Biomedical Application of Commercial Polymers and Novel

July/August 2004

Published by the American Chemical Society

Volume 5, Number 4

© Copyright 2004 by the American Chemical Society

Reviews Biomedical Application of Commercial Polymers and Novel Polyisobutylene-Based Thermoplastic Elastomers for Soft Tissue Replacement† Judit E. Puskas* and Yaohong Chen Department of Chemical & Biochemical Engineering, The University of Western Ontario, London, Ontario, Canada N6A 5B9 Received December 8, 2003; Revised Manuscript Received February 26, 2004

Novel polyisobutylene-based thermoplastic elastomers are introduced as prospective implant materials for soft tissue replacement and reconstruction. In comparison, poly(ethylene terephthalate) (PET), poly(tetrafluoroethylene) (PTFE), polypropylene (PP), polyurethanes (PU), and silicones are outlined from wellestablished implant history as being relatively inert and biocompatible biomaterials for soft tissue replacement, especially in vascular grafts and breast implants. Some general considerations for the design and development of polymers for soft tissue replacement are reviewed from the viewpoint of material science and engineering, with special attention to synthetic materials used in vascular grafts and breast implants. Introduction Bioartificial repair of human organs has been an area of great interest for medical researchers for a long time. The successful application of a medical device or implant in a clinical setting not only depends on the properties of the biomaterials used to fabricate the device or implant, but also on numerous other factors, including the manufacturing and processing history of the materials, and the specific application in which the device is used.1 Most tissues other than bone and cartilage are of the soft category. Implants do not generally interface directly with blood, except for those in the cardiovascular system. Hence, soft tissue replacements include blood-interfacing implants such as vascular grafts as well as non-blood-interfacing implants such as breast implants. Although the ideal “material” to repair human tissues would be the regenerated tissue itself, clinical application of this approach will remain out of reach for a † A version of this paper (#40) was presented at the 160th Meeting of the Rubber Division of the American Chemical Society, San Francisco, CA, April 28-30, 2003, and received the “Best Paper - Honorable Mention” Award. * To whom correspondence should be addressed.

long time; thus, synthetic materials will have to do.2 The development of improved materials for biomedical applications will require close collaboration among chemists, biologists, material scientists, engineers, and clinicians. The vast expertise of scientists and engineers in the rubber industry would help this cause tremendously. Materials used in soft tissue replacement and reconstruction can be classified into four categories: (1) artificial nonresorbable and (2) artificial resorbable polymers, (3) semi-natural or derivatized polymeric materials, and (4) native or natural materials such as grafts and transplants. Elastomers are predominantly used in applications that require compliance with soft or cardiovascular tissue.3,4 The required properties of polymeric biomaterials are similar to other biomaterials, that is, biocompatibility, sterilizability, adequate mechanical and physical properties, and good processability for ease of manufacturing. Cost generally has not been a factor, but with the escalating expenses of medical care and the aging population, this will most probably change. The definition of biocompatibility is vague, compared to definitions material scientists and engineers are used to.

10.1021/bm034513k CCC: $27.50 © 2004 American Chemical Society Published on Web 04/06/2004

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Table 1. Autian Test Scheme Performed for Evaluation of Plastic Devices to Be Used in Pharmacy or Medical Practice6 test for sterilization test for pyrogenicity test for skin and tissue sensitivity test for toxicity (both acute and chronic) test for compatibility with drug products 1. leaching of an ingredient into drug product 2. binding or sorption of a drug ingredient by the plastics 3. permeation of gases into drug systems through the plastic 4. permeation of ingredients from the drug system through the plastic 5. alteration of the potency of the active ingredient after contact with plastic 6. alteration of the physical and chemical properties of the plastic by the drug system 7. other tests

Generally, it means acceptance of an artificial implant by the surrounding tissues, and the body as a whole. The European Society for Biomaterials (ESB) arrived at the following definition: “Biocompatibility: The ability of a material to perform with an appropriate host response in a specific application.” 5 The implant should be compatible with tissues in terms of mechanical, chemical, surface, and pharmacological properties. The evaluation of biocompatibility of synthetic materials, i.e., the evaluation of the suitability of a synthetic material for use in implantable medical devices, has evolved over approximately the last 50 years. Although most biomaterials and implants are inert and nontoxic, they may elicit an adverse response from the body and vice versa. In 1961, Autian was the first to describe a systematic approach to the evaluation of materials in the determination of their biological response.6 The scheme for testing plastics used in biomedical devices, as shown in Table 1, was codified by the U.S. Pharmacopoeia (USP) in 1965 and accepted as the norm for evaluating the biocompatibility of materials by the Food, Drug and Cosmetic Act in 1976.7 The Autian test scheme is still described in the current USP Monograph 88 concerning in vivo biological reactivity tests. Since the biocompatibility of materials is a complex issue, there is a need to broaden the spectrum of the types of tests beyond the concepts in USP Monograph 88; for instance, tests for swelling. There is also a need for closer collaboration between material scientists and medical professionals. As it was recently pointed out by Kennedy,3 “...the most influential contemporary treatise on “Biomaterials Science” edited by Rattner et al.8 does not even have an entry on rubbers or elastomers”. The authors of this review find this crucial since soft tissue replacement is critical to women. For instance, in the U.S., demand for cosmetic breast implants increased dramatically from 32 607 in 1992 to 167 318 in 1999, a 6-fold increase! Demand for post-mastectomy implants also increased dramatically, from 29 607 in 1992 to 82 975 in 1999, which is a 2.5-fold increase, despite the controversy surrounding the silicone implants.9 Without being judgmental as to the reasons why women are desperate to get breast implants, we feel that more efforts should be directed to this area. Today there is no alternative to silicone-based breast implants: a situation unacceptable to more than half of the human population. This review hopes to be a small contribution toward changing this situation. In this review, general considerations for the design and development of polymers for soft tissue replacement will

