Biosynthetic Polymers for Medical Applications

Nondegradable synthetic polymers for medical devices and implants 2 P.A. Gunatillake, R. Adhikari Commonwealth Scientific and Industrial Research Organisation (CSIRO) Manufacturing Flagship, Clayton, VIC, Australia 2.1 Introduction The inert nature of many nondegradable synthetic polymers in biological environments and ease of processing have attracted researchers to explore their utility in biomedical applications. With the advancement of multitude of technologies over the second half of the twentieth century, many medical devices and implants have been developed from synthetic polymers and successfully used to help millions of patients worldwide. Today the biomaterial and medical device industry has developed to an estimated US$150 billion worldwide industry saving lives of millions and improving the quality of life for many millions more (Liu et al., 2012). Nondegradable synthetic polymers have played an important role in these products, often as inert materials, by contributing to the efficient functioning of the devices as well as providing mechanical support in many orthopaedic implants. For example, components in orthopaedic implants such as articulating surfaces, and scaffolds, or as protective coatings on electrically stimulating devices such as cardiac pacemakers. Among the synthetic polymers, the early materials explored for medical applications include silicone rubber, high density polyethylene, and polyesters such as Dacron. Since then, a wide spectrum of devices and implants comprised of inert synthetic polymer components have been developed and successfully used in numerous clinical applications. These applications include coatings on implants to improve blood compatibility, cardiovascular devices such as pacemakers, heart valves, and orthopaedic fixation devices such as knee and hip implants. Other applications include catheters and dialysis tubing, vascular grafts, implantable drug delivery systems such as drug eluting coatings on vascular stents. Since the early introduction of synthetic materials, significant advances have been made in polymer formulations and processing techniques to help optimise the performance and stability of these materials in the biological environment. With the realisation that inert materials may not always be the best in some of these applications, researchers have focused on improving the biological interaction of these materials through various surface modifications techniques and incorporating additives to help improve tissue integration with these materials. With the emergence of tissue engineering and regenerative medicine as the next frontier technology to repair damaged tissues or organs to restore normal biological Biosynthetic Polymers for Medical Applications. http://dx.doi.org/10.1016/B978-1-78242-105-4.00002-X Copyright © 2016 Elsevier Ltd. All rights reserved. 34 Biosynthetic Polymers for Medical Applications functions, many research groups worldwide have focused their attention to develop suitable biodegradable polymers to provide the material needs to further advance these technologies for clinical applications. Over the last two decades many reviews and research publications have appeared in the literature on new biodegradable materials and formulations. It is anticipated that these advancements will translate to clinically useful products and therapies in the near future. However, the use of nondegradable polymers in medical implants will continue to play a major role in many of the existing devices and next generation medical implants. The aim of this chapter is to provide the readers with a review of the current nondegradable polymers used in biomedical applications, particularly covering the advancements over the past decade and to assess their biological performance based on long-term use in humans. A brief introduction to the chemistry, physical properties, biocompatibility and biostability of seven major classes of synthetic polymers will be provided. The term ‘nondegradable’ is used to imply that the polymers are resistant to degradation by hydrolytic and other mechanisms operating in the biological environments. This chapter will not cover biodegradable polymers and readers are referred to many excellent reviews (Martina and Hutmacher, 2007; Nair and Laurencin, 2007; Ulery et al., 2011; Place et al., 2009; Hazer et al., 2012). 2.1.1 Introduction to various classes of synthetic polymers Synthetic polymers with chemical linkages resistant to hydrolytic, oxidative, and other degradation mechanisms operating in biological environments have been evaluated extensively for biomedical applications; however, only a few have made it to devices or implants in clinical use. A major attraction to use synthetic nondegradable polymers in biomedical applications stems from the ability to tailor mechanical properties and biological inertness, as well as a variety of available processing options. The main classes of nondegradable synthetic polymers used in biomedical applications include poly(olefins), poly(urethanes), poly(carbonates), poly(siloxanes), poly(amides), poly(ethers), poly(sulphones) and certain types of poly(esters). Table 2.1 lists representative chemical structures, general material properties and main biomedical applications of nondegradable synthetic polymers. The following sections outline the general chemistry, synthesis, properties and their suitability for biomedical applications based on numerous in vitro and in vivo evaluations. 2.2 Ultra-high molecular weight poly(ethylene) (UHMWPE) Poly(ethylene) (PE) is one of the most commonly used synthetic polymers for industrial and commercial products with millions of metric tons produced annually. Polymerisations using Ziegler Natta or metallocene-based catalysts are the most widely used method for production of many different grades of PE for commercial applications. PE is classified into different grades based on its density and branching; ultra-high Nondegradable synthetic polymers for medical devices and implants 35 molecular weight PE (UHMWPE), high density PE (HDPE) and ultra-low molecular weight PE (ULMWPE) are examples. UHMWPE is a very tough material and has outstanding toughness and resistance to cut, wear and chemicals as well as very low moisture permeability and very low coefficient of friction. For these reasons UHMWPE has been used for the construction of articulating portions of implants used in hip and knee replacements (Kurtz, 2004) and has a history of over four decades of use in bearing surfaces of total joint replacements (Sobieraj and Rimnac, 2009). Table 2.2 lists some of the key properties of UHMWPE and its mechanical properties are dependent on its molecular weight, degree of crosslinking and the relative amounts of crystalline and amorphous phases. 2.2.1 Biomedical applications of UHMWPE Since the first introduction of UHMWPE in hip replacements in 1962, the material is currently used in over 1.4 million patients around the world annually for implants in hip, knee, upper extremities and spine surgery (Kurtz et al., 1999b, Kurtz, 2004). This success is largely due to superior wear resistance of the polymer along with high fracture toughness and biocompatibility compared to other polymers. However, the long-term performance of these implants is less than optimal and in some cases revision surgery is required due to poor material performance. The failure of UHMWPE in joint replacement devices is primarily due to oxidative degradation, resulting in loss of mechanical properties (Kurtz et al., 1999a; Besong et al., 1998; del Prever et al., 2009). Over the last few decades many research groups have investigated the factors responsible for the material failure and have developed PE materials with improved performance. The major approaches investigated include minimisation of processing requirements during fabrication of the implants. These include sterilisation effects, crosslinking of the polymer using high energy radiation, blending with other materials and improvement of the surface properties. Due to the extremely high molecular weight of UHMWPE, the molten polymer produced from granules does not flow like its low molecular weight counterparts; as such the conventional thermoplastic processing equipment cannot be used for its processing. The material must be consolidated using controlled pressure, temperature and time to produce moulded or extruded parts. The consolidation process is diffusion controlled and UHMWPE requires sufficient time at elevated temperature and pressure for the molecular chains to migrate across grain boundaries (Kurtz, 2009). Specially designed and modified techniques of extrusion, compression moulding, direct compression moulding, ram extrusion and hot isotactic pressing allow the polymer to be extruded for the production of medical grade products. 