-4vinylbenzocyclobutene), 135 136 Polyacrylates for biopolymer applications, 81 88 Poly-acrylates/methacrylates, 65 69 as dental materials, 66 67 as hydrogel biomaterials, 67 69 Polyamide-6(PA-6)/UHMWPE composite, 196 Polyamides, 291 292, 487

Polybutylene (PB), 176 Polycaprolactone (PCL), 16, 82, 163 165 Polycarbonate (PC), 292 laser surface texturing (LST) of, 40 42 Polydimethacrylate, 66 Polyelectrolytes, 126 Polyesters, 486 Polyetheretherketone (PEEK), 147, 158 Polyethylene (PE), 8, 176, 297, 481 chemical structures of, 186f laser surface texturing (LST) of, 44 45 molecule alignment in different forms of, 186f typical average properties of different forms of, 188t Polyethylene glycol (PEG), 11 13, 352 353, 357 358, 419, 449 450 PEG-functionalized graphene oxide (PEG-GO), 419, 433 434, 436 437 Polyethylene matrix, 186 198 high-density polyethylene (HDPE)-based biomedical composites, 187 193 ultrahigh molecular weight polyethylene (UHMWPE)-based biomedical composites, 193 198 Polyethylene oxide (PEO), 12 13, 419, 452 Polyethylene terephthalate (PET), 484 Polyglycolide (PGA), 16, 449 450, 484 characterization of copolymers of, 454 459 aqueous solubility and injectability, 455 biocompatibility, cytotoxicity, and biodegradability, 457 459 crystallization behavior, 457 phase transition, 456 structural composition analysis, 454 455 thermal properties, 456 457 history of, 451 452 Polyhedral oligomeric silsesquioxane (POSS), 135 Polyhydroxybutyrate (PHB), 486 Polyhydroxyethyl methacrylate (PHEMA), 541 542 Polyimide (PI), 48 Polylactic acid, 381 385 antigen loading, 403 biomedical applications of, 391 408 DNA and gene delivery, 402 403 drug delivery with PLA particles, 393 395 immunization with PLA particles, 399 402 tissue engineering, 391 393 tumor treatment, 397 399 vaccine delivery, 395 397 imaging and diagnosis, 405 408 modification, 386 graft copolymerization, 386

563

564

Index

Polylactic acid (Continued) high energy radiations and peroxides, modification by, 386 physicochemical properties of, 387 391 biodegradation properties, 390 391 mechanical properties, 388 389 rheological properties, 387 thermal properties, 389 390 polymerization, 385 condensation and coupling of lactic acid, 385 precursors, 384 385 lactic acid, 384 385 lactide, 385 protein delivery, 404 405 Polylactide, 449 450 characterization of copolymers of, 454 459 aqueous solubility and injectability, 455 biocompatibility, cytotoxicity, and biodegradability, 457 459 crystallization behavior, 457 phase transition, 456 structural composition analysis, 454 455 thermal properties, 456 457 history of, 450 451 Polylactide MPs (PLA-MPs), 403 Polylactide/quercetin (PLA/Qt), 398 Poly-L-lactide (PLLA), 352 353 Polymer-based biomaterials, 4 Polymer grafting, 33 34 Polymeric fibers, 191 192 Polymeric materials used in medicine, 479 applications of biomaterials, 483 487 behavior of medical grade UHMWPE in living tissue, 492 495 bioplastics, present and future of, 497 500 biopolymers, present and future of, 497 500 biopolymers in living tissue, background on, 496 497 nanobiomaterials, present and future of, 497 500 UHMWPE behavior under action of external factors, 487 492 UHMWPE versus other biomaterials, 495 496 Polymeric neoglycoconjugates, 314, 325 Polymerization, parameters influencing, 235 extrinsic factors, 235 intrinsic factors, 235 Polymerization inhibitors, 221 Polymerization initiators, activators and, 221 Polymerization mode, 225 227 Polymerization shrinkage, 66 67, 235 236 Polymer/metal nanocomposites, 128 Polymer poly(D,L-lactide) (PDLLA), 352 353 Polymer water interaction, 246