Figure 1. Schematic illustration of the layered construction of an arterial wall. Table 2. Mechanical Properties of Selected Soft Human Tissues10 tensile strength (MPa) tissue inferior vena cava ascending aorta ureter trachea

elongation (%)

transverse longitudinal transverse longitudinal 3.03 1.07 0.47 0.35

1.17 0.069 1.03 2.16

51 77 98 81

84 81 36 61

be reviewed from the viewpoint of material science and engineering, with special attention to synthetic materials used in vascular grafts and breast implants. We will also introduce novel polyisobutylene (PIB)-based polymers as prospective biomaterials. This will be followed by the evaluation of novel PIB-based materials as potential materials for soft tissue replacement, with emphasis on vascular grafts and breast implants. Blood-Interfacing Implants: Vascular Grafts Soft tissues are biological composite structures. The mechanical properties of some soft human tissues are summarized in Table 2.10 When considering replacement of these tissues, very complex requirements should be met. For instance, arteries are anisotropic, pulsatile, compliant, and thrombosis-resistant. Moreover, the impedance, that is, the resistance to pulsatile flow, responds to the frequency of blood pressure and flow rate fluctuation. Figure 1 shows the cross-section of an arterial wall. It is composed of the Adventitia or outer layer of collagen fibers within an extracellular matrix along with elastin and fibroblasts, the Media or middle layer of vascular smooth muscle tissue and elastin, and the Intima or inner monolayer of endothelial cells.11 Collagen in fibrous form is ideally designed to support tensile loads, to which the structural composites are subjected. However, its precise role in the biological composite largely depends on the nature of loading when performing the various functions of the individual soft tissues.12 Elastin,

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Figure 2. Tension-length diagram of human external iliac artery; collagen, after selective digestion of the elastin fibers with trypsin; elastin, after selective digestion of the collageneous fibers with formic acid.13 Table 3. Mechanical Properties of Vascular Wall Components Compared to Stainless Steel and Cured Natural Rubber14

substances elastin collagen smooth muscle relaxed contracted natural rubber (HEVEA) steel (X12CrMoS17)

Young’s modulus (MPa)

tensile strength (MPa)

max. extension (%)

0.3-0.6 0.36-4.4 100-220 1 × 102-2.9 × 103 5-500 5-50 6 × 10-3 1 × 10-2-1.27 1.3 2.16 × 105

300 300 17-25 750-850 540-740 16

on the other hand, behaves more like an elastomeric material. Figure 2 shows stress-strain plots of a human iliac artery, together with those of elastin obtained by selective digestion of the collagen fibers with trypsin, and collagen obtained by selective digestion of the elastin by formic acid.13 This figure demonstrates the drastically different roles of the components of the biological composite material of the arterial wall. Elastin provides the initial stiffness of the curve, which is quite low, and collagen provides a much higher stiffness at higher stress levels when needed. During the transition from low to high stiffness, collagen fibers become fully elongated. Consequently, elastic force in the arterial wall at normal physiological pressure depends on contributions from both elastin and collagen fibers. Table 3 lists mechanical properties of the arterial wall components, in comparison with steel and crosslinked natural rubber.14 It is apparent that collagen has a tensile strength comparable to steel and is much stiffer than crosslinked natural rubber. Elastin, on the other hand, is less elastomeric than crosslinked natural rubber. It must be kept in mind that the values in this table are not necessarily representative of the real properties of the arterial wall in vivo.11,15 As shown in Figure 1, the inner layer, Intima, is lined with a monolayer of endothelial cells (also known as the endothelium). The endothelium is the interfacing structure between the blood vessel lumen and the blood, and participates in the processes of vascular permeability, and also in hemostasis, which is the process that the body uses to stop the flow of blood when the vascular system is damaged. The vascular smooth muscle