2.2.2 Biocompatibility and stability of UHMWPE It is well known that UHMWPE is susceptible to oxidative degradation following radiation-induced sterilisation, leading to loss of mechanical properties affecting the long-term performance of implants with components fabricated from it. Many studies have investigated the effect of gamma irradiation on mechanical properties and wear 36 The chemical structure, general properties and major biomedical applications of nondegradable synthetic polymers Table 2.1 Polymer Chemical structure Major biomedical applications Poly(ethylene) Linear thermoplastic Low and high-density grades Excellent chemical resistance Total hip, knee and spine implants Poly(propylene) Linear thermoplastic Tough, flexible with good fatigue resistance Nonabsorbable suture, medical pouches and hernia mesh Poly(methyl methacrylate) Lightweight transparent thermoplastic, good impact strength, poor resistance to chemicals Orthopaedic prosthesis, dental, encapsulation device for cells, slow release of peptide and protein drug Biosynthetic Polymers for Medical Applications Key properties Transparent, elastic solid, nontoxic, inert, nonflammable, excellent viscoelastic properties Breast implant, contact lens, prosthesis Poly(ether ether ketone) Semicrystalline thermoplastic with excellent mechanical properties and chemical resistance Spinal fusion, disc arthroplasty, pedicle-based rod systems for nonfusion, interspinous spacers, minimally invasive fusion surgery, motion preservation and dynamic stabilisation Polyurethane Linear thermoplastic or crosslinked thermoset can be formulated as soft elastomers or rigid materials to have either biostable or biodegradable properties Pacemaker lead insulators, vascular grafts, catheters Nondegradable synthetic polymers for medical devices and implants Poly(dimethyl siloxane) 37 38 Table 2.2 Properties of nondegradable synthetic polymers Melting temp 8C Ultimate tensile strength (MPa) Ultimate tensile elongation (%) Tg 8C References Ultra-high molecular weight poly(ethylene) (UHMWPE) Extruded GUR 1020 135 54 452 þ 19 110 Kurtz (2009) Poly(propylene) (PP) 120e176 26e32 10e140 8 Handbook of Polymers (2012) Poly(methyl methacrylate) (PMMA) 160 48e76 2e10 105 Livermore and Voldman (2013) Poly(urethane) (PU) Elast-Eon™ 180e185 23 >500 45 Gunatillake et al. (2003) Poly(siloxane) NA 4e12 >500 127 Mata et al. (2005) Poly(ether ether ketone) PEEK 343 90e100 1.5e40 143 Kurtz and Devine (2007) Biosynthetic Polymers for Medical Applications Polymer Nondegradable synthetic polymers for medical devices and implants 39 performance of UHMWPE-based components used in hip, knee and other implants (Adrian et al., 2008; Berry et al., 2012; Affatato et al., 2002; Bracco et al., 2006; Carpentieri et al., 2011; Coote et al., 2000; Costa et al., 1998; Kaneeda et al., 1999; Sugano et al., 2004; Xiong et al., 2007; Xiong and Xiong, 2012). The free radicals generated during radiation sterilisation are responsible for initiating the chemical events leading to polymer degradation and subsequent deterioration of mechanical properties. The free radicals can react with the oxygen present forming hydroperoxides as the first product, which upon decomposition regenerate free radicals (Bracco et al., 2006; Costa et al., 2002, 2008). This autocatalytic process causes polymer main chain degradation leading to fragments with ketone, carboxyl and alcohol functional groups. The degradation process can continue as long as oxygen is available and it can continue without further irradiation (Bracco et al., 2006). This postirradiation ageing process leads to a reduction in the following properties: polymer elongation, modulus, ultimate tensile strength, fracture toughness, crack propagation and overall wear performance. Many studies have focused on the effect of the radiation dose and conditions on the postirradiation ageing process (Adrian et al., 2008; Bracco et al., 2006; Coote et al., 2000; Kaneeda et al., 1999; Carpentieri et al., 2011). Irradiation in an inert atmosphere or in vacuum packaging to eliminate oxygen can reduce or retard the ageing process (Kurtz et al., 1999a) but cannot be completely eliminated. While these approaches help reduce the onset of degradation during storage, the oxidation can still occur once implanted due to available oxygen in vivo. The free radicals produced during sterilisation are considered to have different resident times depending on whether they are residing in the amorphous or the crystalline phase (Carpentieri et al., 2011), with those in the amorphous phase decaying fastest. A similar in vivo oxidation response has been shown for implants sterilised under inert atmospheric conditions (Berry et al., 2012; Sugano et al., 2004; Kaneeda et al., 1999; Kumakura et al., 2009). The design of the components in implants has an influence on the extent and the location of degradation leading to gross fracture and cracking of the UHMWPE components. For example, cracks and gross fracture have occurred in acetabular rims (Berry et al., 2012; Sugano et al., 2004), the stabilising posts in noncruciate sparring tibial components (Hendel et al., 2003; Mariconda et al., 2000), and along rims of total spinal disc replacements (Kurtz et al., 2006a,b). The regions of these components where built-in stress is concentrated due to fabrication technique or the component design are more susceptible to fracture resulting from oxidation. Crosslinking of UHMWPE using ionising radiation (gamma rays or electron beams) has been one of the main approaches employed to improve the wear properties. Typically radiation doses ranging from 50 to 100 kGy (Kurtz, 2004) are used. The irradiation is followed by thermal treatment to destroy any residual free radicals to minimise oxidative degradation. This annealing process is typically conducted by heating the material first below Tm followed by melting above Tm (Kurtz et al., 1999b). However, the residual free radical within crystalline domains may not be completely destroyed by this annealing process (McKellop et al., 2000). Both of these thermal treatments can influence the degree of crystallinity in the polymer. Annealing typically increases crystallinity (Medel et al., 2005) whereas remelting causes an irreversible 40 Biosynthetic Polymers for Medical Applications Figure 2.1 Vivacit-E® highly crosslinked polyethylene liner (containing vitamin E) for use with the Continuum Acetabular System. Courtesy of Zimmer. decrease in crystallinity (McKellop et al., 2000). The crosslinking process can produce the desired improvement in wear resistance but with a compromise on other mechanical properties. A reduction in both ultimate stress and strain has been observed for highly crosslinked UHMWPE (Lewis, 2001; Pruitt, 2005; Sobieraj and Rimnac, 2009). The reduction in these properties is relatively larger for gamma-irradiated materials compared to those irradiated with e-beam. The incorporation of radical scavengers has been another approach to mitigate the effect of residual-free radicals in irradiated UHMWPE. Vitamin E (a-tocopherol) is arguably the most widely investigated radical scavenger (Affatato et al., 2012; Bracco and Oral, 2011; Mehmood et al., 2012; Wolf et al., 2006). Two methods have been employed to incorporate vitamin E: the first before irradiation and the second to allow it to diffuse into the polymer after irradiation (Shibata and Tomita, 2005). These improvements to UHMWPE have shown promise in improving its performance in orthopaedic implants (see Figure 2.1) (Zimmer, 2012); however, there is little information available to-date for their clinical performance. 2.3 Polypropylene (PP) Polypropylene (PP) is a thermoplastic polymer with a wide variety of industrial applications. The major industrial applications include packaging and labelling, carpets, reusable containers, laboratory equipment, automotive components, and bank notes. Nondegradable synthetic polymers for medical devices and implants 41 Due to the unsymmetrical nature of the monomer propylene, the polymer produced can has three different chain conformations: isotactic, syndiotactic and atactic, depending on the special orientation of the methyl group on each monomer unit within the polymer chain. The most widely used commercial PP is isotactic and has intermediate crystallinity with a melting point in the range 160e166  C, whereas the syndiotactic form has a melting point of 130  C with a crystallinity of 30%. PP is hydrophobic and has good resistance to many chemical solvents and acids. PP is commercially produced using ZieglereNatta and certain types of soluble metallocene catalysts. PP can be easily fabricated to various components using a variety of processing techniques, including compression, injection and blow moulding, and extrusion. The properties of PP are listed in Table 2.2. 2.3.1 Biomedical applications of PP In addition to a wide variety of industrial applications, PP has found applications in medically related products. PP is the material of choice for medical supplies such as syringes, catheters, vials, blood transfusion bags, dialysers for blood purification, and for tubing and membranes used in diagnosis and treatment instruments. PP is also used in soft tissue applications and two of the most common applications are nonabsorbable sutures (e.g. Prolene™, manufactured by Ethicon Inc.) and PP meshes for hernia and pelvic organ repair. The understanding of the long-term biological performance of PP-based implants is largely from its use in these applications. PP is liable to chain degradation, particularly when exposed to heat and UV radiation and the degradation is triggered by the free radicals formed on the tertiary carbon atom found in every repeat unit of the polymer; further reaction of the radical with oxygen can produce aldehydes and carboxylic acids leading to chain scission. It is estimated that over 20 million hernia operations are performed globally every year (Sanders and Kingsnorth, 2012). Since its introduction in hernia surgery in 1962, PP has become the most widely used polymer for inguinal hernia repair (67.6%) and incisional hernia repair (44.4%). This is largely due to a few favourable properties of PP including high burst strength and good mechanical properties. In hernia surgery, PP is used as a woven mesh to provide biomechanical strength to prevent the hernia from recurring by reinforcing the abdominal wall. The biological response to the foreign material is critical in the events after surgery leading to the completion of the wound healing process which can take up to 12 weeks (Majercik et al., 2006). The absorption of proteins on the surface of the implanted material occurs almost instantaneously upon implantation and is believed to be the first step in a series of biological events as part of the foreign-body reaction to the material. The material surface properties as well as the extent and the nature of these events influence the magnitude of the inflammatory response (Meintjes et al., 2011). While PP provides good mechanical support, one of the main disadvantages is the intensity of the foreign-body reaction, leading