Polymethacrylates, 217 biocompatibility, 243 244 challenges in improving properties, 227 creep and fatigue, 260 261 dental composites, 222 227 particle size and distribution of fillers, 223 225 polymerization mode, 225 227 viscosity, 225 humid ageing, 246 257 consequences of physical ageing on mechanical properties, 254 256 penetrant composition mixture, effect of, 257 role of the interface, 256 257 water diffusion, 250 254 water solubility, 246 250 hydrolysis, chemical ageing by, 257 259 medical applications, material selection for, 217 218 methacrylate networks, polymerization of, 233 234 mechanistic aspects, 233 polymerization kinetics, 234 monomers, 218 221 dental applications, methacrylates for, 219 220 matrix, composition of, 220 221 methyl methacrylate, 218 physical relaxation, aging by, 244 246 PMMA, radical polymerization reaction of, 228 233 kinetic aspects, 230 233 mechanistic aspects, 228 230 polymerization, parameters influencing, 235 extrinsic factors, 235 intrinsic factors, 235 polymerization shrinkage and its consequences, 235 236 radiolysis, chemical ageing by, 259 structure properties relationships and link with clinical applications, 236 243 glass transition temperature and other transitions, 236 238 short deformation properties, 238 242 ultimate properties, 242 243 Polymethylmethacrylate (PMAA), 218, 230, 242 243, 250, 262, 481, 484, 508 509, 513, 524 527, 531 532, 534, 541 542 chain propagation of, 230f chain termination of, 231f optical properties of, 510t physical properties of, 509t properties and advantages of, 509 Polyolefin-based biocomposites

Index

biocompatibility evaluation of, 177 179 flowchart of processing techniques for, 181f Polyolefin biomedical composites, fabrication techniques for, 180 185 extrusion, 182 183, 183f filament winding, 184 melt electrospinning, 183 184, 184f molding, 181 182, 182f thermoplastic pultrusion, 185, 185f Polyolefin homopolymers, 176 Polyolefins, 175 177 Polyols, 59 60 Polyoxyethylene (POE), 419 Polyphenols, 487 Polypropylene (PP), 176, 541 542 laser surface texturing (LST) of, 42 44 properties of, 201t typical structure of, 200f Polypropylene-based nonwoven fabric membrane (PPNWF), 207 Polypropylene carbonates/poly(lactic acid) (PLA) composite nanofibers, 205 Polypropylene matrix, 199 208 antimicrobial applications, 205 207 bone cement, 201 204 finger joint implants, 200 scaffolds, 204 205 sutures, 208 Polysiloxane chain, structure of, 508f Polytetrafluorethylene (PTFE), 484 laser surface texturing (LST) of, 48 Polyurethanes (PU), 59 64, 90 91, 126 radiation curing in surface modification, 63 64 shape memory polymers, 64 synthetic routes to photocrosslinkable polyurethane polymers, 60 63 Polyurethane thermoset, 137 Polyvinyl alcohol (PVA), 13 15, 359, 467, 470 471, 484 hydrogels based on, 14 Polyvinyl carbonates, degradation of, 72f Polyvinyl chloride (PVC), 275 Polyvinylidene difluoride (PVDF), 325, 335 Polyvinyl pyrrolidone (PVP), 69 70 Porcelain, 273 274 Porous coated motion (PCM), 481 Posterior capsular opacification (PCO), 515 517 PPF-PEG-g-BNNTs, 429 Pristine epoxy systems, 145 146, 154 156 N-Propyl benzoxazine derivatives, 114 Prostate-specific membrane antigen (PSMA), 469 Prosthetic heart valves thermoset metal polymer composites in, 135 136

Prosthetic sockets, polymeric materials in, 137 Prosthodontics, 297 Protein/peptide drug delivery system (DDS), 404 405 Pseudomonas putida, 207 Pseudophakic eyes, 519 PTVT (primary tube versus trabeculectomy), 537 Pulsed lasers, 38 Pultrusion, 185 4-Pyro(idinopyridine), 111 2-Pyrrolidone, 357 358

R Radial polymerization method, 371 372 Radiation curing technology, 57, 90 91 in surface modification, 63 64 Radiation sterilization, 8 Ratner’s correlation, 192 193 Rayon, 497 Reactive oxygen species (ROS), 439 440 Recombinant human growth hormone (rhGH), 404 405 Recombinant viral subunit-based vaccines, 401 Refractive index (RI), 508 Refractory glaucoma, 534 535, 539 ReGel, 465 Regel polymer, 461 462 Resin, defined, 111 112 Resin transfer molding, 153 154, 156 157 Resorbable nanoparticles, 466 472 commercial and investigational examples, 471 472 limitations of, 472 nanoparticle preparation and characterization techniques, 466 467 resorbable nanoparticles-based drug delivery systems, 468 471 Resorbable thermosensitive polymers, 459 466 commercial and investigational examples, 463 465 limitations of, 465 466 thermosensitive polymer-based drug delivery systems, 460 463 Reticuloendothelial system (RES), 450 Ring opening polymerization (ROP), 351, 381, 452 454 Risperidone, 369t Rockwell hardness tests, 151 Rubber, 497 Rubber microparticles, 129 130

