Polyurethane for biomedical applications: A review of recent developments

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Polyurethane for biomedical applications: A review of recent developments Wenshou Wang and Chun Wang, University of Minnesota, USA

Abstract: Polyurethane is a very important polymeric biomaterial, widely used in the preparation of implants and medical devices. This chapter highlights recent research developments in polyurethanes for biomedical applications, including biocompatibility and biostability evaluation, for drug-controlled release carriers, for cardiovascular implants and for medical supplies. We conclude with an outlook of the future of polyurethane. Key words: polyurethane; biocompatibility; medical devices; drug carrier.

3.1 Introduction Polyurethane is a kind of polymer that contains repeating urethane groups. With an enormous diversity of chemical compositions and properties, it has found wide applications in a number of technological areas in our daily life, such as

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automotive parts, footwear, furnishings, construction, coatings, etc. (Oertel, 1994). In the last few decades, biomaterials used for prosthesis and medical devices have seen a rapid development and, due to advances in tissue engineering, polyurethane, as one of the most important biomaterials, finds a niche in this field because of its widely variable mechanical properties and excellent biocompatibility. Polyurethane materials were first introduced in biomedical applications in the late 1950s. In 1958, Pangman described composite breast prostheses covered with a polyesterurethane foam. Later that year, Mandarino and Salvatore (1958) used a rigid polyurethane foam called OstamerTM for in situ bone fixation. Since then, polyurethane, as a biomaterial, has been widely used in medical devices, and a series of biomedical grade polyurethanes were designed and developed accordingly by Bayer MaterialsScience, Lubrizol, BASF, etc. Table 3.1 lists some of the medical grade polyurethanes available on the market, which have been manufactured by these companies. Based on their excellent mechanical properties and biocompatibility, polyurethanes have been widely used in the preparation of all kinds of medical devices, including wound dressings, artificial organs, vascular stents, and so on. Many scientists worldwide are working in this area, to broaden the applications of polyurethane or to optimize its properties to meet the requirements of specific applications, and many papers and research results are published every year. Some very good reviews and books on the biomedical application of polyurethane have been published, which will give readers a comprehensive understanding of the progress of polyurethane in medical devices (Griesser, 1991; Lamba et al., 1997; Zdrahala and Zdrahala, 1999; Vermette et al., 2001). There have been rapid advances in this area over the past ten years and we think it is necessary to summarize the

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Table 3.1

Some commercially available medical grade polyurethane on the market

Trade name

Appearance and Type

Hardness

Tensile Strength (psi)

Tear Manufacturers Strength (pli)

Texin®

Natural clear polyether

85A, 90A, 50D, 5000–7400 65D

500–1200

Bayer MaterialScience

Desmopan® DP 2590A Natural clear polyester

90A

6500

n/a

Bayer MaterialScience

Desmopan® DP 9370A Natural clear polyether

70A

3100

350

Bayer MaterialScience

Texin® 4210

Opaque white TPU/PC blend

70D

6000

900

Bayer MaterialScience

Tecoflex®

Polyether

72A–83D

Lubrizol

Carbothane®

Polycarbonate

73A–75D

Lubrizol

Pellethane®

Polyether

53D–90A

Lubrizol

Elastollan® SP 806

Polyether

87A

6500

342

BASF

ChronoFlex®

Polycarbonate

75A–75D

5500–8000

n/a

AdvanSource Biomaterials

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latest most important advancements in this area. In this chapter, the following topics will be covered: biocompatibility and biostability evaluation; applications in drug-controlled release; application in cardiovascular devices and in medical supplies.

