Stent coating for releasing medicine and method therefor

Technical field to which the invention belongs

This patent application is a partial continuation application of related application No. 08 / 526,273 filed on September 11, 1995, and a part of related application No. 08 / 424,884 filed on April 19, 1995. It is also a continuation application. All matters not included in this patent application and included in these parent applications are incorporated herein by reference for any purpose. For the application No. 08 /, the title of the invention, the "drug release stent coating method", and the partial continuation application of the above-mentioned application, which are common to this case and the inventor and the assignee and filed on the same day. , Referred to each other. Matters not included in this patent application are incorporated herein by reference for any purpose.

The present invention generally relates to providing a biologically stable elastomeric coating on the surface of the implant, wherein the coating comprises a biologically active material having controlled release properties within the coating. Is what you do. The present invention relates specifically to providing a non-thrombogenic surface during or after time-controlled release of a biologically active substance. The present invention is specifically described with respect to coatings on an expandable stent prosthesis in the treatment of intraluminal transplantation, eg, vascular grafting.

Conventional technology

In surgical or other related invasive processes, inserting and dilating a stenting device into a blood vessel, ureter or other inaccessible site prevents restenosis and vascular or luminal wall support or reinforcement. It has become the usual form of long-term treatment to provide the body and for other therapeutic or health-restoring functions. Typically, such a prosthesis is applied to the desired position using a vascular catheter or similar penetrating device to bring the stent to the desired position. There, the stent is then released and expanded or expanded in situ. These instruments are generally designed as permanent implants and may be implanted in blood vessels or other tissues that come into contact at the site of implantation.

One type of self-expanding stent has a flexible tubular body formed from several thread elements, each of which is flexible. Each thread element extends spirally, and the centerline of the tubular body acts as a common axis for each thread element. These thread elements are wound in the same direction but are axially offset from each other. And although it is also arranged along the axis, it intersects with a similar number of elements that are wound in the opposite direction. This arrangement provides an elastic reticulated tubular structure that provides a stable spread upon relaxation. Axial tension causes stretching as well as corresponding diametrical contraction, which causes the stent to be placed on the catheter device and carried through the tubular system as a thin stretched device. Once the tension relaxes in situ, the instrument at least substantially returns to its original form. Class prostheses that include reticulated flexible tubulars are illustrated and described in Wallsten's US Pat. Nos. 4,655,771 and 4,954,126, and Wallsten et al., US Pat. No. 5,061,275.

Implanted stents have been used to carry medical reagents such as thrombus disintegrants. Froix's US Pat. No. 5,163,952 discloses a temperature-dependent shape memory expansion plastic stenting device designed to carry medical reagents in the material of the stent itself. Pinchuk, in US Pat. No. 5,092,877, discloses a stent made of a polymeric material that may have a coating associated with drug delivery. Other patents relating to the class of instruments using biodegradable or bioabsorbable polymers include Tang et al., US Pat. No. 4,916,193 and Mac Gregor, US Pat. No. 4,994,071.

Sahatjian's patent, US Pat. No. 5,304,121, discloses a coating applied to a stent consisting of a hydrogel polymer and a preselected drug such as a cell growth inhibitor or heparin. Further methods for producing coated intravascular stents for delivering therapeutic material are described in Berg et al., US Pat. No. 5,464,650, issued November 7, 1995, and the corresponding November 1994. It is described in European Patent Application No. 0,623,354 A1 published on the 9th. In this disclosure, the polymer coating material is dissolved in a solvent and the therapeutic substance is dispersed in the solvent; the solvent evaporates after application.

The 22nd International Conference on the Development of the Society of Engineering Technology for Materials and Methods, "Medical Instrument Design-System Approach: Central Venous Catheter" by Michael N. Helmus (co-inventor of the present invention), is heparin. Concers about polymer / drug / membrane systems for releasing. These polymer / drug / membrane systems require two different types of layers to function.