to less compliance and increased pain (Sanders and Kingsnorth, 2012). The design of the mesh (Bellon, 2009; Klosterhalfen et al., 2005) and fabrication methods can significantly improve its long-term performance of hernia repair 42 Biosynthetic Polymers for Medical Applications (Nayak et al., 2012). Many research groups have investigated surface modification approaches to improve the overall biocompatibility of PP for implant applications. Low pressure nitrogen plasma surface modification (Gomathi et al., 2012), plasmainduced graft polymerisation of acrylic acid (Gupta et al., 2008), graft polymerisation with other hydrophilic monomers such as N-isopropyl acrylamide (NiPAm) (Desai et al., 2003), conjugation of gold nanoparticles (Grant et al., 2011), surface modification with glycopolymer (Yang et al., 2005), and zwitterionic polymers (Zhao et al., 2011, 2012) are among the main approaches reported to improve PP surface properties in implant applications. The plasma surface modification is one of the main approaches used to surface modify PP as this can be carried out with fibres without modifying its mechanical performance. The approaches range from plasma modification with inert gases such as nitrogen (Gomathi et al., 2012) to plasma-induced polymerisation in the presence of hydrophilic monomers such as acrylic acid (Gupta et al., 2008), NiPAm (ContrerasGarcia et al., 2011; Desai et al., 2003), and acrylonitrile (Gupta et al., 2008). Surface modification with nitrogen plasma introduces nitrogen and oxygen containing functional groups (NH2, NH and CO) onto PP surfaces improving cell compatibility and haemocompatibility (Gomathi et al., 2012). Plasma-induced graft polymerisation of acrylic acid followed by chitosan binding helps to improve antimicrobial properties of PP (Gupta et al., 2008; Saxena et al., 2011). A number of other hydrophilic monomers such as vinylimidazole, NiPAm (Contreras-Garcia et al., 2011), 2-acrylamido2-methylpropane sulphonic acid (Song et al., 2011), D-gluconamidoethyl methacrylate (Yang et al., 2005), and zwitterionic monomers such as 2-methacryloyloxyethyl phosphocholine (Zhao et al., 2011, 2012) have also been reported. Overall, these surface chemical modifications have improved the wettability of PP and cell compatibility and growth. However, long-term in vivo data are needed to fully assess the impact of these modifications to improve the functional performance of membranes used in patients. 2.4 Poly(methyl methacrylate) (PMMA) Poly(methyl methacrylate) is a transparent thermoplastic polymer produced by polymerisation of methyl methacrylate (Table 2.1). The polymer was developed in 1928 in various laboratories and was first brought to the market by Rohm and Hass Company under the trade name Plexiglass. Since then it has been sold under several different names such as Lucite and Perspex. PMMA is currently produced by emulsion, solution and bulk polymerisation methods. Generally, the polymerisation is carried out by free radical polymerisation methods including controlled radical polymerisation methods as well as anionic polymerisation. All commercial PMMA is produced by radical polymerisation and is atactic and completely amorphous. The glass transition temperature of atactic PMMA is 105  C but most commercial PMMA has Tg in a wide temperature range (85e165  C) due to various commercial compositions incorporating comonomers to modify properties (Table 2.2). All Nondegradable synthetic polymers for medical devices and implants 43 common processing techniques such as injection moulding, compression moulding and extrusion used for thermoplastic polymers can be used to process PMMA. PMMA is a strong lightweight material with good impact strength and is transparent, transmitting up to 92% of visible light. PMMA swells and dissolves in many organic solvents and has poor resistance to many other chemicals. The PMMA homopolymer is versatile with many of its uses in a wide range of fields and applications but for most commercial applications the polymer is modified mostly by copolymerising with varying amounts of other monomers. It is widely used as a transparent glass substitute, for artistic and aesthetic uses, and for daylight redirection panels. PMMA has found many applications in medical technologies and implants as well. The major application areas include orthopaedic, cranial and facial reconstruction, ocular lenses, drug delivery and dentistry. 2.4.1 Biomedical applications of PMMA Since the first use of acrylic bone cements in the 1960s (Charnley, 1960), PMMA bone cement has been used in many load-bearing orthopaedic fixation procedures, including anchoring total joint replacement, vertebral body augmentation procedures and in balloon Kyphoplasty (Kenny and Buggy, 2003). The commercial PMMA bone cement is supplied as a two-part system for mixing and delivery at the point of use. One component is a liquid consisting primarily methyl methacrylate monomer and a small amount of an accelerator (N,N-dimethyl-p-toluidine) with traces of hydroquinone whereas the other component is PMMA powder with components like radio-opaque compounds (e.g. barium sulphate or zirconium oxide) as well as the free radical initiator (benzoyl peroxide). The use of prepolymerised PMMA powder in the formulation reduces the volume shrinkage during polymerisation while helping to manage the temperature rise during polymerisation. The formulation details and properties of many commercial PMMA bone cements can be found in review articles (Jaeblon, 2010; Kenny and Buggy, 2003; DiMaio, 2002; Lewis, 2011) and in books (Kuhn, 2000). The two components when mixed form a clear and uniform mixture and the polymerisation of the monomer is initiated generating heat from the reaction exotherm. The peak temperature can reach 113  C and the consistency of the reaction mixture changes from a low viscous sticky phase to a workable putty like consistency (dough) and finally to a hardening phase leading to a high strength solid mass. The mixture has to be delivered during the working phase to the implant site and for most commercial formulations the hardening occurs within 10e20 min. In most implant applications PMMA cement serves as space filler or as an interphase between the implant and the host bone providing stability to the implant. The PMMA layer helps transfer the mechanical forces from the implant to the bone due to comparable mechanical properties of PMMA to bone. The long-term performance of the implants is dependent on the quality of the apposition of the implant-cement and cement-bone. The surgical technique, loading characteristics, as well as the properties of cement, bone and implant can have a significant influence on the long-term performance of the cement (Jaeblon, 2010). 44 Biosynthetic Polymers for Medical Applications PMMA bone cement is also used in the treatment of vertebral defects (Laredo and Hamze, 2004) using a procedure termed vertebroplasty. This procedure consists of percutaneous injection of PMMA into vertebral collapse in order to obtain pain relief and mechanical strengthening of the vertebral body. Kyphoplasty is a variant of vertebroplasty used to help restore the height of vertebral compression fractures (Lieberman et al., 2002). Although the procedure is used extensively, it is not free of complications and the leakage of PMMA during the procedure is the main source of complications. 2.4.2 Biocompatibility and stability of PMMA Despite the success of PMMA bone cement in orthopaedic procedures, several problems have been identified associated with its use. Some of these problems are procedure related and the others are attributed to material properties and the polymerisation reaction employed to set the cement. For example, in the vertebroplasty procedure the leakage of the PMMA cement to spinal canal is very frequent, leading to associated complications. The technical refinements to reduce the risk of PMMA leakage are reviewed by Laredo (Laredo and Hamze, 2004) and readers are referred to this review article for further information. Numerous in vitro studies have reported on the cytotoxicity issues related to the use of PMMA resins in a range of medical implant applications. A review summarised the biocompatibility of PMMA resins used in dentistry (Gautam et al., 2012). PMMA resins used in dentistry are considered cytotoxic on account of leaching of various potential toxic substances such as unreacted monomer and other additives present in various formulations which may include initiator residues, activators, and monomer degradation products and crosslinking agents. The amount of residual monomer can vary depending on the type of polymerisation with the autopolymerised resins exhibiting a higher level compared to heat-polymerised resin activated with benzoyl peroxide (Vallittu et al., 1998). Irritation of the oral mucosa is the main cytotoxic effect of MMA and the systemic toxicity is considered very low consistent with high oral LD50 of MMA in rats (9 g/kg) (Gautam et al., 2012). Although many in vitro studies have investigated the cytotoxicity issues of PMMA leachables, only few in vivo studies have been reported. The heat generated during the polymerisation in PMMA bone cements has also been a concern as this could lead to cell necrosis in tissues closer to cement. It is reported that although the peak temperature during cure can reach about 113  C, in vivo, it is much lower between 40  C and 56  C (Kuehn et al., 2005), largely due to rapid heat dissipation through blood flow around the area as well as heat conduction throughout the implant material. 