3.2 Biocompatibility evaluation As a medical device or implant, the evaluation of biocompatibility is essential, and good biocompatibility is the minimum requirement for any medical device. Biocompatibility, by definition, is the extent to which a foreign, usually implanted, material elicits an immune or other response in a recipient and the ability to co-exist with living organisms without harming them (McGraw-Hill Concise Dictionary of Modern Medicine, 2002). The biocompatibility of polyurethane has been extensively investigated by researchers, which includes the evaluation of in vitro (cell compatibility, toxicity, mutagenicity, as well as additional in vitro tests for a specific application, which will guarantee the function and innocuousness of the device) and in vivo tests. Boretos and Lyman first claimed the blood biocompatibility of polyurethane elastomers (Boretos and Pierce, 1968; Lyman et al., 1971). Of course, blood biocompatibility is only a part of biocompatibility. Recently, Lehle et al. (2009) reported cell-type specific evaluation of biocompatibility of commercially available polyurethanes in more detail. In their study, human saphenous vein endothelial cells (HSVEC) and a mouse fibroblast cell line (L929) were cultivated with different commercially available polyurethane specimens. Tissue-cultured polystyrene (TCP) was used as a reference. The cytotoxic effect was evaluated by morphology, cell viability, cell growth kinetics and proliferation tests.

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Finally, they concluded that commercially available polyurethanes provided an unfavorable support for colonization of patient-derived HSVEC, which demanded a surface modification. As further new kinds of polyurethane are being developed, it is important to expand our understanding on the relationship between the polyurethane structure and its biocompatibility, which will help us to design the new polyurethane materials or modify the existing materials based on our needs. Lyman and Picha pointed out that the blood biocompatibility of polyurethane is related to surface morphology (Lyman et al., 1975; Picha and Gibbons, 1978). The relationship between blood response and hard/ soft segment concentrations was also confirmed with Biomer™12 and other studies, which showed that the hard segments of PUs were highly thrombogenic in platelet retention experiments (Lelah et al., 1985; Takahara et al., 1985). Qu prepared a biodegradable block poly (ester-urethane)s based on poly(3-hydroxybutyrateco-4hydroxybutyrate) copolymers, indicating that crystallinity degree, hydrophobicity, surface free energy and urethane linkage content play important roles in affecting the lactate dehydrogenase activity and hence the platelet adhesion (Qu et al., 2011). Based on these very important findings, a lot of modification work has been done to improve the biocompatibility of polyurethane. With L-lysine diisocyanate and L-lysine chain extender, Han et al.’s results showed that both L-929 cells and HUVECs attached well to and remained visible on these polyurethane scaffolds prepared through electrospining (Han et al., 2011). Incorporation of nanocomposites (i.e. clay, POSS and silver) into polyurethane is another way to improve biocompatibility (Guo et al., 2010; Hsu et al., 2010; Tseng et al., 2011). In our previous study, we found that by incorporation of POSS (polyhedral oligomeric silsesquioxanes) into polyurethane, the surface

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tension decreased and surface morphology changed, which supported the mouse stem cell differentiation and proliferation (Guo et al., 2010). Besides changing the compositions of polyurethane, plasma and ion beam techniques are also convenient and effective ways to improve the biocompatibility of polyurethane (Melnig et al., 2005). Although various techniques have been developed to improve the biocompatibility of polyurethane and very exciting in vitro results have been acquired, it is the in vivo measurements that predominate. Accordingly, many in vivo experiments were conducted on polyurethane based materials. Pereira et al. (2010) synthesized a photopolymerizable and injectable polyurethane for biomedical applications, by employing minimally invasive procedures, the obtained injectable polymer systems could be molded in situ before photo-polymerization with visible light. In vitro tests indicated that the synthesized polyurethanes are cytocompatible; and in vivo tests indicated good biocompatibility during a 4-week period. Santerre, at the University of Toronto, recently reported a polar/hydrophobic/ionic polyurethane (D-PHI) based on lysine diisocyanate (LDI) (McBane et al., 2011). By implanting porous D-PHI scaffolds into a subcutaneous rat model, the in vivo biocompatibility of this polyurethane was assessed with PLGA as a reference. As a result, the elastic D-PHI scaffold demonstrated good biocompatibility within the in vivo environment (cell infiltration and tissue matrix development). Ciardelli et al. (2011) prepared polyurethane guides used for peripheral nerve regeneration, and tested in vivo for the repair of 1.8 cm-long defects in rat sciatic nerves. The results showed that after 45–60 days from the operation, limbs with polyurethane nerve guides showed the same behavior as intact limbs (full recovery), and that the recovery process started to be evident from the 15th day. Based on the above

