**Degradation of Polyurethanes for Cardiovascular Applications**

Juan V. Cauich-Rodríguez, Lerma H. Chan-Chan, Fernando Hernandez-Sánchez and José M. Cervantes-Uc

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/53681

### **1. Introduction**

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Polyurethanes are a family of polymers used in a variety of biomedical applications but mainly in the cardiovascular field due to their good physicochemical and mechanical properties in addition to a good biocompatibility. Traditionally, segmented polyur‐ ethanes (SPUs), have been used in cardiovascular applications (Kuan et al., 2011) as per‐ manent devices such as pacemaker leads and ventricular assisting devices; however, due to their great chemistry versatility, SPUs can also be tailored to render biodegradable systems for the tissue engineering of vascular grafts and heart valves. Therefore many research work have been focused on varying the chemical composition to enhance bio‐ stability or more recently to control the biodegradability of polyurethanes depending on specific applications in the cardiovascular field (Bernacca et al., 2002; Stachelek et al., 2006; Thomas et al., 2009; Wang et al., 2009; Hong et al., 2010; Arjun et al., 2012; Styan et al., 2012). In this way, polyurethanes for biomedical applications can be classified in two main types, according their relative stability in the human body as either biostables or biodegradables. In this chapter, general aspects of polyurethanes chemistry are pre‐ sented first and then, the various types of degradations that can affect these polymers both *in vivo* and simulated *in vitro* conditions. Emphasis is also made on the mecha‐ nism of degradation under various conditions and the techniques used for following the changes in their properties.

© 2013 Cauich-Rodríguez et al.; licensee InTech. This is an open access article distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited. © 2013 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

#### **2. Chemistry of polyurethanes**

#### **2.1. Synthesis of polyurethanes**

Polyurethanes (PU's) properties depend both on the method of preparation and the mono‐ mers used. In general, PU's can be prepared in one shot process or more commonly by a two step method, especially for the case of segmented polyurethanes (SPU's). These materials are thermoplastic block copolymers of the (AB)n type consisting of alternating sections of hard segments, composed of a diisocyanate and a low molecular weight diol chain extender, and soft segments, generally composed of various types of polyols, also called macrodiols. In the two steps method for SPU synthesis, a prepolymer is first obtained and then chain extended as illustrated in Figure 1. In the first step, an excess of the diisocyanate reacts with the soft segment polyol to form the prepolymer. Here, the characteristic urethane linkages are formed through the reaction between the isocyanate groups and the hydroxyl-terminat‐ ed end groups of the polyol. In the second step, the low molecular weight chain extender is used to link the prepolymer segments yielding a high molecular weight polymer. During this stage, additional urethane functional groups are formed when using a diol chain ex‐ tender whereas ureas are produced when a diamine is used.

**Figure 1.** Standard two-step reaction to prepare segmented poly (urethane)s and poly(urethane-urea)s

The properties of polyurethanes as those shown in Figure 1 depend on the various types of monomers that are used during their manufacturing (see Table 1). Historically, biostable polyurethanes were first developed by using polyether type of polyol and dif‐ ferent aromatic diisocyanates. Further developments in this area were focused on the substitution of the polyether macrodiol by novel hydrocarbon, polycarbonate or silox‐ ane macrodiols (Gunatillake et al., 2003) or a combination of these which in general are responsible for the flexibility of the SPUs (Król, 2007). In addition to the polyol chemi‐ cal composition, their molecular weight and concentration have an important effect in the polyurethane behavior. They can be incorporated in various concentrations but up to 50-75% of the polyol is common.

**2. Chemistry of polyurethanes**

52 Advances in Biomaterials Science and Biomedical Applications

tender whereas ureas are produced when a diamine is used.