and elastin in the Media contribute to the bulk of the mechanical strength of an artery, and provide its ability to respond to external stimuli such as pressure. The outer layer, Adventitia, provides stiffness under high tensile loading and has a characteristic laminate structure. Because of its threelayer construction, the artery has the ability to work in concert with heart function and can expand during systole and contract at diastole in order to regulate the blood flow through the vascular system. Vascular grafts are implanted into the body in order to replace damaged vessels or bypass blocked arteries and veins. They are simply conduits for blood flow. Replacement of the artery, a very complex composite material, is a daunting task. About 5% of vascular grafts used today are of natural origin such as saphenous or umbilical veins and bovine heterografts. Although synthetic vessel replacements are available, they are limited to large diameter vessels. To date, there is no satisfactory synthetic small diameter ( 10 mm) and 20% of the medium-size (10> D > 4) grafts, with Gore-Tex (PTFE or extended fluoropolymer) making up the remaining 75% of this latter. The development of synthetic grafts began with the use of nonbiological tubes during the late 40s. These included paraffin-lined silver, vitallium and polished methyl methacrylate as well as polyethylene tubes.17 However, these grafts functioned only in short term implantation trials and failed due to suture line rupture and early graft occlusion. In the early 50s, the concept of porous, fabric vascular grafts was introduced utilizing a poly(vinyl chloride)-polyacrylonitrile copolymer, otherwise known as Vinyon-N.18,19 From the late 50s to early 60s, numerous fabrics were tested experimentally and clinically. However, with the exception of Dacron and Teflon, all the fabrics were eliminated due to lack of stability and durability. During this period, Dacron was the leading material used for vascular reconstruction of all the major large arteries. Nonfabric Teflon grafts, such as expanded poly(tetrafluoroethylene) (e-PTFE) grafts, were developed during the late 60s and first used as small arterial substitutes (Gor-Tex). Two key factors are believed to cause graft failure.20,21 One is intimal tissue overgrowth (hyperplasia) that develops in the artery just proximal and distal to the artificially created connection (anastomosis) between the rigid graft and the compliant artery. Another is the loss of self-cleaning quality as the graft becomes stiff and nonpulsatile, thus pulsation can no longer prevent deposits of fibril and platelets forming on the walls. This relates to blood compatibility. In general, blood compatibility involves two aspects, interfacial and mechanical. Thus, it could be expected that mismatch not only in mechanical, but also in interfacial properties between the graft and host vessel will affect the patency of vascular grafts17,22 (the patency of a vascular graft is defined as the state of being freely open). Clearly, endothelial cells generate the most thrombosis-resistant and tissue proliferation-resistant blood-graft interface. Hence, the development of biocompatible surface modifications and controlled release modalities for vascular implant materials is critical for achieving

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Table 4. Compliance vs Patency Rate of Various Natural and Synthetic Grafts25 graft type

compliance (×10-2 % mmHg1-)

patency rate (% in 2 years)

host artery saphenous vein umbilical vein bovine heterograft Dacron PTFE

5.9 ( 0.5 4.4 ( 0.8 3.7 ( 0.5 2.6 ( 0.3 1.9 ( 0.3 1.6 ( 0.2

84 80 59 42 42

controllable modulation of cellular interactions at the tissuebiomaterial interface. Antithrombogenic coatings such as heparin and benzalkonium chloride were developed, and velour-lined vascular patches reinforced with polyester were suggested.23 Techniques to enhance endothelial cell seeding in order to prevent thrombogenesis were first developed in the 70s. Poly(ether urethanes) were also proposed at this time for use in vascular reconstructions. However, within six months, surface cracking was observed in these materials. The main degradation sites were found in the vicinity of the ether linkage of the polymer.24 The synthetic polymers used today will be discussed in more detail in the section discussing polymers in soft tissue replacement. Synthetic grafts save lives, but it should be noted that, without intervention, the average patency life of vascular grafts is currently 3 years.21 Table 4 shows the relationship between compliance and patency rate of various natural and synthetic grafts.25 The patency rate indicates the percentage of grafts observed to be patent (fully open) for the defined length of time. Vascular compliance is defined as the slope of the pressure-volume relationship and is only a rough index of apparent stiffness and a measure of ease with which a structure can be deformed. It is the ratio of incremental volume change to incremental transmural pressure change, expressed as percent of diameter change per millimeter of mercury. The value depends on the configuration, diameter and thickness of the arterial wall as well as on fundamental material properties such as Young’s modulus and Poisson’s ratio.26 It is an experimentally observable fact that when a test specimen of an isotropically elastic material is subjected to uniaxial stress, the specimen not only deforms in the direction of the stress, but also exhibits deformation of the opposite sign in the perpendicular direction. Poisson’s ratio and Young’s modulus are two fundamental elastic constants relating stress to strain in a biaxial stress field. The Young modulus characterizes the elastic response of the viscoelastic wall tissue, and the Poisson ratio is, by definition, the absolute value of the ratio of transverse strain to the axial strain in a uniaxially stressed specimen. A Poisson ratio of 0.5 is characteristic of natural arteries. The compliance range of 0.08-0.12% mmHg-1 is typical of arteries, whereas the compliance values of Dacron and PTFE, widely used in vascular grafts, are in the range of 0.014-0.019% mmHg-1.25,27,28 Thus, these grafts are 4-to 8 times less compliant than an artery. Certainly, the compliance of an ideal graft should be identical to that of the native artery; both excessive or diminished compliance are detrimental. From the data in Table 4, it is clear that decreasing