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examples, we can see that some polyurethanes present very good in vivo biocompatibility. As we already know from in vitro results that the surface plays a very important role in the bio-compatibility of polyurethanes, is it also true for in vivo? The answer is yes according to Khandwekar’s research, which demonstrated that surface entrapment technique could be used to modify/control the foreign body response on polyurethane surfaces (Khandwekar et al., 2010). Besides the basic research, a lot of polyurethane based medical devices were prepared and placed into animals for in vivo tests, which will be discussed below.

3.3 Biostability evaluation The biostability of polyurethane has been of concern for a long time, which is why most of the polyurethane medical devices were only used as short-term implants. It is widely accepted that polyester based polyurethane is not stable in water and oxygen, so most of the medical grade polyurethanes on the market are polyether based or polycarbonate based. However, experiments show that polyether or polycarbonate based polyurethane was also not stable in vivo, and pacemakers were withdrawn from the market as a result (Vermette, 2001). The biodegradation of polyurethane and polyurethane composites has been extensively studied (Christenso et al., 2004; Lyu et al., 2008). Many modifications were attempted to enhance the biostability of polyurethane accordingly. One way is to replace the soft segment of polyurethane with more stable materials, such as polyolefins, polysiloxane, etc. (Ward et al., 2006; Cozzens et al., 2010; Kang et al., 2010), based on the findings that the soft segment played an important role in biostability (Wiggins et al., 2004). With mixed poly-isobutylene (PIB)/poly (tetramethylene

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oxide) (PTMO) soft segments, Cozzens synthesized a new polyurethane and demonstrated that under accelerated conditions (20% H2O2 solution containing 0.1 M CoCl2 at 50°C to predict resistance to metal ion oxidative degradation in vivo), the PIB-based TPUs showed significant oxidative stability when compared to the commercial controls, Pellethane™ 2363-55D and 2363-80A (Cozzens et al., 2010). Surface modification is another method of improving the biostability of polyurethane (Ward et al., 2007); other methods, including incorporation of nanoparticles (Hsu and Chou, 2004; Chou et al., 2006) and antioxidants (Stachelek et al., 2010), were also reported. Of course, the biostability of polyurethane is only important for long-term biomedical devices, because the loss of mechanical properties and shape usually comes with biodegradation. But some biodegradable polyurethane scaffolds are specifically needed in tissue engineering, and study on the preparation of biodegradable polyurethanes with controlled degradation rates is extremely urgent (Feng et al. 2007; Guelcher, 2008), as they have found important applications in drug delivery, short-term implants, scaffolds, etc., which will be discussed in the following applications.

3.4 Polyurethane for drugcontrolled delivery There has been much study on drug-controlled release recently, because of its great advantages such as lower toxicity and less side effects, fewer injection times and so on, compared with the traditional administration of drugs (Urich et al., 1999; Freiberg and Zhu, 2004; Acharya and Park, 2006). Furthermore, the carrier that encapsulates the drugs could protect them from unnecessary damage or loss, or