**Figure 1.** Standard two-step reaction to prepare segmented poly (urethane)s and poly(urethane-urea)s

Polyurethanes (PU's) properties depend both on the method of preparation and the mono‐ mers used. In general, PU's can be prepared in one shot process or more commonly by a two step method, especially for the case of segmented polyurethanes (SPU's). These materials are thermoplastic block copolymers of the (AB)n type consisting of alternating sections of hard segments, composed of a diisocyanate and a low molecular weight diol chain extender, and soft segments, generally composed of various types of polyols, also called macrodiols. In the two steps method for SPU synthesis, a prepolymer is first obtained and then chain extended as illustrated in Figure 1. In the first step, an excess of the diisocyanate reacts with the soft segment polyol to form the prepolymer. Here, the characteristic urethane linkages are formed through the reaction between the isocyanate groups and the hydroxyl-terminat‐ ed end groups of the polyol. In the second step, the low molecular weight chain extender is used to link the prepolymer segments yielding a high molecular weight polymer. During this stage, additional urethane functional groups are formed when using a diol chain ex‐

**2.1. Synthesis of polyurethanes**

Commercial and experimental polyurethanes have been synthesized by the combination of the aforementioned monomers. Poly(tetramethylene oxide) (PTMO) is the most com‐ mon polyether in conventional medical formulations (Silvestri et al., 2011). Thus, for ex‐ ample, the Pellethane® 2363 80A and ElasthaneTM 80A are poly(ether-urethane)s obtained by the reaction of PTMO, MDI and BD monomers; Tecoflex® by Thermedics is also a poly(ether-urethane) synthesised by the reaction of PTMO, HMDI and BD monomers while Biomer® is a poly(ether-urea-urethane) synthesized from PTMO, MDI and ethylenediamine. Bionate®, Myo LynkTM and Chronoflex are polyurethanes pre‐ pared with polycarbonate diol. These commercial polyurethanes are typical examples of biostables polymers.

The use of vegetable raw materials containing hydroxyl groups such as starch, castor oil, vegetable oil, natural rubber, cellulose, etc, makes possible to obtain biodegradable polyur‐ ethanes (Krol, 2007; Aranguren et al. 2012). However, ester polyol commonly used to syn‐ thesize biodegradable polyurethanes are polycaprolactone, polylactic acid and adipate polyols. Polyethyleneglycol is a polyether which has been copolymerized with poly lactic acid and/or polycaprolactone because its higher hydrofilicity can accelerate the biodegrada‐ tion when this is required (Guan et al., 2005b; Wang et al., 2011b).

The most frequently used diisocyanates in the synthesis of biodegradable polyurethanes for biomedical applications are aliphatic or cycloaliphatic as MDI and TDI which can release carcinogenic and mutagenic aromatic diamines (Heijkants *et. al.*, 2005). Aliphatic diisocya‐ nates are less reactive than the aromatic counterparts but have a greater resistance to hydrol‐ ysis compared to aromatic diisocyanates, although this resistance frequently results in lower mechanical properties (Gogolewski, 1989).

In general, there are two types of compounds that are generally used as chain extenders, di‐ ols or diamines, which can either be aliphatic or aromatic, depending on the required prop‐ erties in the synthesized polyurethanes. New chain extenders, including amino acids have been also used during polyurethane synthesis as isocyanates can react vigorously with amine, alcohol, and carboxylic acids (Thomson, 2005). These novel chain extenders have been used to synthesize biodegradable polyurethanes (Skarja et al., 1998; Marcos-Fernández et al., 2006; Sarkar et al., 2007).