patency correlates with increasing disparity between host artery and graft compliance. It was found that compliance mismatch between the graft and host artery also results in differential mechanical strains and haemodynamic wall shear stress (WSS), due to pulsatile blood flow and the viscoelastic nature of the blood vessel wall itself.28-30 Arteries can adapt their diameter to maintain relatively constant WSS. However, rigid synthetic grafts cannot do this, and an optimum WSS level will not be maintained within the graft unless it is sized properly. It has been determined that there is a thrombotic threshold velocity of blood flow, below which thrombus formation develops. To avoid this, the size of the graft needs to exactly match the size of the host vessel. A synthetic graft that is larger than its host vessel creates a state of low velocity blood flow and shear stress, thus promoting failure. On the other hand, when a graft is too small, the fibrin layer that develops over the artificial surface in the body during healing may obstruct the graft lumen. In addition, the choice of synthetic graft material also causes graft-artery mismatch and leads to late graft failure. Moreover, WSS is known to influence the biochemistry of endothelial cells and the permeability of macromolecules through the arterial wall31-33 and is widely believed to be the principal fluid mechanical mediator of arteriosclerosis, in which the vessel wall is thickened as a result of fat cell deposits.32 Low mean WSS and temporal variation in WSS have been correlated with the distribution of atherosclerotic wall thickening. WSS has also been associated with intimal tissue overgrowth of grafts by several groups.29,34,35 Clearly, the compliant artery is a viscoelastic reservoir absorbing energy during systole and releasing it during diastole; however, the noncompliant graft reduces energy available for distal perfusion by diminishing the pulsatile component of the diastolic shrinking. For elastic vessels, the impedance is dependent on the frequency of the fluctuations in pressure and flow. For rigid grafts with laminar flow and negligible inertial effect, the input impedance is simply the resistance and independent of pressure and flow rate. It was found that impedance change occurring at the interface between the compliant artery and the noncompliant graft results in the propagation of g60% of the original pulsatile energy.36 Change of mechanical characteristics at the interface between host artery and graft also leads to vibratory weakening of the artery wall, loss of endothelial cell viability, and dilatation of a surgically made connection, known as anastomotic aneurysm.32 In addition, calcification of biomaterials is another key factor to be considered, particularly if the materials undergo repeated flexing.37 Approaches to the prevention of calcification, such as localized sustained release of calcium chelating agents, are being studied.38,39 In summary, the key factors that need to be considered when designing the “ideal” vascular graft are mismatch of material properties between the host artery and the graft and surface thrombosis of vascular grafts. The thrombosis of vascular grafts results in a change of mechanical properties promoting late graft failure, whereas artery-graft mismatch of mechanical properties also leads to thrombosis. Hence,


Biomedical Application of Commercial Polymers Table 5. Tensile Strength of Some Commercial Silicone Breast Implant Shells43 implant tensile strength at break (MPa)

Dow Corning Surgitek Dow Corning Surgitek Silastic II SCL standard standard 10.7

8.9

7.1

6.1

the blood compatibility of vascular grafts involves not only the surface interaction of the grafts with blood, but also the mechanical and fatigue properties of the grafts as compared to the host vessel. Non-Blood-Interfacing Implants: Breast Implants When considering replacement or alteration of human breast tissue, one needs to understand the composition and properties of the breast. It is stated that “the materials used in breast implants must have mechanical properties that mimic the human breast”,40,41 but once again the definition is quite vague: “they need to be soft and deformable while maintaining the appropriate shape”. In comparison to blood vessels, no substantial information is available on the “material” properties of the human breast. The topic of breast implants has been extremely controversial over the last number of years and the debate is driven by legal advocacy as well as by “objective” science. It is still not entirely clear how these implants affect the body. The literature is divided on the benefits and complications associated with breast implants. In general, biocompatibility does not just pertain to host response but also to implant response or stability.18,19 It is important to consider both the effects of the physiological environment on the implant and its function and, conversely, the effect of the implant on the physiological environment. A breast implant consists of two or three major components: an elastomeric/rubbery shell, a filler material, and possibly a valve through which the filler material is placed into the shell. The shell provides strength and barrier properties, possibly with a surface layer for improved tissue interaction, whereas the liquid or gel filling supplies bulk and consistency. Materials such as Ivalon (a poly(vinyl alcohol) sponge crosslinked with formaldehyde), Teflon, polyurethanes and silicones have been used for breast implant shells.42 Today silicone rubber is the only “choice” of implant shell available. Silicone-based breast implants can be categorized into three generations. The first generation implants of the 1960s and early 1970s had the thickest shells (>0.0254 cm) and the most viscous gel. As the mechanical properties of silicone improved, the layer thickness was reduced. The second generation, developed in the 1970s and early 1980s, had the thinnest shells and the least viscous gel. This led to rupture of the shell in softer profile implants, which resulted in a general thickening of the shell again, together with efforts to improve the tear strength of silicone rubber. The third generation, developed in the late 1980s and early 1990s, is used currently and has shells of intermediate thickness filled with gels of medium viscosity. The tensile strength of some silicone breast implant shells are given in Table 5.43 The tensile stress at break ranges from 6 to 12 MPa. It is

