**12. Polyurethane (PU)**

PU is a series of biomaterials that contains urethane radical and offers the greatest versatility in compositions and properties of any family of polymers. Especially, a few specific elasto‐ meric PU compositions have demonstrated a combination of toughness, durability, biocom‐ patibility and biostability for being used as implantable medical devices, which is not achieved by any other available materials [181]. Because urethane is available in a very broad hardness range (e.g. eraser-soft to bowling-ball-hard), it allows the engineer to replace rubber, plastic and metal with the ultimate goals in abrasion resistance and physical proper‐ ties. During the last half century, PUs have become and remained the most valuable im‐ plantable elastomers for uses requiring toughness, durability, biocompatibility and biostability [182]. With their inherently stable in the body environment, some of the PUs have been widely used in medical applications such as synthetic heart valves, vascular grafts, and pacemaker electrodes. However, these usages of PUs have been limited by three major complications: calcification, thrombosis, and chemical degradation [183].

ymers (e.g. Pus) [30]. PCL has been approved by the FDA for specific applications, such as a drug delivery devices, sutures, or adhesion barriers. It has been widely used as a scaffold material for tissue engineering with mismatched mechanical properties and slow degrada‐

Various categories of drugs have been encapsulated in PCL, in microsphere, nanosphere or bulk states, for targeted drug delivery and for controlled drug release [174-176]. For exam‐ ple, a PCL scaffold modified by grafting nerve growth factor (NGF) and Tirofiban (TF) has been used as nerve conduits to promote the regeneration of sciatic nerves [177]. Low molec‐ ular weight PCL pieces can be ingested and digested ultimately by phagocyte and giant cell

**Figure 8.** Micrographs illustrating extracellular degradation of biomaterials by macrophage fused multinuclear giant cells. (A) A foreign body giant cell (FBGC) engulfed a fragment of poly(epsilon-caprolactone), PCL polymer *in vivo*. Nu, nuclei of FBGC. The PCL polymer was dissolved during sample preparation. Transmission electron microscopy (TEM), bar = 2 μm. (B) *In situ* cross-section of the interface between a multinuclear giant cell (MnGC) and PLGA film. Note the pseudopodia of the MnGC penetrated deep inside the surface of PLGA film and formed sealed compartments. PLGA polymers are eroded within the compartments. Focused ion beam (FIB) microscopy, bar = 5 μm. (C) *In situ* cross-sec‐ tion of the interface between an osteoclasts-like cell (OC) and calcium phosphate cement. Note the typical ruffled board of OC and vesicles (V) secreting from OC to the sealed extracellular space. FIB microscopy, bar = 2 μm [162].

PU is a series of biomaterials that contains urethane radical and offers the greatest versatility in compositions and properties of any family of polymers. Especially, a few specific elasto‐ meric PU compositions have demonstrated a combination of toughness, durability, biocom‐ patibility and biostability for being used as implantable medical devices, which is not

tion rate [171,172]. In rats the *in vivo* degradation of PCL is about 3 years [173].

without any cumulate vice-products (Figure 8) [178-180].

130 Advances in Biomaterials Science and Biomedical Applications

**12. Polyurethane (PU)**

In the 1970s and 1980s as the PUs became recognized as the blood contacting material and were used in a wide range of cardiovascular devices in long-term implants, they fell under scrutiny with the failure of pacemaker leads and breast implant coatings in the late 1980s. According to the manufacturer's report, high voltage coil fracture and PU defects were the predominant causes of lead failure [184,185]. During the next decade PUs had been exten‐ sively researched for their relative sensitivity to biodegradation and the desire to further un‐ derstand the biological mechanisms for *in vivo* implantation [186,187]. Some investors have seeded autologous sheep blood outgrowth endothelial cells (BOECs) on a cholesterol (Chol) modified PU (PU-Chol) heart valve leaflet to result in an intact, shear-resistant endothelium that would promote resistance to thrombosis [188]. Because of the complex behavior of im‐ plantable PUs in the body environment, special attention to the choice of the constituted components must be paid for designing and manufacturing the PU-containing devices. Sub‐ sequent treatment during sterilization, storage, implantation, *in vivo* operation and explanta‐ tion also determine the performance and provide the means for assessing the efficacy of the PUs implants [189].

The most prominent disadvantages of PUs being used as artificial heart valves include min‐ eralization, environmental stress-cracking and oxidation. While the mechanisms of these forms of degradation are not fully understood, an awareness of their causes and effects that leads to all of the long-term functionality is required for the sophisticated PU-based devices of today and tomorrow [190-191]. Over the last half century, extremely efforts have been paid in the biomedical research field to improve the biocompatibilities and biodurability of the PU implants, but only resulted in very little clinical effects [192-194].