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targeted release could be realized by a smart carrier. With the rapid development of materials preparation and technology, the delivered substances have been extended from small molecules drugs to therapeutic factors such as proteins and genes. Polyurethane, as an important biomaterial, also received much attention for being used as a drug-controlled delivery matrix. Polyurethane has been used to prepare all kinds of medical devices or scaffolds to be used in tissue engineering, because of its excellent mechanical properties and processing ability. The polyurethane stent or scaffold itself is the ideal drug carrier for the local delivery system and many studies have been conducted. Heparin-Deoxycholic acid (DOCA) conjugate was loaded into a polyurethane film by solvent casting (Moon et al., 2001), and it was found that when the heparin-DOCA loaded on the polyurethane films was above 7.5%, the released heparin-DOCA prevented the formation of a fibrin clot and platelet adhesion on the film surface (Figure 3.1). The effects of dexamethasone-loaded polyurethane implants (PU ACT (dexamethasone acetate) implants) on inflammatory angiogenesis in a murine sponge model were investigated by Moura et al. (2011), showing that the local drug delivery systems derived from polyurethane efficiently modulated the key components of inflammation, angiogenesis and fibrosis induced by sponge discs in an experimental animal model. Hafeman et al. (2010) designed injectable polyurethane scaffolds incorporating tobramycin by reactive liquid molding, and the released tobramycin remained biologically active against Staphylococcus aureus. The release behaviors of many other small molecular drugs were also studied, such as ibuprofen, 5-flurouracil, etc. (Ring et al., 2008; Basa et al., 2009; Harisha et al., 2010). Beyond these small molecular drugs, Li et al. (2009) incorporated rhBMP-2 (recombinant human bone morphogenetic protein)

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Figure 3.1

Formation of fibrin clot on the polyurethane surface: (a) 0 wt% heparin-DOCA, (b) 5.0 wt% heparin-DOCA, (c) 7.5 wt% heparin-DOCA, (d) 10.0 wt% heparin-DOC (Moon et al., 2001)

into polyurethane scaffolds implanted in rat femoral plug defects, and the research results clearly showed that at 4 weeks post-implantation, all rhBMP-2 treatment groups showed enhanced new bone formation relative to the scaffolds without rhBMP-2, an impressive improvement (Figure 3.2). Because of the wide reactivity of isocyanate, many drugs can be chemically bonded into the main chain of polyurethane besides physically blending. The release of drugs could be realized by the degradation of polyurethane. Kenawy et al. designed polyurethanes containing azo-linked polymeric prodrugs of 5-aminosaheylic acid (5-ASA) in the main chain (Kenawy et al., 2010), which showed that

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Polyurethane for biomedical applications

Figure 3.2

In vivo evaluation of the effects of PUR/rhBMP-2 scaffolds on new bone formation in a rat femoral plug model. Treatment groups included: PUR control (no rhBMP-2), PUR/rhBMP-2, PUR/ PLGA-L-rhBMP-2, and PUR/PLGA-S-rhBMP-2. The PUR cylinders (5 mm_ 3 mm) were implanted into rat femoral plug defects (A), and harvested for mCT imaging at weeks 2 (B) and 4 (C), respectively (Li et al., 2009)

drugs could be released by the hydrolysis of the urethane bond in the main chain. Ghosh and Mandal (2008) synthesized ibuprofen-based polyurethane, and based on the easy cleavages of ester linkages, ibuprofen could be released. Besides the regular polyurethane materials, many environmental sensitive polyurethanes were also developed

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as drug controlled release carriers, such as temperature sensitive polyurethane (Chen et al., 2011; Sun et al., 2011), pH sensitive polyurethane (Zhou et al., 2011) and pressure sensitive polyurethane (Chen et al., 2009), and so on. Apart from the controlled release of a specific drug from a polyurethane matrix, simultaneous drug release at different rates from biodegradable polyurethane foams were also reported (Sivak et al., 2009); the anti-cancer compounds DB-67 and doxorubicin were covalently incorporated into polyurethane foams and their release behaviors demonstrated that differential release of covalently bound drugs is possible from simple single-phase, degradable polyurethane foams. The incorporation of drugs has no significant effect on the mechanical properties and biological performance of polyurethane. Simmons et al’s results indicated that incorporation of 25 mg/g dexamethasone acetate (DexA) into siloxane-based polyurethane resulted in no significant difference in the biostability and biocompatibility after implantation in an ovine model for 6 months, compared with the pristine polyurethane implant (Simmons et al., 2008). Clearly, the controlled release behavior of drugs from polyurethane is determined by many factors. da Silva et al. (2009) studied the effect of the macromolecular architecture of biodegradable polyurethanes on the controlled delivery of ocular drugs, and found that the presence of poly(ethylene glycol), together with poly(caprolactone) as soft segment in biodegradable PU, was able to increase the rate of dexamethasone acetate release when compared to the rate of drug release from PU having only poly(caprolactone). Also the incorporation of ionic ligands accelerated the drug release from polyurethane (Sivak et al., 2010a). Besides the structure of polyurethane, the functionality of drugs