**Table 1.** Common monomers used in the synthesis of biostable and biodegradable polyurethanes

During polyurethane synthesis, several side reactions may occur leading to branching, crosslinking, or changes in the stoichiometry of reactants. For example, undesirable branch‐ ing and crosslinking may occur at elevated temperatures between isocyanates and urethanes (Allophanate formation) and isocyanates and ureas (Biuret reactions). Furthermore, the presence of water causes isocyanate groups to form unstable carbamic acids, which subse‐ quently decompose to amines with the liberation of CO2 gas (see Figure 2). These newly formed amines react with isocyanates to form ureas, thus changing reactant stoichiometry and leading to lower molecular weight polymers. Additives are sometime used for improv‐ ing specific properties of the polyurethane, for example Vitamin E and Santowhite®, two hindered phenolic antioxidants, prevents oxidative chain scission and crosslinking of poly(ether urethane) by capturing oxygen radicals (Schubert et al., 1997; Christenson et al., 2006). Di-tert-butylphenol and bisphosphonates have been incorporated to promote bro‐ moalkylation of urethane nitrogens in prepolimerized polyurethanes to inhibit the oxidation or calcification (Alferiev et al., 2001; Stachelek et al., 2007). Other compounds as fluorocar‐ bon or polydimethylsiloxane end groups have been attached to the surface of polyurethanes in order to enhance their biostability and hemocompatibility (Ward et al., 2007; Xie et al., 2009; Jiang et al., 2012).

In general, monomer type and stoichiometry, type and concentration of catalyzer, tempera‐ ture and moisture, and the use of additives are important in parameters for controlling the properties of these polymers.

**Figure 2.** Secondary reactions involved during polyurethane synthesis

#### **3.** *In vivo* **degradation**

**Monomeric**

Polyol (macrodiols)

Diisocyanate

Chain Extender Polyethers

54 Advances in Biomaterials Science and Biomedical Applications

Others

Aromatic

Aliphatic

Diols

**component Type Chemical compound Type of polyurethanes**

Poly(tetramethylene oxide)

Methylene diphenyl diisocyanate (MDI)

isocyanate) (HMDI)

1,6- hexamethylene diisocyanate (HDI)

Ethylene glycol (EG)

Aromatic diamines

**Table 1.** Common monomers used in the synthesis of biostable and biodegradable polyurethanes

During polyurethane synthesis, several side reactions may occur leading to branching, crosslinking, or changes in the stoichiometry of reactants. For example, undesirable branch‐ ing and crosslinking may occur at elevated temperatures between isocyanates and urethanes (Allophanate formation) and isocyanates and ureas (Biuret reactions). Furthermore, the

Diethylenglycol 1,4-butanediol (BD) 1,6-hexanediol (HD)

(LDI)

Diamines Aliphatic diamines

Others Amino acids

4,4′-methylene bis(cyclohexyl

L-lysine ethyl ester diisocyanate

Poly(ethylene oxide) (PEO) Biostables, Biodegradables (Sarkar et al., 2009; Lu et al., 2012) Poly(propylen oxide) (PPO) Biostables, Biodegradables (Francolini et al., 2011)

(PTMO) Biostables (Silvestri et al., 2011; Jiang et al., 2012) Poly(caprolactone) (PCL) Biodegradables (Sarkar et al., 2009; Lu et al., 2012)

Poly hydroxyalkanoates (PHA) Biodegradables (Li et al., 2009; Liu et al., 2009) Poly(ethylene adipate) (PEA) Biodegradables (Macocinschi et al., 2009)

Poly(dimethylsiloxane) (PDMS) Biostables (Park et al., 1999; Madhavan et al., 2006)

2,4-toluene diisocyanate (TDI) Biostables (Labow et al., 1996; Basak et al., 2012)

Chan-Chan et al., 2010)

Baudis et al., 2012) 1,4-butane diisocyanate (BDI) Biostables, Biodegradables (Heijkants et al., 2005; Hong et al., 2010) Isophorone diisocyanate (IPDI) Biostables, Biodegradables (Jiang et al., 2007; Ding et al., 2012; He et al., 2012)

Biostables (Gunatillake et al., 1992; Styan et al.,

Biostables, Biodegradables (Thomas et al., 2009;

Biostables, Biodegradables (Wang et al., 2011a;

Biodegradable (Abraham et al., 2006; Guelcher et al., 2008; Han et al., 2009; Wang et al., 2011b)

Biodegradables (Kartvelishvili et al., 1997; Skarja et al., 1998; Marcos-Fernández et al., 2006; Sarkar et