When blood comes into contact with the surfaces of foreign objects in vivo, the thrombus-forming reaction tends to be induced, and the more surface area of the foreign device that comes into contact with the host blood, the more clots and clots form on these surfaces. It has been recognized that the trend will increase. This has resulted in the use of immobilized systemic anticoagulants or thrombolytic agents such as heparin on the surface of oxygen inhalers and the like in contact with blood to reduce this phenomenon. Such techniques are described in US Pat. Nos. 5,182,317, 5,262,451 and 5,338,770 by Winters et al., In which the amine functional group of the active substance uses polyethylene oxide (PEO) on the surface of the siloxane. Is covalently bonded.

Another approach is described in US Pat. No. 4,613,665 granted to Larm. In this approach, heparin is covalently immobilized on a plastic surface material, which contains a primary amino group to confer non-thrombogenicity. Other approaches to fixing heparin are described in the following literature: Barbucci, et al. "Coating of commercially available materials with a". new heparinizable material), Journal of Biomedical Materials Research, Vol. 25, 1259-1274 (1991); Hubbell, JA, "Pharmacologic Modification of Materials" Cardiovascular Pathology, Vol. 2, No. 3. (Suppl.), 121S-127S (1993); Gravlee, GP, "Heparin-Coated Cardiopulmonary Bypss Circuites", Journal of Cardiothroacic and Vascular Anesthesia, Vol 8, No 2, pp 213-222 (1994).

For stents, polymer stents are effective, but may have inferior mechanical properties compared to metal stents of similar thickness and mesh. Metallic vascular stents with a uniform, relatively fine metal mesh can significantly increase their strength to withstand inward circumferential pressure. In order to obtain the same strength with the polymer material, it is necessary to make the side wall thicker, to make the woven structure (weave) of the filament heavier, and to make the woven structure denser. Therefore, the cross-sectional area reserved for flowing through the stent is reduced, and / or the amount of interstitial space in the woven tissue is also reduced. Placing and delivering a polymeric stent using a catheter delivery system is generally more difficult.

Certain types of stents, such as reticulated metal stents, may be preferred for certain applications. However, the coating and coating modification methods of the present invention are not limited in this way and can be used for various prostheses. Thus, in the case of stents, the invention can be applied, for example, to non-self-expandable stents, including non-self-expandable stents that can be expanded with a balloon. The present invention can also be applied to any type of polymeric stent. Medical instruments that can benefit from the present invention include, for example, blood exchange devices, vascular access ports, central venous catheters, cardiovascular catheters, external circuits, vascular grafts, pumps, heart valves, cardiovascular sutures and the like. .. Regardless of the detailed embodiments, the scope of the invention is not limited with respect to the design of the implant, the location of the implant or the material of the components. In addition, the invention can also be used for other types of implantable prostheses.

<p> Therefore, the first object of the present invention is a coating capable of effectively controlling and delivering a biologically active substance over a long period of time, and a stent used as a stent prosthesis placed on the affected area. To provide a method of coating.</p><p> Another object of the present invention is to provide a method for coating a coating and a stent prosthesis with a biologically stable hydrophobic elastomer. Here, the hydrophobic elastomer of the coating comprises a biologically active substance.</p><p> A further object of the present invention is to provide a multi-layer coating and a method of delivering a biologically active material, where the proportion of active material may vary from layer to layer.</p><p> A further object of the present invention is to provide a multilayer coating and a method of delivering a biologically active substance from the coating, where the coating has a non-thrombogenic surface.</p><p> A further object of the present invention is to provide a multi-layer coating that delivers a biologically active substance, such as heparin, that has a fluorosilicone top layer.</p><p> A further object of the present invention is to provide a multi-layer coating that delivers a biologically active substance, such as heparin, with a surface having immobilized polyethylene glycol (PEG).</p><p> Other objects and advantages of the present invention will be further clarified by those skilled in the art understanding the specification and the scope of claims contained therein.</p>