In the later l990s a number of new bioresorbable materials with all the versatility of PUs in terms of physical properties and biocompatibility have been yielded. AorTech Biomaterials was set up in 1997 to commercialise a range of medical grade PUs developed by the Austral‐ ian research group (Commonwealth Scientific and Industrial Research Organization, CSIRO). The company estimates that the worldwide market for surgical heart valve prod‐ ucts is worth more than \$1bn (€705 m) and to be growing at a rate of 8% a year. Meanwhile, the market for catheter-delivered heart valves is worth around \$200 m (€141 mm) [195]. In the authors' own group in Tsinghua University, China, a novel PU made of PCL, PEG, and 1,6-hexmethyldiisocyanate has been synthesized. The hydrolytic degradation property of the PU can be highly tuned by changing the composition and structure of copolymers, such as PEG and PCL. When this kind of PU was used as a small-caliber (1.2 mm inner diameter) vein and nerve repair grafts it demonstrated excellent antithrombogenicity and superior bio‐ compatibility (Figure 9) [196,197].

**Figure 9.** An implantable small-diameter nerve and blood vessel repair PU conduit. (A) PU conduits with different in‐ ner diameters. (B) The PU conduit was connected to the vein of a rabbit. (C) The vein defect repair processes with a very thin layer of fibrin-platelet deposition. (D) The nerve repair processes in rabbits with growing myelinated axons. (E) The PU conduits degraded gradually *in vivo* in 12 weeks [196, 197].

### **13. Polytetrafluoroethylene (PTFE)**

Polytetrafluoroethylene (PTFE), Discovered in 1938 by Roy J. Plunkett, is a synthetic highmolecular-weight compound consisting wholly of carbon and fluorine with numerous ap‐ plications [198]. The best known brand name of PTFE is Teflon made by DuPont Co. It is insoluble in all normally used organic solvents, not biodegradable *in vivo* and can suffer high temperatures as 260 ℃ permanently. Clinically, PTFE has been widely used as a large blood vessels repair materials.

A 5 year research using PTFE-Gore-Tex grafts mainly for superficial femoral occlusion has been conducted. The majority of the grafts were inserted in an elderly poor risk group of patients with critical ischaemia of the lower limb. The overall cumulative patency at 2 years was 29% falling to 18% at 5 years. Perioperative angiographic indicated that inflammatory reaction is the only risk factor significantly affecting the cumulative graft patency. The pres‐ ence of diabetes was found to have a significant detrimental effect on limb salvage [199]. A permanently implantable left ventricular assist device, made of Dacron velour, Teflon felt, and Teflon-coated polyester fiber sutures, has been tested in chronic animal experiments. *In vivo* experiments demonstrated that all components elicited mild to moderate inflammatory reactions. Tissue responses to PTFE are rather passivated. Hematocele occurred only when the components were implanted in the aorta with direct blood contact and exposed to arteri‐ al blood pressures [200]. An 8 cm long PTFE prosthesis was implanted into defects of the abdominal aorta of dogs, and the following changes were found: the blood flow through the vascular prosthesis induced a shortening of the blood clotting time and a slight increase in the prothrombin consumption. It has a favourable effect of the sealing of pores in the pros‐ thesis and covering its internal surface with a fibrin membrane [201].

### **14. Silicone**

vein and nerve repair grafts it demonstrated excellent antithrombogenicity and superior bio‐

**Figure 9.** An implantable small-diameter nerve and blood vessel repair PU conduit. (A) PU conduits with different in‐ ner diameters. (B) The PU conduit was connected to the vein of a rabbit. (C) The vein defect repair processes with a very thin layer of fibrin-platelet deposition. (D) The nerve repair processes in rabbits with growing myelinated axons.

Polytetrafluoroethylene (PTFE), Discovered in 1938 by Roy J. Plunkett, is a synthetic highmolecular-weight compound consisting wholly of carbon and fluorine with numerous ap‐ plications [198]. The best known brand name of PTFE is Teflon made by DuPont Co. It is insoluble in all normally used organic solvents, not biodegradable *in vivo* and can suffer high temperatures as 260 ℃ permanently. Clinically, PTFE has been widely used as a large

(E) The PU conduits degraded gradually *in vivo* in 12 weeks [196, 197].

**13. Polytetrafluoroethylene (PTFE)**

blood vessels repair materials.

compatibility (Figure 9) [196,197].