Biostables, Biodegradables (Król, 2007)

al., 2007; Chan-Chan et al., 2012)

Poly(lactic acid) (PLA) Biodegradable (Wang et al., 2011a)

Poly(carbonate) (PCU) Biostables (Spirkova et al., 2011) Polybutadiene (PBD) Biostables (Thomas et al., 2009)

2012)

SPU´s traditionally has been used in medical devices due to their excellent mechanical prop‐ erties and an acceptable hemocompatibility. However, in the long term they suffer from poor biostability (Santerre et al., 2005). The main reason of this behavior is that living tissues are a very aggressive environment and even when the degradation of these polymers can be simulated by *in vitro* experiments, after *in vivo* usage they can be severally degraded. The *in vivo* failure of polymeric cardiovascular devices has been attributed to a combination of hy‐ drolysis, oxidation, environmental stress cracking and calcification. However, depending on the composition of the polymer one of these predominate over the other.

Polymer degradation in the biological environment results from the synergistic effects of the enzymes present in biological fluids, oxidizing agents and mechanical loads. For ex‐ ample, α-2-macroglobulin, cholesteryl esterase, A2 fosfolipase, K protease and B Cathe‐ psin are enzymes that are known to degrade polyurethanes (Zhao et al., 1993; Dumitriu, 2002). Even when some enzymes require very specific biological substrates, some of them seem to recognise and act over non biological substrates such as polymers (San‐ terre et al., 2005). White blood cells play also an important role in the *in vivo* degrada‐ tion. Some experiments conducted using implanted metallic cages have shown that neutrophiles, monocytes, monocyte derived macrophages (MDM) attach to polymer sur‐ faces, leading to the presence of multinucleated giant cells and foreign body reaction. It is generally accepted that one of the immediate immune responses by the body is the release of reactive oxygen species (ROS). In addition, neutrophils and monocytes release hypochlorous acid (HClO) and lysosomal hydrolases as part of their reaction to foreign surfaces. It has been also reported that activated MDM release ROS leading to the for‐ mation of hydrogen peroxide (Christenson et al., 2006; McBane et al., 2007). In addition, during the inflammatory reaction macrophages are able to lower the local pH up to 4. This condition can be simulated by following the ISO 10993 section 5.

Suntherland et al. (Sutherland et al., 1993) suggested that poly(ether urethanes) (PEU) cannot be significantly degraded by preformed products of phagocytic cells (such as cat‐ ionic proteins and proteases) or by activated oxygen species such as superoxide and hy‐ drogen peroxide. In view of the chemically stable nature of PEU, they hypothesized that the *in vivo* degradation of these materials might involve attack by chlorine-based and/or nitric oxide (NO)-derived oxidants, major oxidative products of activated phagocytes. Therefore, they exposed Pellethane to polymorphonuclear neutrophils (PMN) isolated from heparinized venous blood drawn from normal adult donors. The results reported support the idea that PMN-generated chlorine compounds are likely responsible for ini‐ tial damage to PEU after brief implantation and in addition to macrophage-derived NO and/or peroxynitrite (ONOO-).

Van Minen et al. (van Minnen et al., 2008) studied the *in vivo* (26 weeks of subcutaneous im‐ plantation in rats and 2.5 years in rabbits) degradation of porous aliphatic SPU based on bu‐ tanediisocyanate, DL-lactide-co-caprolactone soft segments and extended with BD-BDI-BD block urethane. After 1 week macrophages were observed along with giant cells and after 4 week phagocytosis was observed. The number of these cells was reduced with time but after 3 years fragments of the polymer remained. Furthermore, few macrophages were observed in the lymph nodes suggesting their local degradation.