2.5 Polyurethane (PU) PUs represent an important class of synthetic polymers used in the manufacture of rigid and flexible foams, adhesives, surface coatings, sealants, and fibres covering a Nondegradable synthetic polymers for medical devices and implants 45 broad range of industrial applications (Ortel, 1994). Polyurethanes were first evaluated for biomedical applications in 1960 and since then numerous studies have explored their applications for a range of medical implants. The current major application is for insulating cardiac pacemakers. Polyurethanes have excellent mechanical properties and are easily formulated to have broad range of mechanical properties for applications as soft elastomers or high strength rigid materials. Polyurethanes are well tolerated in the body and exhibit good compatibility with cells and soft tissues (Lamba et al., 1998). 2.5.1 Synthesis and properties of PU The chemical reaction between an isocyanate group and a hydroxyl or amine group generates urethanes (carbamates) and urea groups, respectively. This reaction has been employed to synthesise a range of thermoplastic polyurethanes (TPU) and thermoset polyurethanes (TS). TPUs are prepared by reacting a diisocyanate, a difunctional polyol (macrodiol) and a dihydroxy or diamine chain extender, whereas TSs are synthesised using the same reagents, except at least one of those reagents has to be tri or higher functional to produce branched or crosslinked polymers. A vast number of potential combinations of the three reagents have been explored to synthesise polyurethanes with a broad range of mechanical properties (Ortel, 1994). However, only a relatively few have been used in medical applications, primarily due to toxicity, stability and mechanical properties. Due to ease of handling, symmetrical structure and high reactivity, 4,40 -methylenediphenyl diisocyante is the most frequently used diisocyanate in formulating biostable polyurethanes for biomedical applications. The chain extender is usually a low molecular weight diol such as 1,4-butanediol, although low molecular weight diamines, such as ethylene diamine may be used. The macrodiol is a diol of higher molecular weight (600e2000 Da) and poly(tetramethylene oxide) (PTMO) is one of the common macrodiols used in polyurethanes such as Pellethane™ used in medical implants such as cardiac pacemakers (Szycher et al., 1996). Polyurethanes can be prepared by one- or two-step batch procedures or by semicontinuous processes such as reactive extrusion (Ortel, 1994). One-step batch synthesis for thermoplastic polyurethanes involves reacting a mixture of the predried macrodiol and the chain extender with the diisocyanate. The reaction is generally catalysed with dibutyltin dilaurate, stannous octoate or amines and is exothermic. The ‘one-step’ process can also be carried out in special continuous mixing machines, or reactive extruders, or in continuous injection moulding machines. The reactive extrusion offers significant advantages in polyurethane synthesis for biomedical applications due to better control of reaction conditions as well as the ability to produce articles or components in one step, avoiding post synthesis processing steps. Compared with the one-step procedure, the two-step procedure gives good control of polymer architecture and can be carried out in bulk or in solvents such as rigorously dried N,N-dimethylformamide or N,N-dimethylacetamide. This procedure is mostly used for laboratory-scale preparations and is less attractive in industry-scale synthesis. The two-step batch procedure involves the synthesis of a prepolymer in the first step by 46 Biosynthetic Polymers for Medical Applications end-capping the polyol with diisocyanate, and chain extending in the second step by reacting with the chain extender. 2.5.2 Biostability and applications of PU The attractive mechanical properties such as high strength, elongation, tear and abrasion resistance of polyurethanes are attributed to the micro-phase separated morphology brought about by the partial mutual insolubility of the different chemical segments in the block copolymers. One microphase is derived from the macrodiol and is generally referred to as ‘soft’ segment which imparts material softness and extensibility. The other segment referred to as the ‘hard’ segment which is derived from the diisocyanate and the chain extender imparts cohesive strength to the polymer matrix. The chemical structures of ‘soft’ and ‘hard’ segments as well as their relative proportions within a polymer chain affect the stability of the polyurethane in the biological environment. It was recognised early that soft segments based on polyester macrodiols are susceptible to hydrolytic degradation and are not suitable for stable biomedical applications (Lamba et al., 1998; Stokes and McVennes, 1995). Accordingly, the polyurethane investigated for biomedical applications were based on polyether macrodiols such as PTMO. For example, Pellethane-80A used in pacemaker insulations was based on PTMO with a molecular weight of 1000. However, it was later realised that PTMO-based polyurethanes are susceptible to oxidative degradation (Szycher et al., 1996). The mechanism of biologically induced degradation of polyurethanes has been the subject of much investigation. The results have been summarised in a number of review articles (Szycher et al., 1996; Santerre et al., 2005; Anderson et al., 2008; Ward et al., 2006b; Christenson et al., 2007). Although the exact mechanisms are not fully understood, it is widely held that oxidative pathways involving the polyether ‘soft’ segment leading to environmental stress cracking as one of the main mechanisms. This form of degradation is found to be concentrated in areas where residual polymer surface stresses are present due to device fabrication and not sufficiently reduced by annealing. Hydrolytic enzymes have also been implicated in polyurethane biodegradation (Santerre et al., 1995). The polyurethanes investigated earlier for biomedical applications were not specifically designed for such applications and as such the understanding of the degradation mechanisms has led to commercial withdrawal of polyurethanes in the 1990s for chronic implants. This provided a stimulus to design more stable polyurethanes for medical applications. The design approaches were focused on the chemical structure of ‘soft’ segment forming macrodiols which are not susceptible to oxidative and hydrolytic degradation. The macrodiols investigated to date can be classified into four main groups based on the type of backbone functional groups. These are hydrocarbon, ether, siloxane and carbonate. In this field, the leading research groups are, Coury et al. (Takahara et al., 1991), Pinchuk (1995), Ward et al. (2006a), Szycher et al. (1996) and Gunatillake et al. (2003). Hydrocarbon backbone macrodiols such as hydrogenated polybutadine diol and poly(isobutylene) results in PUs with significantly improved resistance to thermal, Nondegradable synthetic polymers for medical devices and implants 47 hydrolytic and UV degradation compared to polyether or polyester polyurethanes (Speckhard et al., 1983; Ojha et al., 2009). However, the mechanical properties are inferior to their polyether counterparts and their synthesis is more difficult, making them less attractive for biomedical applications. In another approach C-18 fatty acids were dimerised to form polyols with predominantly hydrocarbon backbone soft segments. The biostability of the resultant PU’s improved but these materials were too stiff for most implant applications. Polyether macrodiols prepared from 1,6-hexanediol, 1,8-octanediol and 1,10 decanediol have shown significantly improved in vivo biostability compared with PTMO-based polyurethanes (Gunatillake et al., 2003); however, the polyurethanes exhibited higher modulus and stiffness as the number of ether linkages was reduced. Polycarbonate-based macrodiols yielded polyurethane with good mechanical properties and improved resistance to degradation in vivo compared to PTMO-based polyurethanes; however, in the longer term carbonate linkages were susceptible to degradation (Christenson et al., 2004). The incorporation of PDMS into the polyurethane structure has the advantage of imparting good haemocompatibility, flexibility, excellent hydrolytic and oxidative stability to polyurethanes. One approach has been to incorporate PDMS segments as polyurethane chain end groups to improve the surface properties of PUs and to reduce oxidative degradation. Another approach was to incorporate PDMS segments to both ‘soft’ and ‘hard’ segments of polyurethane by using appropriate siloxane macrodiols and chain extenders. This second approach has been the most successful in developing polyurethanes with vastly improved biostability, over conventional and other polyurethanes developed in recent times (Figure 2.2). The incorporation of siloxane segments as part of polyurethane structure is not straightforward due to the incompatibility between highly nonpolar siloxane segments with polar urethane segments. This incompatibility results in a high degree of phase separation yielding highly phase separated polyurethanes. This problem has been overcome by incorporating a small amount of a second macrodiol to improve compatibility as well as incorporating short siloxane chain extender segments to the ‘hard’ segment. By appropriate choice of catalysts both one and twostep methods have been adopted to prepare siloxane polyurethanes exhibiting good mechanical properties and processability. This new family (Elast-Eon™) (Table 2.2) of siloxane polyurethanes has been tested for biostability using both in vitro (Choi et al., 2009; Simmons et al., 2006) and in vivo tests and shown to have superior biostability. The in vitro oxidative stability of these polyurethanes is closely related to the morphology resulting from unlike segment demixing (Choi et al., 2009). ElastEon™ polyurethanes are currently used in cardiac pacemakers marketed by St. Jude Medical, USA as Optim™ (Figure 2.3). The biostability of Elast-Eon™ polyurethanes has been evaluated under both in vitro and in vivo conditions by a number of research groups. In a study (Wheatley et al., 2000, 2001) to evaluate the suitability of Elast-Eon™ polyurethane for tri-leaflet heart valves, bileaflet mechanical, Carpentier-Edwards porcine (bioprosthetic) and polyurethane heart valves (Figure 2.4) were implanted in juvenile sheep for six months. The study concluded that in the absence of valve-related deaths, and retention of good haemodynamic function, the PU valve was superior to bioprosthetic. 