interesting to note that implant shells were found to experience maximum strains of 1000% or more before rupture. Currently, there are three major types of synthetic breast implants approved by the FDA in the U.S., namely saline filled, silicone gel filled, and a combination of gel and saline filled.42,44,45 Reconstruction using tissue removed from the patient is an additional option (“flap procedure”),45 but synthetic implants are more “popular”. The saline-filled implant has an external silicone shell and is filled with sterile physiological saltwater (0.8 wt % salt). The silicone gelfilled implant also has an external silicone shell but is filled with silicone gel. The silicone gel consists of a silicone polymer network with low molecular weight silicone oil dispersed through the polymer. The amount of low molecular weight material is important as it controls the viscosity and influences gel bleed, i.e., the passage of the filling gel through the wall of the implant’s shell. It is commonly accepted today that implants with gel bleed, can rupture, and that the gel can migrate.46 Understanding implant stability is crucial, because the host response can be affected by changes in the implant over time. The extent of gel bleed is dependent on the molecular weight of the components and the degree of crosslinking of both the gel and shell and the surface area of the prosthesis. Initial swelling upon implantation results in increased permeability, and, hence more gel bleed. It was shown that diffusion of the dimethylsiloxane small molecules out of the envelope can be reduced by using an additional layer of another elastomer, such as poly(methylphenylsiloxane) or fluorosilicone.47,48 However, the shell itself has been found to release low-molecular-weight moieties even without contacting the gel.49-51 Numerous in vitro studies have been carried out to characterize gel bleed composition, bleed amount, and bleed rate.52 There is some debate as to whether these gel bleed rates decrease over time or whether they are relatively constant, because the implant behaves as an infinite sink.53 As early as 1980, implant manufacturers suggested that the only way to prevent gel bleed and the migration of silicone-containing small molecules into the body would be to use a different shell material.54 However, silicone rubber remains the only shell material available today. FDA approval of silicone gel filled implants is currently suspended. Implants filled with a gel based on soybean oil were approved in Europe, but were subsequently removed from the market.45 Polymers Used in Soft Tissue Replacement From the viewpoint of a well-established implant history as being relatively inert and biocompatible, polyesters (PET), fluoropolymers (PTFE), polypropylene (PP), polyurethanes (PU), and silicones have played the most important role in the development of polymeric materials for soft tissue replacement. New biomaterials considered in the literature are usually a modification of these basic structures, and their full review would exceed the limitations of the current general review. Recently a new alternative material, polyisobutylene-polystyrene (PIB-PS) block copolymer emerged, which is still in the developmental stage. This biomaterial represents a conceptually new soft biomaterial, and has an order of magnitude lower permeability than any other crosslinked elastomer, a property critical for potential breast

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Table 6. Selected Properties of Polymers Used in Soft Tissue Replacement in Comparison with PIB-PS Block Polymer55-57 thermoplastic polymer polymer properties tensile strength (MPa) elongation at break (%) hardness (Shore A) compressive strength (MPa) shear modulus (GPa) flexural modulus (GPa) Young’s modulus (GPa) Possion’s ratio Tm (°C) (crystalline) Tg (°C) (amorphous) permeability coefficient,a 20-25°C (×1018 m2/s Pa) aH

2O vapor transmission rate, 37-40°C (x108g./m.s)

CO2 O2 N2 Air

rubber

thermoplastic elastomer

PET

PTFE

PP

silicone

48-72 30-300 M94-101 (Rockwell) 65-90 0.83-1.1 2.4-3.1 2.2-3.5 0.38-0.43 212-265 60-80 0.54-0.9 0.23-0.45 0.04

21-34 200-400 >100 (D50-65) 30-60 0.11-0.24 0.6 0.3-0.7 0.44-0.47

5-9 300-1000 20-70

327 26.5 25.4 7.9

31-41 100-600 R80-102 (Rockwell) 38-55 0.4-0.6 1.2-1.7 1-1.6 0.4-0.45 160-175 -20 23.8 7.9 1.9

PU

13498 2250 1975

120-160 19.4-194 4.0-13.7 1.0-6.0

2.3-4.6

1.9

0.43-1.7

4.9-9.0

2.6-9.8

PIB-PS

2-58 400-1000 13-94

3-24 250-1800 20-90

33-50 0.0008-0.003 17-43 0.0018-0.009 0.47-0.49

0.0007-0.7

-73

2.5b 4.2b

a m2/s Pa to cm3 mm/m2 day atm multiply by 8.75 × 1018, g mm/m2 day to g/m s multiply by 2.89 × 10-8. b Permeability data of PIB-PS are based on butyl rubber.

Scheme 1

implant shell application. Therefore, this new material will be discussed in more detail, in comparison with wellestablished biomaterials used in soft tissue replacement/ reconstruction, from the material science point of view. This point of view seems to be necessary as even recent surveys (most likely not written or reviewed by polymer scientists) contain major errors. Table 6 shows selected properties of these polymers, which will be discussed in more detail.55-57 Polyesters. Polyesters such as poly(ethylene terephthalate) (PET) are frequently found in medical applications, due to their unique chemical and physical properties. PET is a thermoplastic polymer made by polycondensation of ethylene glycol with either terephthalic acid or dimethyl terephthalate, as shown in Scheme 1.58 PET was first commercialized by DuPont in 1930 as Dacron. The physical properties of PET are largely determined by the degree of crystallinity. It is a strong material with a tensile strength of 170-180 MPa and a tensile modulus of about 14 GPa in oriented form.55 Vascular grafts made of Dacron are fabricated in woven or knitted textile/ mesh form, which is crimped to enhance kink resistance. Dacron was introduced by Debakey as a fabric graft in 1958 and has led vascular reconstructions for all of the major large vessels.59 It has very good material properties as shown in Table 6, but with a hardness of Shore A > 100 (Rockwell M94-101), it can hardly be considered elastomeric. Woven or knitted structures yield increased flexibility, and the pore size and distribution may be controlled by varying the density of the weave or knit which impacts the amount of tissue in-growth and adaptability once implanted. Unfortunately, the approximate life span of polyester grafts is only two years, with additional surgeries required to replace failed grafts.60,61 Dacron arterial prosthesis has been associated with

