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Polyurethanes in vascular grafts.

Polyurethanes have unique mechanical and biological properties that make them ideal for many implantable devices. However, they are subject to some in vivo degradation mechanisms. 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 poly(carbonate)urethane has superior biostability in in vivo qualification tests compared to the traditional polyether polyurethanes.

We describe our approach to the design, development and production of a self-sealing vascular access graft, and the continuing development of a compliant small diameter vascular prosthesis using a unique patented fabrication process with a new generation of a biodurable poly(carbonate)urethane.

Background

Atherosclerotic cardiovascular disease is the number one cause of morbidity and mortality in the western world. This type of disease produces localized reduction in the calibre of arteries (stenosis) which ultimately leads to occlusion. It reduces or even stops the flow of blood through affected vessels. When this occurs, the overall consequences range from gangrene to stroke or to myocardial infarction.

Vascular bypass is one modality effective in the treatment of ischemic syndromes secondary to impeded blood flow. In the low flow small vessel setting, such as the coronary arteries and infrapopliteal vessels, the saphenous vein is the gold standard against which any prosthetic vascular conduit must compare. There is a significant number of cardiac and peripheral vascular cases in which the saphenous vein is either unsuitable (due to intrinsic disease such as sclerosis or varicosities), or is not available (due to stripping or previous vascular surgery). Another need exists in the area of hemodialysis for patients in end-stage renal disease. In such a setting, the need for an off-the-shelf arterial blood vessel substitute is vital.

Polyurethanes have been recognized for some time as an attractive and readily available range of materials for the fabrication of vascular prostheses. The elastic properties of the material, coupled with low thrombogenicity and exceptional physical and mechanical properties, has led to considerable research effort over the last 20 years, aimed at the development of polyurethane vascular grafts (ref. 1).

Vascular access grafts

Vascular access grafts are used in connection with hemodialysis treatments. Several acute and chronic diseases, including polycystic kidney disease, diabetes and hypertension, attack and may destroy normal kidney function, resulting in end stage renal disease (ESRD). ESRD is irreversible and currently approximately 60% of all ESRD patients are maintained by hemodialysis.

Hemodialysis requires that the patient be connected to an artificial kidney machine three times per week, for approximately four hours per session. A portion of the blood pumped from the heart must be routed to the artificial kidney machine, cleansed and returned to the patient. This routing is most efficiently accomplished by accessing the bloodstream at a point where the volume of blood flow is high. Two different methods are typically used. The first method, called an autologous AV shunt, involves cutting one of the arteries in the patient's arm and sewing the artery to an adjacent vein.

The second method requires that one end of a synthetic vascular graft be attached to an artery and the other end be sewn to a vein. In either case, blood flows very rapidly from the high pressure artery into the low pressure vein. A dialysis clinic technician inserts two large needles into the synthetic graft. One needle removes the blood and routes it to the artificial kidney machine and the second returns the blood to the patient. Approximately 54,000 new and replacement AV access grafts will be implanted in the U.S., with an average selling price of $425.

The improved sealing capability of polyurethane access grafts is especially important, as AV access grafts must be punctured in two places three times each week with large gauge needles to withdraw and replace blood cycled through an artificial kidney machine. Currently used synthetic grafts bleed profusely when needles used for hemodialysis treatments are removed and require a technician to apply pressure to the graft for up to 20 minutes to expedite clotting.

Polyurethane degradation

The word polyurethane can be applied to a huge number of different compositions with surprisingly varied applications. Generally speaking, polyurethanes used in medical devices are block copolymers consisting of aromatic or aliphatic polyurethane hard segments and polyester or polyether soft segments. These polymers have unmatched combinations of physical, chemical, electrical and biological properties. As a general rule, they offer high tensile strength and elongation, high tear strength, excellent wear resistance and superior biocompatibility (including blood compatibility). Depending on the ratio of hard to soft segments and the molecular weight of the segments, one can vary the hardness, lubricity, flexibility (elastic modulus) and many other properties.

Thus, polyurethane elastomers have been extensively evaluated for use as artificial heart diaphragms, heart valves, joint prostheses, vascular grafts, urethral catheters, mammary prostheses, penile prostheses and many other devices intended for long-term implant. The polyester polyurethanes were used as coverings for breast prostheses. Polyether polyurethane elastomers have been used as insulation for neurologic leads since 1975 and cardiac leads since 1977. Rigid polyether polyurethane is used in connector modules for implantable cardiac pacemakers, defibrillators and neurologic stimulators. Poly(ether)urethanes were considered to be the state of the art in terms of biocompatibility and mechanical properties. However, pacemaker leads fabricated from ether urethanes exhibited degrees of polymer degradation when implanted into living systems.

In 1983 Szycher (ref. 2) first proposed that poly(ether)urethanes were susceptible to in vivo oxidation of the polyether chain. The most susceptible group is the methyl group in the alpha position to the ether oxygen, which undergoes oxidation, causing eventual chain cleavage, leading to a significant reduction in molecular weight at the surface and eventual surface fissuring.