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also played an important role (Sivak et al., 2010b). By designing the molecular structure of polyurethane and drugs, tailored drug release can be realized. Biodegradable stent coatings based on hybrid polyurethanes were reported by Guo et al. (2009), and the results clearly show that the studied hybrid polyurethane family allows a drug release rate that is effectively manipulated through variation in polymer glass transition temperature, degradation rate, and thickness increment rate. Besides the controlled release of small molecular drugs and proteins, Stachele et al. (2004) first demonstrated that modified polyurethane film could be used as localized and efficient gene delivery systems. By the surface attachment of replication defective adenoviruses, using an anti-adenovirus antibody tethering mechanism, the modified polyurethane implant was studied as a sitespecific gene delivery matrix and thereby demonstrated the potential for intravascular devices that could also function as gene delivery platforms for therapeutic vectors. It was concluded, based on in vivo experimental results from sheep pulmonary valve leaflet replacement study (Figure 3.3), that site-specific intravascular delivery of adenoviral vectors for gene therapy can be achieved with polyurethane implants utilizing the antivector antibody tethering mechanism. Increasing attention was paid to polyurethane based drug-controlled release, as the release mechanism is very important for the prediction of release behavior and in helping us to design the new materials. Reddy et al. (2006) found that the release of drugs from those synthesized polyurethanes can be explained by the Fickian diffusion model. Furthermore, polyurethane copolymers were also investigated as drug carriers (Mathews and Narine, 2010;

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Figure 3.3

Heart valve leaflet gene delivery using polyurethane (PU) pulmonary replacement cusps with antibodytethered AdGFP (108 PFU); (a) The gross appearance (pulmonary outflow side shown) of an AdGFP– polyurethane pulmonary leaflet explant demonstrating a smooth blood contacting surface. (b) A non-modified polyurethane pulmonary leaflet replacement at explant (also shown from the pulmonary outflow side), demonstrating a smooth blood-contacting surface, comparable to (a). (c–f) GFP expression demonstrated by fluorescent microscopy (FITC) with DAPI counterstaining in representative explanted AdGFP– polyurethane pulmonary cusp replacement (as in (a)) with cell localized GFP (c) expression on the surface of the cusp (see arrows) ×400. (d) Absence of autofluorescence in a control polyurethane cusp explant per fluorescent microscopy (FITC) with DAPI counterstaining (as shown in (b), above) ×400.

Polyurethane for biomedical applications

Figure 3.3

Continued (e) GFP expression in myocardium adjacent to the AdGFP–polyurethane cusp implant (as shown in (c). FITC/DAPI ×400. (f) An absence of FITC autofluorescence in myocardium adjacent to control (no adenoviral vector) polyurethane cusp implants (Stachelek et al., 2004)

Schroeder et al., 2007) for specific applications. It was found that polyurethane-14-AMP-acrylates copolymer (DynamX (R)) showed good ethinylestradiol permeation and the drug transport was further increased with the incorporation of oleic acid as a penetration enhancer (Schroeder et al., 2007).