48 Biosynthetic Polymers for Medical Applications CH3 CH3 O HO R Si Si R CH3 CH3 OCN OH m + HO (CH2)4 OH OH n PHMO CH3 CH3 BDO CH3 O O R Si Si CH3 O R O C O NH C O O (CH2)4 O C NH O NH C p q H C H (CH2)6 O C n Soft segment NH m O O NCO MDI α, ω PDMS H O (CH2)6 H C H NH O NH C O O (CH2)4 O C NH O NH C r s Hard segment Figure 2.2 Chemical structure of Elast-Eon™ polyurethane. Courtesy of Gunatillake P and Adhikari R with permission from AorTech Biomaterials Pty Ltd. Figure 2.3 Elast-Eon™ polyurethane lead used for cardiac pacing applications. Courtesy of Dr Ajay Padsalgikar with permission from AorTech Biomaterials Pty Ltd. Nondegradable synthetic polymers for medical devices and implants 49 Figure 2.4 Prototype tri-leaflet heart valve. Courtesy of Dr Robert Kerton with permission from CSIRO Publishing. 2.6 Poly(dimethyl siloxane) (PDMS) Polysiloxanes are characterised by the presence of -Si-O-Si- linkage in the backbone and the most common polysiloxane is poly(dimethyl siloxane) (PDMS). The polymer in its pure form has very poor mechanical properties; PDMS with very high molecular weight tends to clod flow. For most applications, PDMS has to be crosslinked or reinforced with fillers to improve mechanical properties and is generally referred to as silicone rubber. Silicone rubber is nonreactive, stable and resistant to extreme environments and retains useful properties for applications in temperatures between 55  C and þ300  C (Table 2.2). Silicone rubber has also favourable properties for biomedical applications because of its good biocompatibility. Due to these properties silicone rubber is used in a wide range of products, including automotive applications, cooking, baking and apparel such as undergarments, sportswear, footwear, electronics, medical devices and implants. The term ‘silicone’ is used to denote polymerised siloxanes or poly(siloxanes) with the chemical formula [R2SiO]n where R is an organic group such as methyl, ethyl or phenyl present as side groups on Si along the -Si-O-Si-O- backbone. Silicones with a wide variety of properties and compositions can be prepared by varying the length of the backbone, the nature of the side groups or by crosslinking. The substitution of methyl groups along the chain with other groups such as phenyl, vinyl, trifluromethyl leads to silicones with unique properties. The polymers can vary from liquid to gel or to rubber and hard plastic. 2.6.1 Chemistry and synthesis of PDMS PDMS is prepared from dimethyldichlorosilane which reacts with water to produce the linear polymer with either hydroxyl or chloride end groups and under different 50 Biosynthetic Polymers for Medical Applications Linear and cyclic siloxanes CI Si CH3 CH3 CH3 CI HO 2H2O Si O HC1 3,4,5 n CH3 CH3 Si O H CH3 D3, D4, D5 CH3 D3 acid/base catalyst ring opening CH3 HO Si Si CH3 CH3 m CH3 condensation O H O HO Si n CH3 O H (p x n) Functional siloxanes CH3 H 3C H 3C Si O Si CH3 CH3 CH3 CH3 Si CH3 CH3 CH3 O Si CH3 CH3 cyclics D3, D4 H3C Si CH3 O Si CH3 CH3 CH3 cyclics D3, D4 H H 3C H O Si CH3 n CH3 CH3 CH3 x CH3 O Si Si CH3 y CH3 catalyst CH3 H3C Si CH3 H Si O p CH3 CH3 O Si CH3 CH3 q Figure 2.5 Principal chemical reactions in the synthesis of silicone polymers and functionalised silicones. conditions, the polymer formed is cyclic (Figure 2.5). The cyclic monomers hexamethylcyclotrisiloxane (D3) and octamethylcyclotetrasiloxane (D4) can also be used to prepare PDMS using basic or acidic catalysts. Commonly used basic catalysts include potassium, sodium or tetramethylammonium silanolate. The reaction mixture typically consists of a monomer and an end-capper which is used to control molecular weight and stability. The mixture is heated under moderate temperature (60e150  C) conditions to affect the polymerisation. Hexamethyldisiloxane is a common end-capper producing PDMS with trimethylsiloxy end groups which help stabilise the polymer from changes in viscosity and other properties for most fluid Nondegradable synthetic polymers for medical devices and implants 51 applications. The methyl groups on Si are not reactive and prevent further condensation. However, functional groups such as OH (silanols) or H (silanes) are reactive, for example, Si-OH groups are susceptible to condensation under acid or mild base conditions and they are intermediates for room temperature vulcanisable silicone formulations (Noll, 1968). 2.6.2 Biomedical applications of PDMS The unique material properties of silicones and excellent biocompatibility make them well suited for a range of health care products covering numerous personalcare, pharmaceutical and medical device applications. The realisation in the mid-1940s that silicon coatings on glassware and needles prevents blood clotting for hours led to the use of silicon coatings on syringes, needles and blood collecting vials. Subsequently, silicone was used in fabricating implants for bile duct repair and artificial urethra demonstrating that there was no abnormal reaction by the body to silicone and the devices performed well. Perhaps, the most notable first application of silicone in the body is the hydrocephalus shunt in 1957 (Lafay, 1957). The interest of silicone for health care application continued in the 1960 with General Electric and Dow Corning supplying the materials for physicians and researchers, and by the end of the decade silicone was used or investigated in numerous applications, including orthopaedics, catheters, drains and shunts, blood oxygenators, heart-bypass machines, heart valves and aesthetic implants (Curtis and Colas, 2004b; McMillin, 2006). For over 40 years, silicones have been used extensively in aesthetic and reconstructive surgery and despite the numerous allegations in the 1990s with regard to the safety of silicone breast implants, the use of silicones continues. Following the first implantation of a pair of silicone gel-filled breast implants (see Figure 2.6) (Friedman, 2010) in 1962, the popularity of silicone gel for breast reconstruction increased. The litigation related to breast implant safety raised controversial issues such as risk of breast cancer, autoimmune connective tissue disease as well as local or surgical complications such as rupture, infection, or capsular contraction. Many subsequent studies have found that these allegations are without sound scientific evidence and epidemiology studies have found that there was no association between breast implants and breast cancer. Many reports, papers and monographs have been published on this subject and the readers are referred to these excellent reviews for detailed information (Curtis et al., 2000; Brandon et al., 2003; Friedman, 2010; Brook, 2006). 2.6.3 Biocompatibility and stability of PDMS Among the synthetic polymers, silicone can be considered as the material with the longest history in medical implant applications (De Nicola, 1950). Numerous in vitro and in vivo studies have provided evidence to claim that silicone materials are ‘biocompatible’ (Curtis and Colas, 2004a). However, such claims and the data supporting those claims should be carefully considered in view of the modern definition of biocompatibility, which was the subject of much discussion. The modern definition ‘the ability of a material to perform with an appropriate host in specific 52 Biosynthetic Polymers for Medical Applications Figure 2.6 Silicon breast implant gels. Courtesy of Richard Szabo with permission from Get Media, Leadewrs in Niche Digital Publishing. situation’ (Williams, 2003) warrants careful review of the type of biological fluids/ tissues that the material comes into contact with before making such claims. Silicone materials in various forms such as liquids, gels and elastomers have been used as implants in contact with biological tissues and fluids in different parts of the body. While the bulk polymer may not induce adverse reaction from the body, with synthetic materials it is often the impurities present in the material that causes biocompatibility issues. Silicones are no exception as catalysts are used in the polymerisation and subsequent crosslinking reactions to manufacture the silicone for various applications. However, in addition to the wealth of information from in vitro test results for various silicone formulations and applications, results of long-term in vivo studies as well as explant analysis results substantiate the claim that silicone biocompatibility is excellent. Among the many studies to evaluate the biodurability of silicone elastomers and implants, the following can be cited as major studies to confirm long-term durability of silicon. In a comprehensive study to evaluate 42 silicone breast explants including human implantation duration of up to 32 years, Branden et al. (Brandon et al., 2003) reported that there was little or no degradation of the base polydimethylsiloxane during in vivo ageing of the explants they examined. In another study (Curtis et al., 2000) silicone breast implants were surgically excised and examined after 13.8e19.3 years. Only minor changes in the tensile strength of the shell were observed and the gel extract molecular weight remains unchanged. Despite its excellent biocompatibility, the relatively poor mechanical properties limit applications of silicones in medical implants. In particular, poor tensile strength, abrasion and tear strength limit the application to less load-bearing capacity. Nondegradable synthetic polymers for medical devices and implants 2.7 53 Polyether ether ketone (PEEK) PEEK is a semicrystalline thermoplastic with excellent mechanical and chemical resistance properties which are retained at high temperatures. The polymer is very rigid with a Young’s modulus of 3.6 GPa and tensile strength in the range 90e100 MPa. The glass transition temperature is around 143  C (Table 2.2). The polymer is highly resistant to thermal degradation as well as degradation from organic and aqueous environments. The polymer can be dissolved in concentrated sulphuric acid. PEEK is the leading polymer candidate used in orthopaedic applications from a family of polymers generally referred to as polyaryl ether ketones (Kurtz and Devine, 2007). 