Adhikari et al. (Adhikari et al., 2008) studied the *in vivo* degradation of two-part injecta‐ ble biodegradable polyurethane prepolymer systems (prepolymer A and B) consisting of lactic acid and glycolic acid based polyester star polyols, pentaerythritol (PE) and ethyl lysine diisocyanate (ELDI) using sheep femoral cortical defect model. No adverse acute or chronic inflammatory tissue response was noted in the interface tissues of the pre‐ cured polymer implants. By 6 weeks, there was direct apposition of new bone to the polymers. New bone and fibrovascular tissue was also observed within the porous spaces of the precured polymers by 6 weeks, and fluorochrome analysis suggested that this bone had started to be laid down at between 4 and 5 weeks. The polymer without β-tricalcium phosphate (TCP) showed histological evidence of some degradation by 6 weeks with progressive increase in polymer loss by 12 and 24 weeks. The polymer with β-TCP showed no evidence of degradation at 6 weeks and only minimal loss at 12 weeks. By 12 weeks, there had been considerable degradation of the polymers and at week 24, polymer was completely degraded.

Polymer degradation in the biological environment results from the synergistic effects of the enzymes present in biological fluids, oxidizing agents and mechanical loads. For ex‐ ample, α-2-macroglobulin, cholesteryl esterase, A2 fosfolipase, K protease and B Cathe‐ psin are enzymes that are known to degrade polyurethanes (Zhao et al., 1993; Dumitriu, 2002). Even when some enzymes require very specific biological substrates, some of them seem to recognise and act over non biological substrates such as polymers (San‐ terre et al., 2005). White blood cells play also an important role in the *in vivo* degrada‐ tion. Some experiments conducted using implanted metallic cages have shown that neutrophiles, monocytes, monocyte derived macrophages (MDM) attach to polymer sur‐ faces, leading to the presence of multinucleated giant cells and foreign body reaction. It is generally accepted that one of the immediate immune responses by the body is the release of reactive oxygen species (ROS). In addition, neutrophils and monocytes release hypochlorous acid (HClO) and lysosomal hydrolases as part of their reaction to foreign surfaces. It has been also reported that activated MDM release ROS leading to the for‐ mation of hydrogen peroxide (Christenson et al., 2006; McBane et al., 2007). In addition, during the inflammatory reaction macrophages are able to lower the local pH up to 4.

This condition can be simulated by following the ISO 10993 section 5.

and/or peroxynitrite (ONOO-).

56 Advances in Biomaterials Science and Biomedical Applications

in the lymph nodes suggesting their local degradation.

Suntherland et al. (Sutherland et al., 1993) suggested that poly(ether urethanes) (PEU) cannot be significantly degraded by preformed products of phagocytic cells (such as cat‐ ionic proteins and proteases) or by activated oxygen species such as superoxide and hy‐ drogen peroxide. In view of the chemically stable nature of PEU, they hypothesized that the *in vivo* degradation of these materials might involve attack by chlorine-based and/or nitric oxide (NO)-derived oxidants, major oxidative products of activated phagocytes. Therefore, they exposed Pellethane to polymorphonuclear neutrophils (PMN) isolated from heparinized venous blood drawn from normal adult donors. The results reported support the idea that PMN-generated chlorine compounds are likely responsible for ini‐ tial damage to PEU after brief implantation and in addition to macrophage-derived NO

Van Minen et al. (van Minnen et al., 2008) studied the *in vivo* (26 weeks of subcutaneous im‐ plantation in rats and 2.5 years in rabbits) degradation of porous aliphatic SPU based on bu‐ tanediisocyanate, DL-lactide-co-caprolactone soft segments and extended with BD-BDI-BD block urethane. After 1 week macrophages were observed along with giant cells and after 4 week phagocytosis was observed. The number of these cells was reduced with time but after 3 years fragments of the polymer remained. Furthermore, few macrophages were observed

