Polyurethanes have unique mechanical and biologic properties that make them ideal for many implantable devices. They are subject to some in vivo degradation mechanisms, however. Polyester polyurethanes are subject to hydrolytic degradation and are no longer used in long-term implanted devices. Polyether polyurethanes, while hydrolytically stable, are subject to oxidative degradation in several forms, including environmental stress cracking and metal ion oxidation. Mineralization is also known to occur. A new polycarbonate polyurethane has superior biostability in early in vivo qualification tests compared to the polyether polyurethanes, including no evidence of hydrolysis, ESC or MIO.

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Polyether polyurethanes are extremely interesting for use in implantable devices. They are, however, susceptible to autoxidative degradation and stress cracking. One approach to improving biostability is to replace some of the polyether with polysiloxane. Shore 80A polyether polyurethanes with 20% (PS-20) and 35% (PS-35) polysiloxane were strained to 400% elongation and implanted in rabbits. Twelve weeks implant showed that both were significantly more biostable than their polysiloxane-free controls. After 18 months implant, PS-20 developed some localized tensile fractures. PS-35 showed no sign of visual damage. Infrared surface analysis does not allow direct evaluation of autoxidation because the Si--O--Si stretch peaks mask the polyether bands. Secondary indicators suggest possible very slight autoxidation of both PS-20 and PS-35 surfaces, but not enough to develop cracks. The polysiloxane-free controls did show substantial infrared evidence of autoxidation. Molecular weights of long-term PS-20 and PS-35 explants were negligibly lower. In comparison, the polysiloxane-free control suffered 35% molecular weight loss. Positive and negative controls performed as expected. PS-20 is recommended for devices that do not sustain high fixed loads. PS-35 is dramatically more biostable than its unmodified polyether analogues and is recommended for use in chronically implantable devices.

Polyether polyurethanes are subject to oxidation catalyzed by and through direct (redox) reaction with transition metal ions (cobalt), released by corrosion of metallic parts in an implanted device. Replacing part of the polyether with polysiloxane appears to reduce susceptibility to metal ion oxidation (MIO). In vitro studies indicated that polyurethanes containing 20-35% polysiloxane (PS-20 and PS-35) are about optimum. We implanted tubing samples containing cobalt mandrels in the subcutis of rabbits for periods up to 2 years. After 2 years, only traces of microscopic cracks were seen on half the PS-35 samples, PS-20 significantly delayed MIO, while the polysiloxane-free control was very severely degraded. Infrared spectroscopy established that polyether soft segment oxidation was occurring in PS-20. We could not directly evaluate oxidation in PS-35 because siloxane bands mask the aliphatic ether. Indirect FTIR evidence suggests that there is very slight polyether oxidation that develops early, and then seems to stabilize. The molecular weight of degraded PS-20 decreased. That of microcracked PS-35 decreased negligibly while that of undamaged PS-35 increased slightly after 2-year in vivo. The polysiloxane-free control was profoundly degraded. PS-20 has much improved MIO resistance, while that for PS-35 is highly MIO resistant compared with its polysiloxane-free control.

In recent years, pacemaker lead failure due to compressive damage has been reported with increasing frequency. To document the mechanism of this failure, we evaluated explanted mechanically damaged leads with electrical testing, optical microscopy, and in some cases, scanning electron microscopy (SEM). In addition, we performed an autopsy study to measure the compressive loads on catheters placed percutaneously through the costoclavicular angle, as well as by cephalic cutdown. Of the 49 explanted compression damaged leads with enough clinical data for analysis, all had been placed by percutaneous subclavian puncture. Our autopsy data confirmed the significant increase in pressures generated in the costoclavicular angle for medial percutaneous subclavian catheterization (126 +/- 26 mmHg) compared to a more lateral percutaneous subclavian puncture (63 +/- 15 mmHg) or a cephalic cutdown (38 +/- 13 mmHg) (P < 0.01). In vivo coil compression testing documented loads up to 100 pounds per linear inch of coil and a compressive morphology by SEM identical to that seen in the clinical explants. Pacemaker leads appear to be susceptible to compression damage when placed by subclavian venipuncture. When possible, leads should be placed such that they avoid the tight costoclavicular angle.