2.7.1 Synthesis and properties of PEEK PEEK polymers are synthesised by the dialkylation of the bisphenolate salts using step-growth polymerisation method. In a typical polymerisation procedure, 4,40 difluorobenzophenone is reacted with disodium salt of hydroquinone and the reaction is carried out in polar aprotic solvents such as diphenylsulphone at approximately 300  C to complete the nucleophilic substitution reaction (Figure 2.7). PEEK can be processed using a range of commercial processing techniques, including injection moulding, extrusion and compression moulding at temperatures between 390 and 420  C, which are significantly higher compared to that used for conventional thermoplastic polymers. The physical properties of PEEK are shown in Table 2.2. The extremely high thermal and chemical resistance is attributed to its resonance stabilised chemical structure of PEEK. Its outstanding chemical resistance is also reflected in its resistance to postirradiation degradation, allowing PEEK to be sterilised by gamma and electron beam irradiation. Although free radicals may be formed during irradiation, unlike in polymers like UHMWPE the free radicals formed are believed to decay rapidly due to their mobility along the polymer chain. In a study by Li et al. (1999), no evidence of residual free radicals were observed in PEEK exposed to Figure 2.7 Chemical synthesis of PEEK. 54 Biosynthetic Polymers for Medical Applications gamma irradiation of up to 600 kGy, indicating that the life time of the free radical is less than 20 minutes. In another study (Kwarteng and Starck, 1990), repeated sterilisation doses up to four 25e40 kGy of gamma radiation in air confirmed that no significant changes in mechanical properties of PEEK resulted. As such, postirradiation degradation is not expected to be a clinically relevant concern with PEEK-based medical implants as it has been the case with UHMWPE. Unmodified PEEK has considerable ductility and can accommodate large deformation plastic flow under both uniaxial tension and compression. Detailed studies of the true stressestrain behaviour of PEEK are reported in the literature (Rae et al., 2007; Hamdan and Swallowe, 1996). At room temperature and under very low strain (0.03) PEEK displays a linear relationship between stress and strain in both tension and compression. As the strain is increased the material exhibits a clear yield transition in the stressestrain curve. For industrial engineering applications, the mechanical properties of PEEK generally decreases with elevated temperatures up to 250  C, with a significant decrease above 150  C, which is slightly above the glass transition temperature (Rae et al., 2007). However, for biomedical applications where the application temperature is body temperature, PEEK elastic properties are relatively insensitive to temperature. For implant applications that involve heat generation, particularly load-bearing applications such as joint replacement require more detailed characterisation of PEEK with respect to plastic flow, and fracture behaviour is required (Kurtz and Devine, 2007). PEEK is gaining acceptance as a suitable material for fabrication of components for certain orthopaedic procedures, for example, in cranioplasty (Figure 2.8). Figure 2.8 Custom skull implant fabricated from PEEK for use in cranioplasty. Courtesy of Xilloc, the Netherlands. Nondegradable synthetic polymers for medical devices and implants 2.7.2 55 Biocompatibility and implant applications of PEEK Numerous studies on the systemic and intracutaneous toxicity and intramuscular implantation have shown that PEEK and its composites do not illicit adverse tissue reactions (Williams et al., 1987; Petillo et al., 1994). The first (Williams et al., 1987; Toth, 2012) in vivo study in mice reported that PEEK elicited no adverse tissue reaction. A subsequent in vitro study using mouse fibroblasts by Wenz et al. (1990) showed that the cell cultures were healthy with no difference to negative controls. Numerous other in vitro studies using different cell types including human-derived osteoblasts, fibroblasts (Wenz et al., 1990; Hunter et al., 1995), murine macrophages (Scotchford et al., 2003) have confirmed the noncyctotoxicity of PEEK. The early investigations of PEEK polymers for orthopaedic implants occurred in the mid-tolate 1980s (Skinner, 1988; Brown et al., 1990) but it was only in the 1990s that PEEK received consideration in the field of spine implants. A book (Kurtz, 2012) titled ‘PEEK Biomaterials Handbook’ provides a comprehensive account of the development of PEEK as a biomaterial. 2.8 Future directions For many decades nondegradable synthetic polymers have found applications in a variety of medical implants. The acceptance of these polymers for a wide variety of implants is based on the consideration of the mechanical properties, relative inertness of the polymers in their respective biological environments, and their ability to retain a required mechanical strength under dynamic environments for a long period of time. Since the first introduction of the synthetic polymers in implants in clinical use and subsequent findings from clinical experience about the material deficiencies, many research groups have investigated various approaches to improve their performance. In particular, the focus since the mid-1990s was to explore methods to minimise processing-related factors, imperfections at a molecular level and material surface properties to enhance biological tissue compatibility. Similar research is expected to continue as most of these polymers continue to play a major role in many of the existing and emerging implants. The adaption of techniques such as additive manufacturing will play a major role in minimising process-related defects in fabrication of implants and components based on synthetic polymers. And the incorporation of additives to control infections, improve biocompatibility and to deliver therapeutic agents will continue to be areas for further research. The wear performance of crosslinked UHMWPEs has significantly improved over that of conventional UHMWPE based on short-term clinical studies. One of the main improvements is post processing following radiation-induced crosslinking to eliminate residual radicals which has helped the long-term stability of UHMWPE implants. The incorporation of radical scavengers such as vitamin E may also help to reduce polymer degradation. These developments are encouraging and can expect to extend the functional life of UHMWPE-based implants. However, long-term clinical data is still 56 Biosynthetic Polymers for Medical Applications required to understand and assess the performance enhancements associated with these improvements. Further research into understanding the molecular-level changes resulting from these modifications and effects on mechanical performance needs to be conducted. PMMA bone cement formulations have also been improved, enabling better curing characteristics (Copal®, Osteopal®) and bioactive loading capabilities (Palamed®, Refobacin®-Palacos®). Improvements are still needed as the toxicity associated with long-term exposure to unreacted methacrylate monomer will continue to be an issue of concern. Silicone materials remain as a class of synthetic polymers thoroughly tested for a range of important medical implants due to their biocompatibility and biodurability. The limitation of the use of silicone materials in a wide range of medical implants is the relatively poor mechanical properties. Some of the recent advances in incorporating silicone segments to other high strength materials such as polyurethanes have resulted in materials with improved mechanical properties, expanding the range of applications. Further efforts in this area should lead to development of materials that can fulfil the materials needs for next generation medical implants, particularly in neural and vascular environments. Based on recent developments, PEEK has been recognised as a next generation high strength biomaterial, particularly for spine implants. The attractiveness of PEEK stems from the fact that it is biocompatible, inert, radiolucent and inherently strong. The use of PEEK in more advanced fields such as total joint replacement and fracture fixation may take many more years as the introduction of new materials in the biomedical field is slow requiring extensive testing. However, as new implant designs progress PEEK may be considered as a material that can offer a number of advantages over the more conventional synthetic polymers used in orthopaedic implants. The on-going development in PEEK composites, particularly those incorporating additive such as hydroxyapatite may provide novel materials to improve the functional performance of orthopaedic implants. Silicone-based polyurethanes such as Elast-Eon™ have shown to be significantly more biostable than conventional polyether polyurethanes and have shown excellent performance in devices such as cardiac pacemakers. With the expansion of the range of implants based on silicone polyurethanes, better understanding of the long-term biostability will emerge. These results will provide the researchers valuable information to broaden the application range as well as to address any materials deficiencies in specific applications. References Adrian, M., Zaharescu, T., Jipa, S., Silva, L.E., Oros, C., 2008. The assessment of thermal and radiation stability of UHMWPE. J. Optoelectron. Adv. Mater. 