S Salinomycin (SLM), 361 364 Sandblasting, 33

565

566

Index

Scanning electron microscopy (SEM), 151, 158 159 Self-assembled monolayers (SAMs), 34 Self-reinforced polymer composites, 197 198 Semicrystalline polymers, 9 S-glass fibers, 297 298 Shape memory polymers (SMPs), 15 17, 64 for biomedical applications, 163 166 Silane treatment, 153 154, 296 Silica nanoparticles, 129 130 Silicone, 507 508, 514 515, 522t, 538 539, 541 542 Silicone rubber composites, SEM micrographs of, 202f Silver (or gold)-labeled neoglycoconjugate, 334 341 lectin sensors, development of, 334 338 UV-visible absorbance spectroscopy, 338 341 Silver/gold glyconanoparticles, synthesis of, 325 Silver nanoparticles, 130, 135 Single edge notch (SEN), 285 Single polymer composites. See Self-reinforced polymer composites Single walled carbon nanotubes (SWCNT), 168 Sisal fiber, 155t, 157 158 Small cell lung cancer (SCLC), 398 Sorbents, 341 342 Spray drying, 358, 467 Staphylococcus aureus, 207, 529 531 Starch, 82, 486 Starch-based bioplastics, 486 Steam sterilization, 8 Step-growth polymerization, 110 Stereolithography, 89 90 Sterilization of plastics, 8 Stimulated body fluid (SBF) immersion time, 189 190 Streptococcus equi., 403 Succinic anhydride (SA), 453 454, 461 Supercritical fluid (SCF) technique, 358 359 Super paramagnetic iron oxide nanoparticles (SPIONs), 371 Surface chemical modification, 34 Surface modification of nanodiamonds, 168 Surface roughening, 32 34 Sutures, 16, 208 Swine influenza virus (SwIV) vaccine, 368 SwIV H1N2 antigens (KAg), 368

T Tack Quotient, 512 513 Tapered cleavage (TC) methods, 285 Tetrahydrofuran, 357 358

Tetramethylsilane (TMS), 454 455 Textured surfaces, 30 Th1 lymphocytes, 401 Thermal melting, 6 Thermodynamic theory, 254 Thermogelling polymers, 14 15 Thermogravimetric analysis (TGA), 120 121, 434 Thermogravimetric and differential thermal analysis (TG-DTA), 456 457 Thermoplastic and thermosetting polymers, 5 7 Thermoplastic polymers, 6, 8, 29, 48, 105 Thermoplastic pultrusion, 185, 185f Thermoplastics, 6 Thermoresponsive biomaterials, 14 15 Thermoresponsive polymers, 1, 459 Thermosensitive polymer-based drug delivery systems, 460 463 Thermoset metal polymer composites, 127 134 applications of, 133 134 in bone grafting, 136 137 in bones, 136 characterization of, 132 133 in dentistry, 134 135 in medical devices, 137 properties of, 130 132 dynamic mechanical properties, 132 fracture surface, 131 stress strain behavior, 131 132 tensile strength, 131 wear performance, 132 in prosthetic heart valves, 135 136 in prosthetic sockets, 137 synthesis, 128 130 Thermoset polymers, 105 127 applications of, 124 127 epoxy resins, 127 melamine formaldehyde(MF) resins, 125 126 phenol formaldehyde resin, 126 polyelectrolytes, 126 polyurethane, 126 unsaturated polyester resin (UPEs), 127 urea formaldehyde resin, 125 characterization of, 116 124 differential scanning colorimetry (DSC), 119 120 dynamic mechanical thermal analysis (DMTA), 122 124 Fourier transform infrared spectroscopy (FTIR), 117 118 nuclear magnetic resonance (NMR) spectroscopy, 118 119 thermogravimetric analysis (TGA), 120 121 X-ray fluorescence (XRF) spectroscopy, 124

Index

identification of, in chemistry, 106f properties of, 114 116 fiber impregnation, 116 formulations, 114 115 mechanical properties, 116 melt viscosity, 116 processing cycle, 116 solvent resistant, 115 116 synthesis of, 105 109 by crosslinking or curing, 111 114 by polymerization, 110 111 Thermoset polymethacrylate-based materials for dental applications, 273 aluminum, 274 275 bakelite, 275 base metal alloys, 276 celluloid, 275 chemical modification of PMMA, 293 gold, 273 poly(methyl methacrylate) (PMMA), contemporary denture base materials and modifications of, 291 292 epoxy resins, 292 polyamides, 291 292 polycarbonates, 292 poly(methyl methacrylate) (PMMA), properties of, 282 290 biocompatibility and cytotoxicity, 290 color stability, 289 crosslinking, 286 flexural strength (FS), 282 285 fracture toughness (FT), 285 286 impact strength, 286 radiopacity, 289 290 residual monomer, 288 sorption and solubility, 286 288 thermal conductivity, 288 poly(methyl methacrylate) (PMMA) resins, classification of, 276 282 according to method of polymerization, 277 282 according to the ISO standards, 276 polyvinyl chloride (PVC), 275 porcelain, 273 274 reinforcement of PMMA denture base materials, 293 300 aramid fibers, 297 carbon fibers (CFs), 296 fiber reinforcement, 294 296 glass fibers, 297 300 polyethylene (UHMWPE) fibers, 297 reinforcement with metal wires or mesh, 293 294 vulcanite, 274