Scheme 2

early and late graft failures due to complications involving hemorrhaging, false aneurisms, and especially graft dilatation, which affects graft compliance.59,61 An increase in length and diameter of the grafts results in flattening of crimps created by the weave/knit compositions as the graft pressurizes in vivo, and this contributes to immediate dilatation. Finally, elongation and degradation causes permanent deformation due to fatigue of the material, resulting in further dilatation. These factors are only some of the physical and chemical mechanisms involved in the degradation processes related to polyester grafts that are inherently susceptible to degradation in the body. Modification of the chemical structure of polyesters (poly(ester-urethane)s, poly(esterether-ester-ester)s, and poly(ester-ether)s) led to improvements for biomedical applications.61 Biocompatible soft polyesters containing dimer fatty acids and/or poly(dimethysiloxane), with hardness comparable to medical grade silicone filled with silica (Shore A 85), were reported recently.62 Fluoropolymers. The best known fluoropolymer is poly(tetrafluoroethylene) (PTFE), commonly known as Teflon. PTFE is made from tetrafluoroethylene under pressure with a peroxide catalyst in the presence of excess water for removal of heat, as shown in Scheme 2. Teflon is linear, free from any significant amount of branching. The highly compact structure leads to a molecule of great stiffness and results in a high crystalline melting point and thermal stability.


Biomedical Application of Commercial Polymers Scheme 3

The polymer is formed into shapes by sintering, giving a moderate stiffness with tensile modulus of elasticity of 0.5 GPa and tensile strength of 14 MPa. PTFE is a tough and flexible material with excellent resistance to heat and chemicals and is a good electric insulator. Teflon is highly hydrophobic, with undetectable water absorption. The water vapor transmission rate of PTFE is approximately half of that of PET. Implantable forms of the material used for prosthesis include “expanded” and textile versions. In 1963, Shinsaburo Oslinge of Sumitomo Industries in Japan discovered a process for expanding PTFE (e-PTFE) during extrusion.63 Textile technology for both knitted and woven structures is most often used together with expansion of PTFE for production of e-PTFE grafts. These are produced by extrusion and sintering into a porous walled tube, consisting of fibrils and nodules, with controllable pore size. e-PFTE has been used for the fabrication of vascular grafts, especially for small diameter arterial substitutes. The porosity can be used advantageously to promote tissue in-growth and the formation and retention of an endothelial layer, but at the same time unwanted substances can also adhere to the surface. e-PTFE-based grafts, often with special external reinforcement via supporting coils and wraps to limit bleeding, are used primary for medium-sized artificial grafts. Again, with Shore A hardness over 100, PTFE can hardly be considered to be a soft material. Compared to PET, e-PTFE is less thrombogenic and inflammatory once implanted. However, PET is commonly used over e-PTFE due to its ease of handling and suturability.64,65 Polypropylene. Polypropylene (PP) is an addition polymer of propylene, produced by Ziegler-Natta or metallocane catalysis. During polymerization as shown in Scheme 3, the CH3 groups characteristic of this olefin can be incorporated spatially into the macromolecule in different ways, resulting in isotactic, syndiotactic, and atactic polypropylene. Stereospecific catalysts can control the isotactic position of the methyl group where the CH3 groups are on the same side of the main chain. PP is a strong, crystalline, high modulus thermoplastic with a tensile strength of 400 MPa in the isotactic form, whereas atactic PP is an amorphous, somewhat rubbery material, considered to be of little value commercially. Commercial PPs are usually about 90-95% isotactic. The greater the amount of isotactic material, the greater the crystallinity, and hence the higher the softening point, stiffness, tensile strength, modulus, and hardness are. Although polypropylene and polyethylene are structurally similar, PP has a lower density and higher Tg and Tm. The chemical resistance of polypropylene is similar to highdensity polyethylene, but compared to polyethylene, it is more susceptible to oxidation, chemical degradation, and crosslinking by physical means such as UV and high-energy irradiation. Moreover, PP has an exceptionally long flex life and excellent environmental stress-cracking resistance; hence,

