Nonthrombogen1c Treatments And Strategies

No rinse, heparin burst 15 min rinse, no heparin hurst

No rinse, heparin burst 15 min rinse, no heparin hurst

Shunt Patency Times (hrs)

FIG. 3. APTT measurements during A-A shunt patency.

(ADP) remained relatively constant when adjusted for platelet number (Fig. 4). The occlusion time for the control, nonhepa-rinized polyurethane was 30 ± 4 min. This compares well with Biomer shunts, which occluded in 45 ± 7 min. The heparin-releasing matrix tubing had prolonged patency times ranging from 1 to 5 hr, decreasing with increased rinse times. Thus, heparin released from the shunt was effective in improving the blood compatibility of the device.

The ex vtvo heparin level at the time of shunt occlusion was estimated from parallel elution of [,4C]heparin. For experiments in which the shunts were prerinsed, the heparin release level to prevent shunt occlusion was consistently around 4 X 10 ^ g/cm2/min. This release rate was termed the minimum critical heparin release rate for nonthrombogen-

Time (hrs)

FIG. 4. Platelet counts during A-A shunt implantation.

Time (hrs)

FIG. 4. Platelet counts during A-A shunt implantation.

icity of the heparin-releasing matrix (Lin et al., 1987). Tanzawa et al. (1973) studied heparin elution from cationic hydrophilic copolymers composed of vinyl chloride, ethylene, and vinyl acetate that bind heparin by coulombic interactions. The in vivo test of these heparin-releasing polymers showed that a constant heparin elution rate of 4 x 10~8 g/cm2/min will keep an inferior vena cava indwelling catheter thrombus free for 2 weeks. Basmadjian and Sefton (1983) concurred with this value by using a theoretical flow model to show that heparin release will maintain the interfacial heparin concentration above the minimum concentration for prophylactic minidose heparin therapy.

The synergism between platelet aggregation and fibrin formation in vivo stimulated the synthesis of a covalently bonded conjugate of commercial-grade heparin and PGE[. This dual-acting drug conjugate was evaluated for use as a controlled release system for blood-contacting surfaces in order to improve the blood compatibility of polymer surfaces. The compound was synthesized using a modified mixed carbonic anhydride method (Meienhofer, 1972) of amide bond formation between the carboxylic acid moiety of PGE! and a primary amine group on heparin. Bioactivity tests on PGE!—heparin conjugates (Jacobs and Kim, 1986) (APTT and platelet aggregation) confirmed that the antithrombic activity of heparin was maintained. However, PGEi bioactivity, as measured by ADP-induced platelet aggregation, decreased, but was still active. Rabbit A-A shunt experiments revealed that heparin—PGE, released from the polyurethane device prevented both fibrin formation and platelet activation (Jacobs et al., 1989). This approach is promising owing to the dual activity of the released compound which prevents both platelet activation and fibrin formation at the blood—polymer interface.

A new approach to create heparin-releasing systems was presented by Gutowska et al. (1992). They incorporated heparin into thermosensitive hydrogels (TSH), hydrogels that exhibit higher swelling at low temperature than at body temperature, as illustrated in Fig. 5. TSH were fabricated with N-isopropyl acrylamide (NiPAAm) and copolymerized with butyl methacrylate or acrylic acid. These NiPAAm/TSH were combined with polyurethane to form a novel interpenetrating polymer network (1PN) (Gutowska et al., 1992), and heparin loading and release was studied from this device. Equilibrium swelling studies showed that modification of NiPAAm gel with polyurethane via a semi-IPN formation did not affect the gel collapse point, but resulted in decreased thermosensitivity and lower swelling levels. It was hypothesized that NiPAAm/poly-urethane semi-IPNs, in which the polyurethane network is not cross-linked and therefore more susceptible to phase separation, will form a microporous structure due to the enhanced phase separation of components in the hydrated state (Gutow-saka et al., 1994). This TSH-hepann-IPN was coated onto polyurethane catheters followed by heparin loading in low temperature solution and implanted as vascular access catheters in dogs. The device was removed after 1 hr to evaluate surface-induced thrombosis. The amount of thrombus attached to control surfaces (bare polyurethane or IPN without heparin) was significantly greater than to heparin-releasing surfaces (Gutowska et al., 1995). Therefore, the TSH material was able to absorb heparin at low temperature and release sufficient

polymer coating (below LCST), gel gel collapse, swell, entrapping heparin released heparin

FIG. 5. Surface-grafted thermosensitive polymer release mechanism. Polymer swelling and heparin loading at low temperature; polymer collapse and heparin release at body temperature.

polymer coating (below LCST), gel gel collapse, swell, entrapping heparin released heparin

FIG. 5. Surface-grafted thermosensitive polymer release mechanism. Polymer swelling and heparin loading at low temperature; polymer collapse and heparin release at body temperature.

amounts at body temperature to minimize surface-induced thrombus formation in vivo.

The fact that heparin is a highly negatively charged polysaccharide led researchers to bind heparin onto a cationic surface through ionic binding. A summary of ionically bound heparin surfaces is presented in Table 2. However, the usefulness of ionically bound heparin surfaces is limited, primarily due to the short-term release of heparin.

As another approach, Kwon et al. (1995) prepared and investigated an electroerodible polyelectrolyte complex for the pulsatile release of heparin. An insoluble polyelectrolyte com plex was formed by combining two water-soluble polymers, poly(allylamine) and heparin. Upon the application of an electric current, a rapid structural change of the complex occurred, dissolving the polymer matrix in proportion to the intensity of an applied electric current. The disruption of ionic bonds in the polymer matrix attached to the cathode and subsequent release of heparin was due to the locally increased pH near the cathode from hydroxyl ion production. Thus, the release pattern of heparin followed the applied electric current, primarily due to surface erosion of the polymer matrix, as illustrated in Fig. 6.

TABU 1 Ionically Bound Heparin Surfaces

Investigators

Method

Gott et al. (1966) Leninger et al. (1966) Grode et al. (1972)

Tanzawa et al. (1973, 1978)

Barbucci et al. (1984)

Tanzi and Levi (1989) Kwon et al. (1994)

Prior adsorption of cationic surfactant onto graphite-coated surfaces, followed by ionic heparin binding.

Formation of quaternary amine groups on the surface of polymers, followed by ionic heparin binding.

Prior adsorption of tridodecyl methyl ammonium chloride onto the preswol-len material, followed by heparin binding.

Formation of quaternary amine groups in the bulk copolymer preparation, followed by ionic binding of heparin.

Synthesis of polymers containing amido and tertiary amine groups or premodi-fication of polymer by grafting on poly(amido amines), followed by ionic binding of heparin.

Synthesis of polyurethane-containing polyiamido amine) blocks.

Heparin binding to cationic polymer to form insoluble complex; complex dissolution and heparin release under electric current.

Heparin Immobilized onto Polyurethane

The main alternative to a heparin-release device is the covalent coupling of heparin or other anticoagulant drugs directly onto the polymer surface. This procedure effectively

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