3.5 Polyurethane for cardiovascular applications A perfectly functioning cardiovascular system is critical for the human body. With the advance of tissue engineering, scientists are able to restore or partly restore the malfunctioned tissues under the help of medical devices. As a result, far more medical devices are needed to meet the requirement of advances in cardiovascular technology. Because of its excellent biocompatibility, blood compatibility and duration,

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polyurethane has been the top candidate in the preparation of all kinds of biomedical devices for cardiovascular applications, such as catheters, pacemaker leads insulation, vascular prostheses, heart valves, cardiac assist devices, etc. (Vermette et al., 2001). Many polyurethane based devices have been commercialized and used in all kinds of surgery, for example, artificial blood vessel (CorthaneTM), Vascugraft® prosthesis, and Pellethane™ 2363-80A insulation, because most polyurethanes are inexpensive and reliable for a short duration of usage. Studies have continued to improve the properties of polyurethane and to make the best use of it in the cardiovascular area (Solis-Correa et al., 2007). Ashton et al. (2011) reported the preparation of polymeric endo-aortic paving with polycaprolactone and polyurethane blends, by characterizing the mechanical, thermoforming and degradation properties of the blends, which may be useful in developing the next generation of endo-aortic therapy. Shape memory materials based on polyurethane are also proposed as cardiovascular implants, due to their selfexpanding ability (Hassan et al., 2009). As shape memory polymer blends of PCL and polyurethane were prepared and measured, results showed that this material supports the cell adhesion and proliferation, so it might be a potential material for implant stents. Artificial heart valves are another important application of polyurethanes in the cardiovascular area. There are several recently published articles reviewing polymeric heart valves and polyurethane heart valves (Xue and Greisler, 2003; Venkatraman et al., 2008; Ghanbari et al., 2009; Kutting et al., 2011). Biostability in the human body has been discussed regarding the polyurethane made device. Xie et al. (2010) reported recent study results of inserting five kinds of polyurethane vascular grafts into dogs. The polyurethane grafts displayed substantial degradation after a 6-month

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period, and caused tissue deposition in the degradation area, indicating that polyurethane-made vascular grafts without further modification are not ideal for long-term use. Silicon was introduced into the polyurethane, in order to address the biodegradation problems occurring in long-term implantation. Briganti et al. (2006) synthesized silicon based polyurethane materials and studied the biocompatibility of the resulted polyurethane. It was concluded that the incorporation of silicone in the investigated range did not introduce any toxicity into the materials; however, further studies related to biostability and mechanical properties of the new material are necessary. Please refer to Section 3.3 on Biostability evaluation, for more information on how to improve the long-term biostability of polyurethane. Bioactivity is another desired property for biomedical stents, which will eventually enhance cell-matrix interaction and its effects. Kidane et al. (2006) incorporated lauric acid modified RGD into polyurethane, which showed a significant increase in metabolism but had no adverse effect on the platelet adhesion and hemolysis when compared to unmodified polyurethane. De Nardo et al. (2007) modified the polyurethane surface with heparin and tested the bacterial colonization. MTT tests and SEM observations showed a decrease in colonization of the different strains on the heparinized polyurethane surfaces, confirming that preadsorbed heparin plays a role in mediating biomaterial surface/bacterial cells interactions. The cellular behaviors of human umbilical vascular endothelial cells, such as attachment, growth and proliferation, were significantly increased onto the PU films surface modified by microwaveinduced argon plasma treatment (Lim et al., 2008). Most small diameter cardiovascular bypass products are made from expanded polytetrafluoroethylene (ePTFE). Because of the low long-term patency compliance of ePTFE,

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the small diameter cardiovascular bypass prostheses cause thrombosis and intimal hyperplasia (Sarkar et al., 2009). Therefore, polyurethane is a better candidate than ePTFE in the view of its biocompatibility, but the manufacture of a bypass with such a small diameter is a great challenge. Alexander et al. from University College London developed an extrusion-phase-inversion technique to manufacture uniform-walled porous conduits with polyurethane, named UCL-NANO™ (Sarkar et al., 2009); the equipment used is shown in Figure 3.4. This automated vertical extrusionphase-inversion device can reproducibly fabricate uniformwalled small caliber conduits with polyurethane and the resulting elastic micro-porous grafts demonstrate favorable mechanical integrity for haemodynamic exposure and are currently undergoing in vivo evaluation of durability and healing properties.