<p> The present invention provides a relatively thin layered coating, the coating of which is made of a biologically stable elastomeric material. The elastomeric material also contains a predetermined amount of biologically active substance within it and has a non-thrombogenic surface. Such surfaces are useful for coating the surface of deployable prostheses, such as stents.</p><p> The stent to be coated is preferably a tubular stent prosthesis that is self-expanding and has an open end. Other materials, including polymeric materials, can be used, but in a preferred embodiment, the tubular body is a self-expanding, interstitial network of metal wires of a single or multiple filaments. Consists of. This metal wire bends without depression, easily deforms axially into an elongated shape, is inserted into the lumen through a vascular catheter, and is removed in situ. Stable expansion to a predetermined diameter.</p><p> In this method, the initial coating is preferably applied as a mixture, solution or suspension of the polymeric material with a finely divided biologically active material. Here, such a substance may be dispersed in an organic medium. It may also be a solution of a solvent or medium or a partial solution, where the medium is a medium of a polymer and / or a biologically active substance. For the purposes of this patent application, the term "finely divided" means an inclusive material of any kind or size, which is a molecule dissolved from a mixture of suspensions, colloids and microparticles. is there. The active material is dispersed inside the carrier material, which carrier material may be a polymer, a solvent, or both. The coating is preferably applied as a plurality of relatively thin layers, which are preferably applied sequentially and in a relatively quick order, and are applied to the stent in a warp-expanded state. Is preferable.</p><p> In many applications, a layered coating means having an undercoat and a topcoat, and is also characterized by having an undercoat and a topcoat. The ratio of the coating thickness of the undercoat to the topcoat depends on the desired effect and / or the elution system. Typically, they have different compositions from each other, the undercoat contains almost or all active material, and a non-thrombogenic surface is found in the topcoat.</p><p> The coating can be applied by dipping or spraying, in which case a volatile solvent material with a relatively high vapor pressure is used to obtain the desired viscosity and quickly increase the thickness of the coating layer. Determine. A method of reciprocally spray coating using an air brush device while rotating the longitudinally expanding stent is preferred. By this coating method, it is possible to adhere to the entire surface of the filament having a structure with a gap in the stent so that the shape can be adapted and the entire surface of the filament can be coated. At the same time, this coating method allows the coating apparatus to maintain a mesh-like or other pattern of interstitial lattice structure.</p><p> The coating is exposed to the ventilator for a predetermined time (eg, 1 hour or longer) at room temperature to evaporate the solvent medium. In the case of certain undercoat materials, the polymer material is then cured at room temperature or at elevated temperatures. Curing is defined as the conversion of an elastomeric or polymeric material into a final or useful state by inducing physicochemical changes by heating and / or using chemicals. For example, a polyurethane thermoplastic elastomer can be used as the undercoat material and the solvent can be evaporated at room temperature. This allows the undercoat material to be useful for controlled drug release without further curing.</p><p> The ventilation time and temperature that can be applied for curing is determined by the particular polymer involved and the particular agent used. For example, a silicone material, i.e. a polysiloxane material (eg, polydimethylsiloxane) can be preferably used. Urethane prepolymers can also be used. Unlike polyurethane thermoplastic elastomers, some of these materials are applied as prepolymers in the coating composition, so it is essential that they be cured by heat. Preferred silicone species have a relatively low curing temperature and are known, for example, as materials that can be vulcanized at room temperature (RTV). Some of the polydimethylsiloxane materials can be cured, for example, by exposure to air at about 90 ° C, for example for a period of 16 hours. The curing step may be performed both after the undercoat or some lower layer has been applied and after the top layer has been applied, or once after the coating is complete. Only the curing step may be performed.</p><p> The coated stent is exposed to a post-curing treatment step, which treatment steps include inert gas plasma treatment and sterilization, which may include, for example, gamma irradiation, ETO treatment, electron beam treatment, vapor treatment. Can be mentioned.</p><p> In plasma treatment, an unconstrained coated stent is placed in the reaction chamber and the system is purged with nitrogen gas to a vacuum of 20-50 milittle. An inert gas (argon, helium or a mixture thereof is then introduced into the reaction chamber for plasma treatment. One method uses argon (Ar) gas, with an output range of 200-400 watts, 150-150-per minute. Treat with a flow rate of 650 standard milliliters (this flow rate corresponds to about 100-450 milittles) and an exposure time of 30 seconds to about 5 minutes. Immediate removal of the stent after plasma treatment. Or can be retained in an argon atmosphere for an additional predetermined time, typically 5 minutes.</p><p> In the present invention, the topcoat, i.e., the surface coating, may be applied by any of several methods, which can further control thrombus formation and, if desired, release profiles, especially It is possible to control the initial significantly higher release rates associated with heparin elution.</p><p> In one embodiment, an outer layer of fluorosilicone is applied to the undercoat as a topcoat. The outer layer may also contain heparin. In another embodiment, polyethylene glycol (PEG) is immobilized on the surface of the coating. In this method, the underlayer is treated with an inert gas plasma and then immediately ammonia (NH).