10, 826e829. Affatato, S., Bordini, B., Fagnano, C., Taddei, P., Tinti, A., Toni, A., 2002. Effects of the sterilisation method on the wear of UHMWPE acetabular cups tested in a hip joint simulator. Biomaterials 23, 1439e1446. Nondegradable synthetic polymers for medical devices and implants 57 Affatato, S., Bracco, P., Costa, L., Villa, T., Quaglini, V., Toni, A., 2012. In vitro wear performance of standard, crosslinked, and vitamin-E-blended UHMWPE. J. Biomed. Mater. Res. Part A 100A, 554e560. Anderson, J.M., Rodriguez, A., Chang, D.T., 2008. Foreign body reaction to biomaterials. Semin. Immunol. 20, 86e100. Bellon, J.M., 2009. Role of the new lightweight prostheses in improving hernia repair. Cir. Esp. 85, 268e273. Berry, D.J., Currier, B.H., Mayor, M.B., Collier, J.P., 2012. Gamma-irradiation sterilization in an inert environment: a partial solution. Clin. Orthop. Relat. Res. 470, 1805e1813. Besong, A.A., Tipper, J.L., Ingham, E., Al, E., 1998. Quantitative comparison of wear debries from UHMWPE that has and has not been sterilized by gamma irradiation. J. Bone Jt. Surg. 80, 340e344. Bracco, P., Del Prever, E.M.B., Cannas, M., Luda, M.P., Costa, L., 2006. Oxidation behaviour in prosthetic UHMWPE components sterilised with high energy radiation in a low-oxygen environment. Polym. Degrad. Stab. 91, 2030e2038. Bracco, P., Oral, E., 2011. Vitamin E-stabilized UHMWPE for total joint implants: a review. Clin. Orthop. Relat. Res. 469, 2286e2293. Brandon, H.J., Jerina, K.L., Wolf, C.J., Young, V.L., 2003. Biodurability of retrieved silicone get breast implants. Plast. Reconstr. Surg. 11. Brook, M.A., 2006. Platinum in silicone breast implants. Biomaterials 27, 3274e3286. Brown, S.A., Hastings, R.S., Mason, J.J., Moet, A., 1990. Characterisation of short-fibre reinforced thermoplastics for fracture fixation devices. Biomaterials 11, 541e547. Carpentieri, I., Brunella, V., Bracco, P., Paganini, M.C., Del Prever, E.M.B., Luda, M.P., Bonomi, S., Costa, L., 2011. Post-irradiation oxidation of different polyethylenes. Polym. Degrad. Stab. 96, 624e629. Charnley, J., 1960. Bone cements and fillers: a review. J. Bone Jt. Surg. 42-B, 28. Choi, T., Weksler, J., Padsalgikar, A., Hernandez, R., Runt, J., 2009. Polydimethylsiloxanebased polyurethanes: phase-separated morphology and in vitro oxidative biostability. Aust. J. Chem. 62, 794e798. Christenson, E.M., Anderson, J.M., Hittner, A., 2007. Biodegradation mechanisms of polyurethane elastomers. Corros. Eng. Sci. Technol. 42, 312e323. Christenson, E.M., Dadsetan, M., Wiggins, M., Anderson, J.M., Hiltner, A., 2004. Poly(carbonate urethane) and poly(ether urethane) biodegradation: in vivo studies. J. Biomed. Mater. Res. Part A 69A, 407e416. Contreras-Garcia, A., Bucio, E., Concheiro, A., Alvarez-Lorenzo, C., 2011. Surface functionalization of polypropylene devices with hemocompatible DMAAm and NIPAAm grafts for norfloxacin sustained release. J. Bioact. Compat. Polym. 26, 405e419. Coote, C.F., Hamilton, J.V., Mcgimpsey, W.G., Thompson, R.W., 2000. Oxidation of gammairradiated ultrahigh molecular weight polyethylene. J. Appl. Polym. Sci. 77, 2525e2542. Costa, L., Bracco, P., Brach, E.M., Luda, M.P., 2002. Oxidation in prosthetic UHMWPE. Abstr. Pap. Am. Chem. Soc. 224, U507. Costa, L., Carpentieri, I., Bracco, P., 2008. Post electron-beam irradiation oxidation of orthopaedic UHMWPE. Polym. Degrad. Stab. 93, 1695e1703. Costa, L., Luda, M.P., Trossarelli, L., Del Prever, E.M.B., Crova, M., Gallinaro, P., 1998. Oxidation in orthopaedic UHMWPE sterilized by gamma-radiation and ethylene oxide. Biomaterials 19, 659e668. Curtis, J., Colas, A., 2004a. Medical Applications of Silicones. Academic Press. Curtis, J., Colas, A., 2004b. Silicone Biomaterials: History and Chemistry & Medical Applications of Silicones. Academic Press. 58 Biosynthetic Polymers for Medical Applications Curtis, J.M., Peters, Y.A., Swarthout, D.E., Kennan, J.J., Vandyke, M.E., 2000. Mechanical and chemical analysis of retrieved breast implants demonstrate material durability. In: Transactions of the Sixth World Biomaterials Congress. Del Prever, E.M.B., Bistolfi, A., Bracco, P., Costa, L., 2009. UHMWPE for arthroplasty: past or future. J. Orthop. Traumatol. 10, 1e8. De Nicola, R.R., 1950. Permanent artificial (silicone) urethra. J. Urol. 63, 168e172. Desai, S.M., Solanky, S.S., Mandale, A.B., Rathore, K., Singh, R.P., 2003. Controlled grafting of N-isopropyl acrylamide brushes onto self-standing isotactic polypropylene thin films: surface initiated atom transfer radical polymerization. Polymer 44, 7645e7649. DiMaio, F.R., 2002. The science of bone cement: a historical review. Orthopedics 25, 1399e1407. Friedman, R.M., 2010. Silicon Breast Implants [Online]. Available: http://www.plasticsurgerydallas. com/blog/which-should-i-choose-saline-breast-implants-or-silicone-gel-breast-implants/. Gautam, R., Singh, R.D., Sharma, V.P., Siddhartha, R., Chand, P., Kumar, R., 2012. Biocompatibility of polymethylmethacrylate resins used in dentistry. J. Biomed. Mater. Res. Part B Appl. Biomater. 100B, 1444e1450. Gomathi, N., Rajasekar, R., Babu, R.R., Mishra, D., Neogi, S., 2012. Development of bio/blood compatible polypropylene through low pressure nitrogen plasma surface modification. Mater. Sci. Eng. C Mater. Biol. Appl. 32, 1767e1778. Grant, D.N., Benson, J., Cozad, M.J., Whelove, O.E., Bachman, S.L., Ramshaw, B.J., Grant, D.A., Grant, S.A., 2011. Conjugation of gold nanoparticles to polypropylene mesh for enhanced biocompatibility. J. Mater. Sci. Mater. Med. 22, 2803e2812. Gunatillake, P.A., Martin, D.J., Meijs, G.F., Mccarthy, S.J., Adhikari, R., 2003. Designing biostable polyurethane elastomers for biomedical implants. Aust. J. Chem. 56, 545e557. Gupta, B., Saxena, S., Ray, A., 2008. Plasma induced graft polymerization of acrylic acid onto polypropylene monofilament. J. Appl. Polym. Sci. 107, 324e330. Hamdan, S., Swallowe, G.M., 1996. Strain rate and temperature dependence of the mechanical properties of polyetherketone polyetheretherketone. J. Mater. Sci. 31, 1415e1423. Handbook of Polymer, Copyright, 2012. ChemTec Publishing, p. 482. Hazer, D.B., Kilicay, E., Hazer, B., 2012. Poly(3-hydroxyalkanoate)s: diversification and biomedical applications. A state of the art review. Mater. Sci. Eng. C Mater. Biol. Appl. 32, 637e647. Hendel, G., Garti, A., Al, E., 2003. Fracture of the central polyethylene tibial spine in posterior stabilized total knee arthroplasty. J. Arthroplasty 15, 672e674. Hunter, A., Archer, C.W., Walker, P.S., Blunn, G.W., 1995. Attachment and proliferation of osteoblasts and fibroblasts on biomaterials for orthopedic use. Biomaterials 16, 635e646. Jaeblon, T., 2010. Polymethylmethacrylate: properties and contemporary uses in orthopaedics. J. Am. Acad. Orthop. Surg. 18, 297e305. Kaneeda, T., Utsumi, M., Watanabe, E., Aspe, 1999. Effects of gamma irradiation and accelerated aging on wear properties of UHMWPE artificial joint. In: Proceedings of the Fourteenth Annual Meeting of the American Society for Precision Engineering, pp. 167e170. Kenny, S.M., Buggy, M., 2003. Bone cements and fillers: a review. J. Mater. Sci. Mater. Med. 14, 923e938. Klosterhalfen, B., Junge, K., Klinge, U., 2005. The lightweight and large porous mesh concept for hernia repair. Expert Rev. Med. Devices 2, 103e117. Kuehn, K.D., Ege, W., Gopp, U., 2005. Acrylic bone cements: composition and properties. Orthop. Clin. North Am. 36, 17e28. Kuhn, K.-D., 2000. Bone Cements. Springer Valley, Heidelberg, Germany. Nondegradable synthetic polymers for medical devices and implants 59 Kumakura, T., Puppulin, L., Yamamoto, K., Takahashi, Y., Pezzotti, G., 2009. In-depth oxidation and strain profiles in UHMWPE acetabular cups non-destructively studied by confocal Raman microprobe spectroscopy. J. Biomater. Sci. Polym. Ed. 20, 1809e1822. Kurtz, S.M., 2004. Ultrahigh Molecular Weight Polyethylenes (UHMWPE). Academic Press, London. Kurtz, S., Muratoglu, O., Mark, E., Al, E., 1999a. Advances in the processing, sterilization, and crosslinking of ultra-high molecular weight polyethylene for total joint arthroplasty. Biomaterials 20, 1659e1688. Kurtz, S., Van Ooij, A., Al, E., 2006a. Polyethylene wear and rim fracture in total disc arthroplasty. Spine J. 7, 12e21. Kurtz, S.M., 2009. UHDMWPE Biomaterials Handbook, Chapter2. Academic Press, London, UK, pp. 7e19. Kurtz, S.M., 2012. PEEK Biomaterials Handbook. William Andrew Publication, Oxford, UK. Kurtz, S.M., Devine, J.N., 2007. PEEK biomaterials in trauma, orthopedic, and spinal implants. Biomaterials 28, 4845e4869. Kurtz, S.M., Hozack, W.J., Purtill, J.J., Marcolongo, M., Kraay, M.J., Goldberg, V.M., Sharkey, P.F., Parvizi, J., Rimnac, C.M., Edidin, A.A., 2006b. 2006 Otto Aufranc Award paper e significance of in vivo degradation for polyethylene in total hip arthroplasty. Clin. Orthop. Relat. Res. 47e57. Kurtz, S.M., Muratoglu, O.K., Evans, M., Edidin, A.A., 1999b. Advances in the processing, sterilization, and crosslinking of ultra-high molecular weight polyethylene for total joint arthroplasty. Biomaterials 20, 1659e1688. Kwarteng, K.B., Starck, C., 1990. Carbon fibre reinforced PEEK (APC-2/AS-4) composites for orthopaedic implants. SAMPE Q. 10e14. Lafay, H., 1957. A father’s last chance invention saves his son. Reader’s Digest. Lamba, K., Woodhouse, K., Coper, S.L., 1998. Polyurethanes in Biomedical Applications. CRC Press, New York. Laredo, J.D., Hamze, B., 2004. Complications of percutaneous vertebroplasty and their prevention. Skeletal Radiol. 33, 493e505. Lewis, G., 2001. Properties of crosslinked ultra-high-molecular-weight polyethylene. Biomaterials 22, 371e401. Lewis, G., 2011. Viscoelastic properties of injectable bone cements for orthopaedic applications: state-of-the-art review. J. Biomed. Mater. Res. Part B Appl. Biomater. 98B, 171e191. Li, H.M., Fouracre, R.A., Given, M.J., Banforn, H.M., Wysocki, S., Karolczak, S., 1999. Effects on polyetheretherketone and polyether sulfone of electron and gamma irradiation. Dielectr. Electr. Insul. IEEE Trans. 6, 295e303. Lieberman, I., Dudeney, S., Reinhardt, M.