Figure 3.4

(a) Extruder device showing vertical mechanical arm with mandrel attached, (b) polymer chamber with mandrel entering superiorly and polymer introduction channel laterally and (c) under surface of polymer chamber showing adaptors enabling control of exit aperture size (Sarket et al., 2009)

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The corrosion of metal stents implanted inside an artery can have several adverse effects, such as possible tissue reaction or toxic effect caused by the metal ions leaking from the stent (Halwani et al., 2010), loss of mechanical strength, and so on. Besides the stent itself, polyurethane was also used as the coating layer of metal stents, to prevent the metal stent from corrosion and increase the biocompatibility of the materials. Trigwell et al.’s (2006) research results indicate that a polyurethane film of less than 25 µm was found to be sufficient for corrosion resistance and flexibility, without producing any excess stress on the stent structure. Straining the film to 225% and plasma modification did not affect the mechanical and surface properties, but allowed for improved biocompatibility as determined by the critical surface tension, surface chemistry and roughness. Mazumder et al. (2003) reported that the corrosion rate decreased rapidly from 275 µm/year for an uncoated surface, down to less than 13 µm/year for a 30 gm thick polyurethane coating. Stainless steel (316L) and Nitinol both contain potentially toxic elements, and both are subject to stress corrosion. Minimization of corrosion can significantly reduce both tissue reaction and structural degradation. Overall, polyurethane has showed great advantages and potential in applications in the cardiovascular area; much research work has been done and the potential of polyurethane is still being explored. However, some issues still need attention, such as long-term biostability in the human body, etc.

3.6 Polyurethane for medical supplies Dressings have been playing an important role in the healing of all kinds of wounds. The main purpose of a wound dressing is to provide a moist environment and to encourage the establishment of the best conditions for natural healing

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(Choi et al., 2011). It absorbs the exudates usually composed of blood, serous fluids and proteinaceous liquids and accelerates healing. Some medicinal agents, such as pharmacological active agents, antibacterial agents, topical anesthetics, bacteriostatic agents and antifungal agents, are often incorporated into the dressings, dependent on the specific applications. Many wound dressings made of different materials have been studied and reported, such as hydrocolloid (Jeong et al., 2011), hydrogel (Zohdi et al., 2011), medicated dressings (Yu et al., 2006) and liquid bandages (Martin-Garcia et al., 2005). Klode et al. (2011) investigated all kinds of wound dressings, totaling 56 used nowadays, and presented a useful data source that can be used to develop a specific wound dressing. Polyurethane is often used in wound dressings, because of its good barrier properties and oxygen permeability. Some commercially available polyurethane based wound dressings, such as Opsite®, have demonstrated that they will accelerate the inflammatory and proliferation phases of dermal repair when used in full thickness injuries, and also enhance contraction, revascularization and earlier remodeling of the wound (Vermette et al., 2001). A considerable amount of work is being conducted to make polyurethane dressings more suitable to specific applications. Dornseifer et al. (2011) reported the modification of polyurethane dressings for splitthickness skin graft donor (STSG) sites by perforation, which permits controlled leakage into a secondary absorbent dressing. By comparison with Aquace (a hydrofiber wound dressing), the modified polyurethane dressing was significantly less painful until and during removal. Incorporating ibuprofen into polyurethane dressings and the application in the management of STSG donor sites were also studied (Cigna et al., 2009), and the results demonstrated that the Biatain (a kind of polyurethane based dressing)-Ibu dressing is a useful

