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thane heparin solution coating.

Gutowska et a!.

Heparin loading and release frum tem

(1992, 1994, 1995)

pera tu re-sensitive polymers.

polymer, and the heterogeneity ot heparin within the polv-mer matrix.

Heparin was dispersed in silicone rubber or silicone rubber/ graphite and fabricated as canine arterial grafts (Hufnagel et at, 1968). This device was able to prevent thrombus formation in the animal for a minimum of 2 hr. Heparin was also incorporated into a poly(hydroxyethyl methacrylate) monolithic device and the release of heparin was detected for —10 hr (Ebert et ai, 1980). In addition, low-molecular-weight heparin fractions were released more rapidly than poly dispersed heparin from polyutethane and polyvinyl chloride membranes, demonstrating a rapid initial release, followed by a more gradual release (Ebert and Kim, 1984). Hep arm-releasing polyurethane catheters were also implanted in canine femoral and jugular veins, exhibiting significant reduction of thrombus formation after 1 hr (Heyman et ai, 1985).

Lin et ai. (1987) and Nojlri et ai. (1987) coated the lumen of polyurethane tubing with a Biomer-heparin solution and implanted the device as an A-A shunt in rabbits. The heparin release profile demonstrated an initial burst effect, followed by first-order release (see Chapter 7.8). The systemic effects of heparin released from shunts were determined by APTT, platelet number, and platelet aggregation assays. Without rinsing the heparin monolithic coating, heparin released from the device caused changes in systemic blood parameters for 30 mm, and extended APT!" values were observed for 15 min after insertion of the shunt. Using a 15-min rinse protocol, the release of heparin did not affect systemic blood hemostasis, and the APTT values were within control values (Fig. 3), There was a consistent decrease in platelet numbers 15 min after the shunts were inserted, which subsequently rebounded to normal control values. Platelet aggregation with adenosine diphosphate

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