-K., Bell, G., 2002. Initial outcome and efficacy of “kyphoplasty” in the treatment of painful osteoporotic vertebral compression fractures. Spine 26, 1631e1638. Liu, Q.Y., Jiang, L., Shi, R., Zhang, L.Q., 2012. Synthesis, preparation, in vitro degradation, and application of novel degradable bioelastomers-a review. Prog. Polym. Sci. 37, 715e765. Livermore, C., Voldman, J., 2004. Design and Fabrication of Microelectromechanical Devices Material Property Database, 2004 [cit 14.05.2013]. Majercik, S., Tsikitis, V., Iannitti, D.A., 2006. Strength and tissue attachment to mesh after ventral hernia repair with synthetic composite mesh in a porcine model. Surg. Endosc. 20, 1671e1674. Mariconda, M., Lotti, G., Al, E., 2000. Fracture of posterior-stabilized tibial insert in a genesis knee prosthesis. J. Arthroplasty 15, 529e530. 60 Biosynthetic Polymers for Medical Applications Martina, M., Hutmacher, D.W., 2007. Biodegradable polymers applied in tissue engineering research: a review. Polym. Int. 56, 145e157. Mata, A., Fleischman, A.J., Roy, S., 2005. Biomed. Microdevices 7, 281e293. Mckellop, H., Shen, F.W., Lu, B., Campbell, P., Salovey, R., 2000. Effect of sterilization method and other modifications on the wear resistance of acetabular cups made of ultrahigh molecular weight polyethylene e a hip-simulator study. J. Bone Jt. Surg. Am. Vol. 82A, 1708e1725. Mcmillin, C.R., 2006. Biomedical applications of rubbers and elastomers. Rubber Chem. Technol. 79, 500e519. Medel, F.J., Garcia-Alvarez, F., Gomez-Barrena, E., Puertolas, J.A., 2005. Microstructure changes of extruded ultra high molecular weight polyethylene after gamma irradiation and shelf-aging. Polym. Degrad. Stab. 88, 435e443. Mehmood, M.S., Hafeez, U., Jahan, M.S., Mishra, S., Walters, B.M., Ikram, M., 2012. The effect of high dose of gamma-irradiation on residual radicals concentration in ultra-high molecular weight polyethylene (UHMWPE) in the presence of vitamin E. Polym. Sci. Ser. A 54, 343e348. Meintjes, J., Yan, S., Zhou, L., Zheng, S.S., Zheng, M.H., 2011. Synthetic, biological and composite scaffolds for abdominal wall reconstruction. Expert Rev. Med. Devices 8, 275e288. Nair, L.S., Laurencin, C.T., 2007. Biodegradable polymers as biomaterials. Prog. Polym. Sci. 32, 762e798. Nayak, R., Kyratzis, I.L., Truong, Y.B., Padhye, R., Arnold, L., Peeters, G., Nichols, L., O’shea, M., 2012. Fabrication and characterisation of nanofibres by meltblowing and melt electrospinning. In: Chen, W.Z., Xu, X.P., Dai, P.Q., Chen, Y.L. (Eds.), Advanced Manufacturing Technology, Pts 1e4. Noll, W., 1968. Chemistry and Technology of Silicones. New York. Ojha, U., Kulkarni, P., Faust, R., 2009. Syntheses and characterization of novel biostable polyisobutylene based thermoplastic polyurethanes. Polymer 50, 3448e3457. Ortel, G., 1994. Polyurethane Handbook. Hanser Gardner, Berlin. Petillo, O., Peluso, G., Ambrosio, L., Nicolais, L., Kao, W.J., Anderson, J.M., 1994. In vivo induction of macrophage Ia antigen (MHC Class II) expression by biomedical polymers in the cage implant system. J. Biomed. Mater. Res. 28, 635e646. Pinchuk, L., 1995. A review of the biostability and carcinogenicity of polyurethanes in medicine and the new generation of biostable polyurethanes. J. Biomater. Sci. Polym. Ed. 6, 225e267. Place, E.S., George, J.H., Williams, C.K., Stevens, M.M., 2009. Synthetic polymer scaffolds for tissue engineering. Chem. Soc. Rev. 38, 1139e1151. Pruitt, L.A., 2005. Deformation, yielding, fracture and fatigue behavior of conventional and highly cross-linked ultra high molecular weight polyethylene. Biomaterials 26, 905e915. Rae, P.J., Brown, E.N., Orler, E.B., 2007. The mechanical properties of poly(ether-ether-ketone) (PEEK) with emphasis on the large compressive strain response. Polymer 48, 598e615. Sanders, D.L., Kingsnorth, A.N., 2012. Prosthetic mesh materials used in hernia surgery. Expert Rev. Med. Devices 9, 159e179. Santerre, J.P., Woodhouse, K., Laroche, G., Labow, R.S., 2005. Understanding the biodegradation of polyurethanes: from classical implants to tissue engineering materials. Biomaterials 26, 7457e7470. Santerre, P., Duguay, D., Labow, R., Brash, J., 1995. Interaction of hydrolytic enzymes at an aqueous-polyurethane interface. Proteins at interfaces II. J. Am. Chem. Soc. 352e370. Nondegradable synthetic polymers for medical devices and implants 61 Saxena, S., Ray, A.R., Kapil, A., Pavon-Djavid, G., Letourneur, D., Gupta, B., Meddahi-Pelle, A., 2011. Development of a new polypropylene-based suture: plasma grafting, surface treatment, characterization, and biocompatibility studies. Macromol. Biosci. 11, 373e382. Scotchford, C.A., Garle, M.J., Batchelor, J., Bradley, J., Grant, D.M., 2003. Use of a novel carbon fibre composite material for the femoral stem component of a THR system: in vitro biological assessment. Biomaterials 24, 4871e4879. Shibata, N., Tomita, N., 2005. The anti-oxidative properties of alpha-tocopherol in gammairradiated UHMWPE with respect to fatigue and oxidation resistance. Biomaterials 26, 5755e5762. Simmons, A., Hyvarinen, J., Poole-Warren, L., 2006. The effect of sterilisation on a poly (dimethylsiloxane)/poly(hexamethylene oxide) mixed macrodiol-based polyurethane elastomer. Biomaterials 27, 4484e4497. Skinner, H.B., 1988. Composite technology for total hip arthroplasty. Clin. Orthop. Relat. Res. 235, 224e236. Sobieraj, M.C., Rimnac, C.M., 2009. Ultra high molecular weight polyethylene: mechanics, morphology, and clinical behavior. J. Mech. Behav. Biomed. Mater. 2, 433e443. Song, L.J., Zhao, J., Yang, H.W., Jin, J., Li, X.M., Stagnaro, P., Yin, J.H., 2011. Biocompatibility of polypropylene non-woven fabric membrane via UV-induced graft polymerization of 2-acrylamido-2-methylpropane sulfonic acid. Appl. Surf. Sci. 258, 425e430. Speckhard, T., Ver Strate, G., Gibson, P., Sl, C., 1983. Properties of polyisobutylenepolyurethane block copolymers: I. Macroglycols from ozonolysis of isobutylene copolymer. Polym. Eng. Sci. 23, 337e349. Stokes, K., Mcvennes, R., 1995. Polyurethane elastomer biostability. J. Biomater. Appl. 9, 321e335. Sugano, N., Saito, A., Yamamoto, T., Nishii, T., Yau, S.S., Wang, A., 2004. Analysis of a retrieved UHMWPE acetabular cup crosslinked in air with 1000 kGy of gamma radiation. J. Orthop. Res. 22, 828e831. Szycher, M., Reed, A.M., Soc Plast Engineers, I.N.C., 1996. Medical-grade polyurethanes: a critical review. In: Antec ’96: Plastics e Racing into the Future, Vols I-III: Vol I: Processing; Vol II: Materials; Vol III: Spacial Areas, 42, pp. 2758e2766. Takahara, A., Coury, A., Hergenrother, R., Cooper, S., 1991. Effect of soft segment chemistry on the biostability of segmented polyurethanes. I. In vitro oxidation. J. Biomed. Mater. Res. 25, 341e356. Toth, J.M., 2012. Biocompatibility of Polyaryletheretherketone Polymers. Elsevier Inc. Ulery, B.D., Nair, L.S., Laurencin, C.T., 2011. Biomedical applications of biodegradable polymers. J. Polym. Sci. Part B Polym. Phys. 49, 832e864. Vallittu, P.K., Ruyter, I.E., Buykuilmaz, S., 1998. Effect of polymerization temperature and time on the residual monomer content of denture base polymers. Eur. J. Oral Sci. 106, 588e593. Ward, R., Anderson, J., Mcvenes, R., Stokes, K., 2006a. In vivo biostability of polysiloxane polyether polyurethanes: resistance to biologic oxidation and stress cracking. J. Biomed. Mater. Res. Part A 77A, 580e589. Ward, R., Anderson, J., Mcvenes, R., Stokes, K., 2006b. In vivo biostability of shore 55D polyether polyurethanes with and without fluoropolymer surface modifying endgroups. J. Biomed. Mater. Res. Part A 79A, 836e845. Wenz, L.M., Merritt, K., Brown, S.A., Moet, A., Steffe, A.D., 1990. In vitro biocompatibility of polyetheretherketone and polysulfone composites. J. Biomed. Mater. Res. 24, 207e215. Wheatley, D.J., Bernacca, G.M., Tolland, M.M., O’connor, B., Fisher, J., Williams, D.F., 2001. Hydrodynamic function of a biostable polyurethane flexible heart valve after six months in sheep. Int. J. Artif. Organs 24, 95e101. 62 Biosynthetic Polymers for Medical Applications Wheatley, D.J., Raco, L., Bernacca, G.M., Sim, I., Belcher, P.R., Boyd, J.S., 2000. Polyurethane: material for the next generation of heart valve prostheses? Eur. J. Cardiothorac. Surg. 17, 440e447. Williams, D.F., 2003. Revisiting the definition of biocompatibility. Med. Device Technol. 14, 10e13. Williams, D.F., Mcnmara, A., Turner, R.M., 1987. Potential of polyetheretherketone (PEEK) and carbon-fibre-reinforced PEEK in medical applications. J. Mater. Sci. Lett. 6, 188e190. Wolf, C., Maninger, J., Lederer, K., Fruhwirth-Smounig, H., Gamse, T., Marr, R., 2006. Stabilisation of crosslinked ultra-high molecular weight polyethylene (UHMW-PE)-acetabular components with alpha-tocopherol. J. Mater. Sci. Mater. Med. 17, 1323e1331. Xiong, D., Ma, R., Lin, J., Wang, N., Jin, Z.M., 2007. Tribological properties and structure of ultra-high molecular weight polyethylene after gamma irradiation. Proc. Inst. Mech. Eng. Part J J. Eng. Tribol. 221, 315e320. Xiong, L., Xiong, D.S., 2012. The influence of irradiation dose on mechanical properties and wear resistance of molded and extruded ultra high molecular weight polyethylene. J. Mech. Behav. Biomed. Mater. 9, 73e82. Yang, Q., Xu, Z.K., Dai, Z.W., Wang, J.L., Ulbricht, M., 2005. Surface modification of polypropylene microporous membranes with a novel glycopolymer. Chem. Mater. 17, 3050e3058. Zhao, J., Shi, Q., Luan, S.F., Song, L.J., Yang, H.W., Stagnaro, P., Yin, J.H., 2012. Polypropylene non-woven fabric membrane via surface modification with biomimetic phosphorylcholine in Ce(IV)/HNO3 redox system. Mater. Sci. Eng. C Mater. Biol. Appl. 32, 1785e1789. Zhao, J., Shi, Q.A., Luan, S.F., Song, L.J., Yang, H.W., Shi, H.C., Jin, J., Li, X.L., Yin, J.H., Stagnaro, P., 2011. Improved biocompatibility and antifouling property of polypropylene non-woven fabric membrane by surface grafting zwitterionic polymer. J. Membr. Sci. 369, 5e12. Zimmer, 2012. Vitamin E Containing UHMWPE in Continuum Acetabular Shell System (Vivacit). Zimmer, Inc. http://www.zimmer.com/content/pdf/en-US/zimmer_continuum_ acetabular_system_with_vivacit-e_surgical_technique.pdf.