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tool in the management of STSG donor sites, by providing an optimal environment for wound healing, due to its bioocclusive properties and by minimizing pain and discomfort. Ring et al. (2008) did further work on the Biatain-Ibu polyurethane dressing, by studying the in vivo microvascular response of murine cutaneous muscle to this dressing. With Biatain (without Ibu) as the control, this study showed that local release of small-dose ibuprofen from a polyurethane dressing does not decrease new blood vessel growth during the implantation time of 12 days. Finally, the microvascularization of the implant’s border zones in both groups was found to be comparatively undisturbed. Choi et al. (2011) developed a liquid bandage based on polyurethane dispersion and it worked well in rats compared with gauze. Figure 3.5 shows the healing process of rat wounds with polyurethane based liquid bandage and traditional gauze; clearly, the polyurethane dressing had a better effect. Varma et al. (2006) also assessed the efficacy of polyurethane foam dressing on debrided diabetic lower limb wounds, which showed a significant reduction in the time taken for wounds to heal when sterile, non-medicated polyurethane foam dressings were used, compared to conventional dressings. Humid environments may help wounds heal faster, but accumulation of exudates under the dressing can cause infection (Matsuda et al., 1993). Nanofibrous membrane (NFM) is a wound dressing that has high gas permeation and could protect the wound from infection and dehydration at the same time. Various kinds of nanofibrous membrane dressings have been developed and reported recently (Khil et al., 2003; Chen, 2010; Lakshman et al., 2010; Liu et al., 2010). Chen prepared a bioactive polyurethane nanofibrous dressing containing silver nanoparticles by electrospinning (Chen, 2010). After modification, the NFM’s antimicrobial

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Figure 3.5

Comparison of wound healing by (a) gauze and (b) liquid bandage (sample PD2) dressings (Choi et al., 2011)

Polyurethane for biomedical applications

activity improved to 100% inhibition of bacterial growth and the in vivo results also showed that this NFM was better than gauze and commercial collagen sponge wound dressings in the wound healing rate. Khil’s results also indicated that the nanofibrous polyurethane membrane showed controlled evaporative water loss, excellent oxygen permeability, and promoted fluid drainage due to the porosity and inherent property of polyurethane (Khil et al., 2003). Using male adult guinea pigs as the host, without any further modification, neither toxicity nor permeability to exogenous microorganisms was observed with the nanofibrous membrane. Histological examination also confirmed that the epithelialization rate was increased, as shown in Figure 3.6, and the exudate in the dermis was well controlled by covering the wound with the electrospun membrane. Besides the widely studied and used wound dressings, polyurethane is also a good candidate for the medical adhesive that finds applications in tissue engineering. Tissue adhesives are a valuable alternative to mechanical tissue fixation by sutures or staples, especially for the regions where damage and bleeding must be avoided. Biodegradable and biocompatible polyurethane based bio-adhesives have been studied and show promising results (Ferreira et al., 2007; Sternberg et al., 2010). Many polyurethane hydrogels were used to fabricate contact lenses (Novartis, 2005), indicating a bright future for polyurethane in the medical supplies area.

3.7 Future outlook The need for increasing numbers and types of biomedical devices will be a feature of our society as the population continues to increase and age. The future of synthetic

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Figure 3.6

Histological findings of wound (a) at 3rd day postwounding of the control group (HE stain ×200); (b) at 3rd day postwounding, polyurethane-membrane group (HE stain ×200); (c,d) at 6th day postwounding of the polyurethane-membrane treated group (HE stain × 200); (e) at 15th day postwounding of the control group (HE stain ×200); (f) at 15th day postwounding of the polyurethane-membrane treated group (HE stain ×200). (Khil et al., 2003)

polymeric biomaterials is bright because of their advantages as polymers, such as low cost, ease of processing, properties that are adjustable, etc. Polyurethanes, by virtue of their range of properties, will continue to play an important role among

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polymeric biomaterials. Despite the great success of polyurethanes in some areas however, the requirements of other applications are not fully met by existing polyurethanes, and new requirements will appear continually with the advance of implantology and biomedical devices. The biostability of polyurethane will continue to be one of the focuses for medical implants in the long term. However, biodegradable polyurethane will continue to receive significant attention because of its great potential as a drug controlled release carrier, wound dressing, scaffold or stent in soft tissue engineering, and much research still needs to be done on controlling the degradation rate, long-term biocompatibility and biological effects of degradation products, etc. Bioactive or functionalized polyurethanes are ideal for application in tissue engineering, and how to modify them properly and effectively at low cost will be a challenge.

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