Adhikari et al. (Adhikari et al., 2008) studied the *in vivo* degradation of two-part injecta‐ ble biodegradable polyurethane prepolymer systems (prepolymer A and B) consisting of lactic acid and glycolic acid based polyester star polyols, pentaerythritol (PE) and ethyl lysine diisocyanate (ELDI) using sheep femoral cortical defect model. No adverse acute or chronic inflammatory tissue response was noted in the interface tissues of the pre‐ cured polymer implants. By 6 weeks, there was direct apposition of new bone to the polymers. New bone and fibrovascular tissue was also observed within the porous The *in vivo* degradation of segmented poly(urethane urea)s (SPUUs) with hard segments de‐ rived only from methyl 2,6-diisocyantohexanoate (LDI) and PCL, PTMC (polytrimethylene carbonate), P(TMC-co-CL), P(CL-co-DLLA), or P(TMC-co-DLLA) as soft segment was con‐ ducted by Asplund et al. (Asplund et al., 2008). The *in vivo* study of SPUU-PCL using male Sprague-Dawley rats displayed the typical foreign body response seen with most inert poly‐ meric implant materials. The reaction at 1 week thus displayed an infiltration of ED1 posi‐ tive macrophages closest to the implant surface, an outside layer of fibroblasts and some collagen formation. At 6 weeks, the foreign body capsule had matured, displaying lower numbers of interfacial macrophages and an increased amount of collagen in the fibrotic cap‐ sule. The thickness of the foreign body capsule was similar to the controls. These observa‐ tions seemed also to be reflected in the number of ED1 positive macrophages, as well as in the total number of cells throughout the reactive capsule.

Hafeman et al. (Hafeman et al., 2008) synthesized polyurethane scaffolds by one-shot reac‐ tive liquid molding of hexamethylene diisocyanate trimer (HDIt) or lysine triisocyanate (LTI) and a polyol as hardener. Trifunctional polyester polyols of 900-Da and 1,800-Da mo‐ lecular weight were prepared from a glycerol starter and 60% ε-caprolactone, 30% glycolide, and 10% D,L-lactide monomers, and stannous octoate catalyst. Tissue response was evaluat‐ ed by subcutaneous implantation in male Sprague-Dawley rats for up to 21 days. During this time, initial infiltration of plasma progressed to the formation of dense granulation tis‐ sue. All of the implants showed progressive invasion of granulation tissue with little evi‐ dence of an overt inflammatory response or cytotoxicity. Fibroplasia and angiogenesis appeared to be equivalent among the different formulations. Extracellular matrix with dense collagen fibers progressively replaced the characteristic, early cellular response. The LTI scaffolds exhibited a greater extent of degradation at 21 days, although the incorporation of PEG into the HDIt scaffold accelerated its degradation significantly. Degradation rates were much higher *in vivo*. With time, each of the materials showed signs of fragmentation and en‐ gulfment by a transient, giant cell, foreign body response. After the remnant material was resorbed, giant cells were no longer evident.

Khouw et al. (Khouw et al., *2000*) reported that the foreign body response to degradable ma‐ terials differs between rats and mice. van Minnen et al., (van Minnen et al., 2008) also sug‐ gested that it is possible that the response between rats and rabbits differs as well, due to the faster degradation in the rabbit. This may be related to differences at the enzymatic or cellu‐ lar level, but also to the highly mobile and well vascularized skin of the rabbit, as compared to the rat.

#### **3.1. Calcification**

Mineralization or calcification (formation of various types of calcium phosphates such as apatite) is a well documented event in various medical devices, especially in those used in the cardiovascular field. Calcification is in fact, the most common macroscopic cause of failure in heart valves including those made of polyurethanes (Santerre et al., 2005). Even when calcification has been identified in heart valves, *in vitro* experiments on SPU showed little mineralization and associated exclusively to failure regions, indicating that the SPU´s have a lower intrinsic capacity for calcification compared to bovine bioprosthe‐ sis (Bernacca et al., 1997).