₃) Plasma treatment to aminate the surface. As used herein, amination means producing imino groups and other nitro containing species on the surface. It is then immediately immersed in an electrophileically activated polyethylene glycol (PEG) solution with a reducing agent, such as sodium cyanoborohydride.</p><p> Coated and cured stents with a modified outer layer or surface are ultimately sterilized by gamma irradiation, usually 2.5-3.5 millirad. Argon (Ar) plasma treated stents, whether restrained upon exposure, enjoy sufficient elasticity after irradiation. However, when a stent that has not been previously treated with argon plasma is gamma-sterilized while being restrained, it loses its elasticity and does not recover at a sufficient or appropriate rate.</p><p> The elastomeric material that forms the underlayer of the stent coating should have certain properties. The lower layer should preferably be composed of a hydrophobic and biologically stable elastomeric material that does not decompose. The surface layer material is preferably one that minimizes tissue rejection and tissue inflammation and is capable of being encapsulated by tissue adjacent to the site where the stent was implanted. The material to be exposed should be designed to be less prone to coagulation by contacting blood, and the surface is preferably modified accordingly. Therefore, the lower layer of the material is preferably provided with a fluorosilicone outer coating layer, which may or may contain an embedded bioactive substance, such as heparin. You don't have to. Alternatively, the outer coating may be substantially composed of polyethylene glycol (PEG), polysaccharides, phospholipids or combinations thereof.</p><p> Polymers generally suitable for undercoating or underlayers include silicones (eg, polysiloxanes and substituted polysiloxanes), polyurethanes, thermoplastic elastomers in general, ethylene-vinyl acetate copolymers, polyolefin elastomers, polyamides. Examples include elastomers and EPDM rubbers. The material is considered hydrophobic in the environment in which the present invention is considered. Materials for the surface layer include fluorosilicone, polyethylene glycol (PEG), polysaccharides, phospholipids, and combinations thereof.</p><p> Heparin is preferred as the active substance to be included. However, the drugs included include antithrombotic agents, anticoagulants, antibiotics antiplatelet agents, thrombolytic agents, antiproliferative agents, steroidal anti-inflammatory agents, non-steroidal anti-inflammatory agents, hyperplasia, in particular, Antithrombotic agents, smooth muscle cell inhibitors, growth factors, growth factor inhibitors, cell adhesion inhibitors, cell adhesion promoters, and healthy neointimal tissue (including endothelial cell regeneration) formation promoters Can be mentioned. Positive effects may be obtained by suppressing certain cells (eg, smooth muscle) or tissue formation (eg, fibrous muscular tissue), while different cell migration (eg, endothelium) and tissue formation (eg, eg, fibrous muscle). , New intima tissue) may be promoted.</p><p> Suitable materials for obtaining reticulated stents include stainless steel, tantalum, titanium alloys and certain cobalt alloys. Here, the titanium alloy includes nitinol, that is, a nickel-titanium alloy material that memorizes its shape by heat. Cobalt alloys also include cobalt-chromium-nickel alloys such as Elgiloy® and Phynox®. For more information on the manufacture and other aspects of the stent itself, see US Pat. Nos. 4,655,771 and 4,954,126 (both patents granted to Wallsten) and US Pat. No. 5,061,275 (Wallsten el al), which have already been referenced. These documents are incorporated herein by reference.</p><p> Various combinations of polymeric coating materials can be used with the biologically active material of interest and can achieve the desired effect when coated on the stent of the invention to be implanted. There are many types of filling of therapeutic materials. The mechanism by which a biologically active substance is incorporated into a surface coating and the mechanism by which it is expressed depends on the properties of both the surface coating polymer and the active substance to be incorporated. The release mechanism also depends on the built-in mode. The bioactive material may be eluted through the passage between the particles, or may be administered by transport or diffusion through the material contained therein.</p><p> As used herein, "elution" is the extraction or release of a substance through direct contact with body fluids, involving a release process involving particles that connect to the outside of the coating and pass through a passage between the particles. Is defined as. "Transportation" or "diffusion" is defined as including a release mechanism by which a released substance passes through another substance.</p><p> The desired release rate profile is the thickness of the coating, the longitudinal (ie, layer-to-layer) distribution of the bioactive material, the mixing method, the amount of the bioactive material, the combination of different matrix polymer materials in different layers and the polymer. It can be adjusted by changing the crosslink density of the material. The crosslink density is related to the amount of crosslinks that actually occur and the relative air density of the matrix, which is caused by the particular crosslinker used. During the curing process, the amount of cross-linking is determined, and thus the cross-linking density of the polymeric material. For bioactive substances released from the crosslinked matrix, such as heparin, when the density of the crosslinked structure increases, the release time becomes longer and the burst effect decreases.</p><p> A thin top layer of silicone that does not carry the drug has advantages and can further control the elution of the drug. However, in the case of heparin, for example, when the topcoat or surface coating is modified, the initial heparin release profile can be further controlled, or significant advantages can be gained by making the surface more non-thrombogenic. It should be understood that it is done.</p>