Thermoset polyolefins, 135 136 Thermosets, 6 7, 105 106 Thiol-alkyne, 88 89 Thiol-ene materials, 88 89 Thiol-norbornene, 88 89 Third-generation tissue engineering techniques, 533 534 Tissue engineered-corneal analogs, 523 524 Tissue engineering, 167 168, 391 393, 417 418 polyethylene glycol (PEG) in, 419 Toluene diisocyanate, 59 60 Tool Narayanaswamy Moynihan model, 245 Total hip arthoplasties (THAs), 494 Total hip replacement (THR), 148 149, 153 Total joint arthroplasty (TJA), 494 Total joint replacement (TJR), 494 Total knee replacements (TKR), 148 149, 153, 495 Transition metal-based catalysts, 175 176 Transmission electron microscopy (TEM), 133, 321, 492 Transverse strength. See Flexural strength (FS) Triethylbenzylammonium (TEBAC), 117 118, 120 Triethylene glycol dimethacrylate (TEGDMA), 66 67, 135, 219 220, 258f Triethylenetetramine (TETA), 113 Trifluoroacetic acid, 323 324 Triglycerides, 487 1,3,5-Triglycidyl isocyanurate (TGIC), 149 150 3-(Trimethoxysilyl)propyl methacrylate, 221, 222f, 296 4,4 N Trimethylaniline, 222f 2,4,6-Trimethyl benzoyl-diphenyl phosphine oxide, 75 78 Trimethylolpropane (TMP), 453 454 Trimethylolpropane trimethacrylate, 223 Triphenyl 2,4,6-trimethylbenzoyldiphenylphosphine oxide (TPO) photoinitiator, 75 78 Triphenylbismuth (TPH), 289 290 Triple negative breast cancer (TNBC), 398 399 Triptorelin, 369t Tris-buffered saline (TBS), 325 Trithiophosphonate phloroglucinol (P3SP), 120 121 Tube versus trabeculectomy (TVT), 537

U Ultra-high molecular weight polyethylene (UHMWPE), 147, 158, 176, 187, 193 194, 196 197, 297 behavior under action of external factors, 487 492

567

568

Index

Ultra-high molecular weight polyethylene (UHMWPE) (Continued) laser surface texturing (LST) of, 46 47 UHMWPE-based biomedical composites, 193 198 versus other biomaterials, 495 496 Ultralow density polyethylene (ULDPE), 176 Ultraviolet (UV) lasers, 35 36 Uncorrected visual acuity (UCVA), 543, 545 Universal testing machine (UTM), 284 Unsaturated polyester resin (UPEs), 127 Upper critical transition temperature (UCST), 456 Urea formaldehyde resin, 125 Urethane acrylates, 60 63, 91 Urethane di(meth)acrylate (UDMA), 135 Urethane dimethacrylate monomers, 221f Urethanes, 60 UV and blue rays elimination from intraocular lenses, 520 523 UV-curing technology, 58, 88 89 UV radiation, 59 60, 69 70, 87 88, 521 UV-visible absorbance measurements, 325 UV-visible absorbance spectroscopy, 338 341

V Vaccination, 368 370 Vacuum-assisted resin transfer molding (VARTM), 153 154 Van der Waals bonds, 235 236

Van’t Hoff law, 246 248 Vinyl carbonate, 70 71 Vinyl esters, 72 73 Vinyl monomers, 69, 91 low photoreactivity of, 72f N-Vinyl pyrrolidone, 57 58, 69 70, 91 Viscosity, 225 Viscosity control agents, 79 Visible-light photoredox catalysis, 90 Vitamin E, 394 395, 482, 488 492 Vitrification, 109 110 Vivathane, 541 542 Vulcanite, 274

W Wang’s model, 192 193 Water contact angle (WCA), 31 32, 31f, 47 48 Water diffusion, 250 254 Water solubility, 246 250 Wenzel model, 32 Wide angle X-ray diffraction (WAXD), 457

X X-ray fluorescence (XRF) spectroscopy, 124

Y Yield stress, 241 Young’s modulus, 239, 242 243, 388