Scheme 4

Scheme 5

it had been tried for finger joint prostheses with an integrally molded hinge design.66 PP yarns have been woven into multifilament tubes and used for single-component vascular prosthesis. Chronic in vivo studies suggest that PP offers advantages over e-PTFE and Dacron in small diameter vascular grafts, because their biomechanical behavior can be modulated by varying their diameter and weaving conditions, enhancing tissue ingrowth.67,68 Silicones. Silicone is the generic description for synthetic polymers containing repeat Si-O bonds in their backbone and organic groups attached to the silicone atom via Si-C bonds. There are a variety of functional groups that can appear on silicone. Some of the important examples are shown in Scheme 4. Silicones developed by the Dow Corning company are based on dimethylsiloxane, which is polymerized by condensation polymerization, as shown in Scheme 5.47 Cyclic siloxanes can also be used to obtain silicone elastomers.48 Low molecular weight polymers have low viscosity and can be crosslinked to produce higher molecular weight, rubber-like materials. The silicone elastomers most commonly used for medical applications are the high consistency (HC) types with high-temperature vulcanization (HTV) as well as liquid injection molding (LIM) types with room-temperature vulcanization (RTV).47,48,57 The former is crosslinked with peroxides or platinum, and the latter is cured by platinum as shown in Scheme 6. The first biomedical applications for silicone rubber can be traced back to the 1950s, whereas industrial-scale manufacturing dates to 1965.69,70 Today, silicone rubbers are the most widely used polymers in medical applications because of the strong Si-O-Si (siloxane) backbone, which provides enhanced chemical inertness and exceptional flex-

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Scheme 7

Scheme 8

ibility. Little degradation is expected to occur on implantation, whereas most other elastomeric materials show some degradation when left in the body for long periods of time. The unique stability of silicone rubbers in the biological environment has been a driving force for their continued use. Silicone rubbers also have exceptional properties such as high tear strength and outstanding elasticity over a wide temperature range. However, the mechanical properties of silicones are worse than those of other crosslinked elastomers. Medical grade silicone rubber is usually filled with silica particles, which very often may have a more profound influence on mechanical properties and biocompatibility than the nature of the silicone polymer itself.71 Recently, there has been increasing interest in the possible adverse effects of silicones used in implantation. Problems in the past have been associated in particular with adsorption of oxidized lipids, which causes swelling and slight dimensional change.71-73 Thus, for dynamic functioning such as vascular grafts, silicone rubber alone is not suitable, but this was not expected to be critical for breast implants. Although information about the degradation characteristics of silicones in the human body is limited, the potential breakdown of silicone can be examined from both physiochemical and chemical standpoints. There has been some debate regarding the chemical stability of siloxane bonds in specific physiological environments. For example, it was suggested that within a macrophage free radicals are abound, which may catalyze the hydrolysis of siloxane. It is known that, under certain conditions, the siloxane bond can be hydrolyzed to form silanols,74 as shown in Scheme 7. This is a reversible reaction catalyzed by acids, bases, and amines as well. Moreover, immunological reactions to silicone can also develop that can be local, regional due to silicone migration, or systemic. Migration of silicone has been documented on numerous occasions in the literature.62 Systemic reactions, such as acute renal insufficiency and respiratory compromise, etc., have been reported following the introduction of silicone into the body. However, the mechanism underlying the systemic immunological reactions to silicone has not been thoroughly investigated. Legal advocacy have been more vocal about possible adverse effects than biomedicine and material science, which further underlines the necessity of more detailed and more interdisciplinary investigations. Polyurethanes. Polyurethanes are copolymers consisting of soft and hard segments connected by the urethane linkage

-CO(O)NH-. The rest of the polymer may contain a variety of other functional groups such as polyethers, polyesters, ureas, epoxides, silicones, aromatic or aliphatic hydrocarbon groups, and polyolefins. Scheme 8 demonstrates the simplified version of polyurethane formation.24,42 The soft segments are usually derived from a macromonomer (macroglycol), ranging from several hundred to several thousand Daltons in molecular weight. The macroglycol imparts softness or stretchiness to the polyurethane and contributes to the relative hydrophilicity, lubricity, and biostability of the polyurethanes. The hard segments generally consist of the reaction product of a diisocyanate and a diol/diamine. The urethane linkage of the hard segment shows excellent secondary bonding that provides greatly enhanced mechanical properties. Based on hard to soft segment ratio, the polyurethane can be soft and rubbery or tough thermoplastic elastomer, or hard and strong engineering polymer. Thermoplastic elastomers (TPEs) are a class of materials that combine the material properties of rubbers and plastics. The soft elastomeric and hard amorphous or crystalline plastic phases separate at a nanoscale level, creating a physically crosslinked network structure shown in Figure 3. At room-temperature, TPEs behave like crosslinked rubbers but can be melted at higher temperatures as plastics and processed with conventional plastic processing methods. Polyurethanes with 60-85% soft segment content are considered to be segmented TPEs.24 Polyurethane TPEs are characterized by high elongation at break, high flexibility, and low permanent deformation on static and dynamic loading. They are very versatile materials that have been used in the body for over 20 years. However, they are sensitive to strong acids, strong alkalis, aromatics, alcohols, hot water, hot moist air, and saturated steam. Thus, polyurethanes are subject to biodegradation.

Biomedical Application of Commercial Polymers

Figure 3. Schematical illustration of TPE physical crosslinking.