#### **3.2. Thrombosis**

Plasma protein adsorption is well accepted as one of the first events to occur when blood is in contacts with a biomaterial. These adsorbed proteins mediate the subse‐ quent interactions of cells and platelets with the surface and may induce thrombus for‐ mation, which remains one of the major problems associated with the long-term use of blood-contacting medical devices. The surface properties of the implanted materials are determinant in protein adsorption and biological interactions with the material. The ef‐ fects of various physicochemical properties such as surface hydrophilicity/hydrophobici‐ ty balance, surface charge density, ability to form hydrogen bonds, and chemical composition of biomaterials on protein adsorption as well as subsequent blood platelet adhesion have been investigated (Xu et al., 2010).

Antithrombogenicity is one of the essential requirements for a vascular graft, but it is very difficult to achieve. There are two common approaches employed to attain this goal. One is to develop biomaterials with inherent antithrombogenicity or to use sur‐ face modified biomaterials with an anticoagulant. The other approach is to quickly and completely endothelialize the inner surface of the tubular scaffolds, thereby, reducing thrombogenicity (Yan et al., 2007).

Thrombosis is a leading cause of vascular graft failure in small-diameter prostheses, where it leads to decreased flow or occlusion. In addition to inducing acute or suba‐ cute failure of grafts, it may be a cause of late failure owing to thrombosis superim‐ posed on stenosis due to other causes of vessel narrowing, such as intimal hyperplasia. Methods to improve vascular grafts (e.g., antithrombotic therapy) have been shown to be beneficial in decreasing graft occlusion after surgery. Agents known to inhibit thrombogenesis or promote anticoagulation (e.g., heparin, prostaglandin E1, hirudin, di‐ pyridamole, tissue factor pathway inhibitor and aspirin) have also been bound to the lumen of the synthetic vessels (Wang et al., 2007; Lu et al., 2012).

#### **3.3. Environmental stress cracking and metal ion oxidation**

Traditionally, SPUs have been used as permanent devices such as pacemaker leads insula‐ tion and ventricular assisting devices. When used as pacemaker lead insulators, they substi‐ tute silicone rubbers and have been used as biostable polymer for outer or inner insulated coating of coaxial bipolar pacemakers. Unfortunately, decades of experience showed that they were degraded by environmental stress cracking (ESC) or metal ion oxidation (MIO) or even autooxidation (AO) within a period of 28 and 34 months.

Environmental stress cracking includes crack formation and propagation on the surface of the polyurethane (Santerre et al., 2005). However, this type of degradation is a combination of the *in vivo* chemical degradation with the presence of mechanical stresses. In other words, polymer chain scission caused by the chemical degradation, create microscopic defects that are augmented by the presence of mechanical loads, leading to the formation of cracks on the surface (Wiggins et al., 2003). ESC it is also enhanced by the presence of residual stresses in the polymeric surface introduced during manufacturing and not eliminated during poly‐ mer annealing (Santerre et al., 2005).

The generally accepted *in vivo* degradation MIO mechanism involves the presence of hydro‐ gen peroxide (H2O2) produced by a variety of inflammatory cells (McBane et al., 2007; Chan‐ dy et al., 2009) and divalent metal ion such as Co2+ released from the lead. This reaction is known as the Haber-Weiss reaction and yields hydroxyl radicals that can attack α-methyl‐ ene groups in the polyether (PTMO based polyurethanes) to render hydroperoxydes with decompose in the presence of divalent cations rendering carbonyl groups that can accelerate (catalyse) further this decomposition (Kehrer, 2000; Wiggins et al., 2001). Polyether diol based polyurethanes are prone to oxidation and environmental stress cracking (ESC) (Król, 2007). However, polycarbonate based-polyurethanes (PCNUs) have been proven superior to polyether and polyester PUs, especially in terms of reduced ESC and metal ion oxidation (MIO), although they are still susceptible to hydrolysis (McBane et al., 2007).

Environmental stress cracking, calcification and thrombosis only became evident after a sufficiently long-term implantation of several years. For biodegradable PUs, which are designed to degrade in a relative short period (several months), the effective degrada‐ tion mechanisms are hydrolysis with or without the assistance of enzymatic catalysis (Chen et al., 2012).