While it is not normally supposed that enzymes catalyze the degradation of synthetic polymers, the possibility exists that ether-based polyurethanes are degraded by enzymes in vivo. Williams (ref. 3) remarked that since enzymes have the ability to reduce the activation energy of chemical reactions, a degradation reaction that may normally only occur at elevated temperatures or in the presence of actinic radiation may conceivably take place under physiological conditions in the presence of the correct enzymes.

In a landmark study, Phua and Anderson (ref. 4) tested the biodegradation of polyurethanes by in vitro exposure to enzymes. Ultra-thin samples were exposed to two proteolytic enzymes, papain and urinase, at 37 [degrees] C for one to six months. Both enzymes were found to be capable of degrading poly(ether)urethanes.

Since papain is closely related to cathepsin B, a thiol endopeptidase which is released by cells of the inflammatory response, the authors concluded that segmented poly-etherurethanes can be degraded by enzymes which are present during the inflammatory response.

New generation polyurethanes

It has been reported by both Capone (ref. 5) and Szycher (ref. 6) that polycarbonate-urea-urethane is resistant to biological oxidation (surface fissuring) for periods of up to six months. Poly(carbonate) elastomers were fabricated as tubing stretched to 300% over mandrels and implanted subcutaneously in experimental animals. The tubing was retrieved at three and six months, evaluated by S.E.M. and FTIR analysis, with no evidence of degradation.

Stokes et al (ref. 7) compared poly(carbonate)urethanes to poly(ether)urethanes in vivo using the "Stokes test" which is designed to accelerate environmental stress cracking, using strain as the time accelerant. Extruded tubing was stretched over mandrels to 400% elongation. Poly(carbonate) urethanes showed no evidence of environmental stress cracking at eighteen months post-implantation.

Processing

Our process involves what we term `low temperature cast coagulation." In this system we gently extrude polymer solution onto the smooth surface of a mandrel. The mandrel and extrusion head rotate in synchronization, minimizing shear and residual stress, while a pair of 2.5 meter long mandrels are drawn through the twin extrusion heads into a coagulant maintained at 40 [degrees] C.

Mandrel rotation and transverse speed, extrusion head rotation speed and polymer pump pressure are all electronically controlled and coordinated to give the desired structure and geometry to the graft wall. By controlling the process conditions, grafts can be produced with a wide range of physical and mechanical characteristics; an example would be a very porous graft allowing rapid cellular ingrowth which could be suited to venous applications.

The polymer solution comprises a solution grade poly(carbonate)urethane, a water soluble filler of between 10% and 60% by weight and a surfactant in an amount between 1% and 10% by weight. During phased coagulation the fillers prevent collapse of the structure as the solvent disperses and the filler dissolves into the coagulant, resulting in a single layer uniform microporous structure. A single layer structure avoids the danger of delamination and subsequent loss of strength, maintains cross section and allows the surgeon to cut the graft evenly and cleanly). The specification of the polyurethane-based vascular grafts is shown in table 1.

Experimental implants - vascular access

To assess initial biological response, we have implanted four 5 mm ID grafts into canines as bilateral aorto - femoral bypasses with the iliac arteries tied off. Patency was monitored monthly by color-coded Duplex examination.

All grafts were patent at explant. Light microscopy showed a well developed neointima lining the anastomotic region of the graft without apparent lumenal reduction. The surrounding capsule was composed of tissue macrophages and foreign body type giant cells.

Conclusions

We have presented our methods and materials used to fabricate vascular grafts. We expect that a combination of proven technology, product design, and the use of a second generation poly(carbonate) urethane, may lead to clinically successful compliant polyurethane vascular grafts.

References

(1.) Underwood, C.J., Tait, W.F. and Charlesworth, D., "Design considerations for a small diameter vascular prosthesis," Int. J. Artif. Organs., 1988, 272-276.

(2.) Szycher, M., "Surface fissuring of polyurethanes following in vivo exposure," ASTM, STP 859, 308-321, Fraker and Griffin, eds., Philadelphia.

(3.) Williams, D.F., J. Bioeng., 1,279-294,1977.

(4.) Phua, S.K. and Anderson, J.M., "Biodegradation of a polyurethane in vitro," J. Biomed. Mat. Res., 21, 231-246, 1987.

(5.) Capone, C.D., "Biostability of anon ether polyurethane," J. Biomat. App., Vol. 7, Oct., 1992.

(6.) Szycher, M., et al., "In vivo testing of biostable polyurethane," J. Biomiat. App., Vol. 6, Oct., 1991.

(7.) Stokes, K., et al, "Polyurethane elastomer biostability," J. Biomat. App., Vol. 9, April, 1995.
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Author:Szycher, Michael
Publication:Rubber World
Date:Apr 1, 1998
Words:1737
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