Scheme 8 demonstrates the in vitro hydrolysis of polyurethanes; the diamine forming in this reaction can be toxic.24 Unfortunately, today there are no TPE alternatives commercially available for biomedical application, which demonstrate as high tensile strength, good abrasion resistance, fatigue properties and lubricity, easy handling and good biocompatibility as polyurethanes. Polyurethanes can have either aromatic or aliphatic hard segments as well as internal hard segment urethane or urea linkages. Aliphatic polyurethanes showed significant changes in molecular weight after implantation under strain, with polyurethane-urea demonstrating the largest losses in tensile properties over a 12 week implantation period.75 In addition, the aliphatic formulations often soften significantly at or near body temperature and become sticky and “slump” at 80 °C. White et al.76 showed that aromatic polyurethanes had better flex fatigue life, significantly higher wet tensile strength, and better heat stability than aliphatic polyurethanes. Aromatic polyurethanes also demonstrate less water absorption, higher dimensional stability when wet, and better solvent resistance than aliphatic polyurethanes. The effect of soft segment chemistry on the biostability of polyurethanes has been investigated.77 Polyurethanes containing hydrogenated poly(butadiene) or aliphatic hydrocarbon soft segments showed higher stability in oxidative environments. The chemical structure of aliphatic hydrocarbon polyols plays an important role in their susceptibility toward oxidation. The susceptibility of hydrocarbon moieties of polyol soft segments to oxidation is on the order of -[CH2CH2]n- < [CH2CH(CH2CH3)]n CH2- < [(CH2)x-O]n-. Incorporating poly(dimethyl siloxane), poly(ethylene oxide), or sulfonated soft segments into polyurethane has also been investigated.77-80 Poly(dimethyl siloxane)-containing polyurethanes showed severe degradation under oxidative conditions in the presence of organic carboxylic salts but not in pure oxidative environments, whereas poly(ethylene oxide) degraded by hydrolysis and oxidative ways. Sulfonated polyurethanes with antioxidants performed well in short-term tests; however, they demonstrated cracking after 1 month without these stabilizers.80 The earliest attempt to introduce polyurethanes into biomedical application was in the late 1950s when Pangman described a composite polyurethane foam breast prosthesis.81 Many reports are published on the degradation of the polyester urethane component of breast implants.82-84 The large pore size of the breast implant’s coating is sufficient for cells to migrate into and release acids, oxidants and


enzymes, which can catalyze hydrolysis.85 Furthermore, once the ester linkage begins to hydrolyze, it produces acid groups which further increase the acidity surrounding the degrading polyurethane and may autocatalyze its destruction.86 Virtually all of the polyurethanes used for prolonged periods in the body have been the polyether urethane type, specifically aromatic polyether urethanes. Studying the degradation mechanisms of the polyester and polyether urethanes has led to the observation that the weak links in these materials are the ester and ether bonds 87-89 and that aromatic polyurethanes are more stable than aliphatic polyurethanes.78,90-92 Reformulating these polyurethanes without ester and ether bonds, or with fewer ether bonds, has resulted in a new generation of more biostable elastomeric polyurethanes. Pinchuk et al.93 reported that proprietary polyurethanes without ether or ester linkages did not demonstrate biodegradation. The chemical structure of the novel soft segment was a polycarbonate glycol.94-96 The most biodegradation-resistant polycarbonate urethane formulations with the best physical properties are achieved with chemistries similar to that of the aromatic Pellethane formulations where the polyether glycol intermediate is replaced with a suitable polycarbonate glycol. New Materials for Soft Tissue Replacement: Polyisobutylene-Based Thermoplastic Elastomers Polyisobutylene (PIB) and butyl rubber (a copolymer of isobutylene and a small amount of isoprene) have been used in the food industry and are FDA-approved for chewing gum application. Butyl rubber is also used for pharmaceutical stoppers and blood bags, due to its excellent barrier properties. It should be emphasized here again that butyl rubber has an order of magnitude lower permeability than any other crosslinked elastomer. This could also be useful to prevent or delay enzymatic degradation, together with the excellent chemical and oxidative stability of PIB. However, PIB-based biomaterials have only been introduced recently by the Kennedy school at the University of Akron. Poly(methylmethacrylate) bone cement was modified with PIB, increasing the fracture toughness of the material.97 Amphiphilic networks (hydrogels) developed by combining PIB and hydrophilic polymer segments were used for coating GorTex vascular grafts and showed good biocompatibility.98 These hydrogels were also used as membrane carriers for insulin-producing porcine platelet implants.99 The size of the holes in the hydrogel allows small nutrient molecules and growth factors necessary for insulin production to reach the platelet and the insulin produced to exit, while preventing large “attack” molecules (antibodies, immunoglobulins) from entering the holes to generate an immune response toward the foreign platelet. The implants restored normoglycemia in diabetic rats for extended periods.3 A more detailed review of potential PIB-based biomaterials will be published elsewhere.100 A very important class of polymers is block copolymers because they give us a means to combine materials with different properties that may be incompatible by simple mixing. Scheme 9 demonstrates the formation of PIB blocked with polystyrene (PS) by living carbocationic polymerization.

1150


Figure 4. Schematical illustration of PIB-PS macromolecular structures. Scheme 9

Puskas and Chen

Figure 5. Stress-strain curves of arborescent PIB-PS block polymers with various PS contents106 Table 7. Oxidative Stability of Various Polymersa 109 polymer

time to dissolve

tensile strength remaining

remarks

PET PU Silicone PP PTFE PIB-PS

Biomedical Application of Commercial Polymers and Novel

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