132 Advances in Biomaterials Science and Biomedical Applications

Silicon is a metal in the same column as carbon in the periodic table with the symbol Si and atomic number 14 [30]. It is the most abundant element on earth and does not oc‐ cur naturally in its pure metallic state. Dimethylsiloxane is the building block for most medical-grade silicone products, including breast implants. This FDA Grade Silicone sheeting is commonly used in applications where food or consumables are present. For more than 20 years silicone miami breast implants have gone through a lot of changes since their first uses. After the mid-1980s many reports concerns the rupture rate of the thinner-shell products, the risk of subsequent breast cancer, and the connective-tissue diseases or symptoms in women with silicone gel-filled breast implants appeared. In the United States a moratorium (in place since 1992) on the use of these prostheses has been maintained by the pressure of overwhelming litigation. At the same time, Australian au‐ thorities also restricted the availability of silicone breast implants. Huge damages award‐ ed by United States courts forced Dow Corning, manufacturer of a large percentage of breast prostheses, to file for Chapter 11 bankruptcy in May 1995 [202].

As with any implantable medical devices or drugs, the risk of possible adverse effects must always be weighed against the ability to provide benefits. A great deal of safety research combined with more than 40 years of clinical experience has proven the efficacy and relative safety of the silicone gel breast implants. A rough estimate of implant shell rupture rate is ~10% at 10 years with both biocompatibility and biodurability problems [203]. A fibrocon‐ nective tissue capsule was found around all the samples [204]. The capsule formed around implanted mammary prosthesis is highly differentiated and organized, consisting of three layers: interface layer in three variations, intermediate fibrous layer of dense rough collagen fibers and light elongated cells with oval nucleus between them and adventitious layer. Be‐ tween the fibers of the interface and the middle strata intra- and extracellular silicone drop‐ lets and bulks were observed, representing the location where further pathological processes can take place [205]. It is said by Dr. Sidney Wolfe, director of Public Citizen's Health Research Group, in a statement that: "Public Citizen continues to oppose the FDA's 2006 decision to return silicone breast implants to the market for cosmetic use in women for augmentation. The agency's newer information about the risk of implant-associated lym‐ phoma and the previously known risks are serious enough to warrant advising women against having these implanted."

On March 9, 2012 a new silicone breast implant, which joins the two other silicone breast implants on the market - one made by Allergan and the other by Mentor, was approved by the FDA of the United States of America. Recommended monitoring after initially silicone breast implantation is 3 years and then every two years thereafter. In a review Roach and coworkers concern the importance of length and time on physicochemical interactions be‐ tween living tissue and biomaterials that occur on implantation. The review provides de‐ tailed information on material host interactions, dynamic material/cell surface states, surface chemistry and topological roles during the first stage of implant integration, namely protein adsorption. Generally, after the first contact of material with host tissue a state of flux due to protein adsorption, cell adhesion and physical and chemical alteration of the implanted ma‐ terial is followed (Figure 10) [206]. This model can answer many questions concerning the conformational form and bound proteins and therefore has instruction meanings in new im‐ plantable biomaterial design field.

**Figure 10.** Schematic of protein–surface interactions: Chemistry—adsorption onto biotinylated stripes which appear white, whilst adsorption is hindered on square oligoethylene-glycol regions, the white box shows an intentionally bleached area Topography—albumin adsorption onto hydrophilic silica spheres of varying dimensions as a model of surface curvature [206].

Beside the breast implants a silicon-silk transistor about one millimeter long and 250 nano‐ meters was created. So far the technique has been tested on mice with no adverse effects. Electrical, bending, water dissolution, and animal toxicity studies suggest that this approach might provide many opportunities for future biomedical devices and clinical applications [207]. A silicone catheter attached to a 2-5 x 1-3 cm stainless steel chamber with a self sealing injection port had been intravenously for antimicrobial chemotherapy. Peripheral venous ac‐ cess had become unsatisfactory in all of patients, and six had required central venous cathe‐ terisation [208]. More recently, a silicon-based neural probe with microfluidic channels was developed [209].

Origins of controlled release of implantable drug delivery dates back to 1964 when silicone implants were used to prolong a drug effect. Over 40 years, the progress to a safe, effective and acceptable implant system has been slow. The critical factors in implant research which need to be addressed include: erodibility, reproducibility, lack of irritation and carcinogenic‐ ity, lack of dose dumping, duration and pulses. While it is possible to surgically implant and remove drug-containing devices or polymeric matrices, the requirement for such interven‐ tion could have a significant negative impact on the acceptability of a product candidate. In recent years, two implant systems have been approved for human use; (a) a silicone-based device (NorplantR), and (b) a system based on lactide/glycolide copolymers to release a lutei‐ nizing hormone - releasing hormone (LHRH) agonist for treatment of male reproductive tract tumours. This drug delivery approach is very appealing for a number of classes of drugs, particularly those that cannot be given via the oral route, and drug candidates whose therapeutic index is relatively large [210].