Embodiment of the invention

In the present invention, the stent coating incorporates a biologically active substance for time-controlled delivery to the body lumen of interest in vivo, in situ. In such stent coatings, spraying the prepared coating solution or suspension results in many thin layers, in which biologically active substances are incorporated. The process of such a method is outlined in FIG. The coating solution or suspension is prepared at 10 as described below. If a cross-linking agent is added, the desired amount of cross-linking agent is added to the suspension or solution at 12 and then the mixture is stirred to give a homogeneous coating composition at 14. After that, the composition is copied to a coating container or a coating device, and the coating container or the coating device may be a container for 16 spray coatings. Typical preparation examples of coating solutions used for heparin and dexamethasone are described below.<u style="single">General manufacture of heparin undercoat composition</u> Silicone was obtained as a mixture of polymer precursors in solvent (xylene). For example, the weight of 35% solid silicone in xylene was obtained from Applied Silicone, Part # 40,000. First, the silicone-xylene mixture was weighed. The solid silicone content was measured according to Vendor's method of analysis. A pre-calculated amount of finely divided heparin (2-6 microns) was added to the silicone, followed by tetrahydrofuran (THF) HPCL grade (Aldrich or EM). For example, for a 37.5% heparin coating, W_silicone= 5g; Solid% = 35%; W_hep= 5x0.35x0.375 / (0.625) = 1.05g. The required amount of THF (44 ml) in this coating solution is the formula: W for a 37.5% heparin coating solution._silicone/ W_he_pCalculated using = 0.04. Finally, the manufacturer's cross-linking agent solution was added using a Pasteur P-pipette. The amount of cross-linking agent added was determined to provide a release rate profile. Typically, 5 drops of cross-linking agent solution were added to each 5 g of silicone-xylene mixture. The solution was stirred using a stirring rod until the suspension became a homogeneous milky form. The coating solution was then transferred into a paint jar under conditions for airbrush application.<u style="single">General production of dexamethasone undercoat composition</u> Silicone (35% solution as above) was weighed into a beaker on a Metler scale. The weight of the dexamethasone free alcohol or acetate form was calculated from the weight of the silicone multiplied by 0.35 and the desired proportion of dexamethasone (1-40%) and weighed the required amount. Example: W_silicone= 5g; W for 10% dexamethasone coating_dex= 5x0.35x0.1 / (0.9) = 0.194g, and the required amount of THF in this coating solution was calculated. For 10% dexamethasone coating solution, W_silicone_solid/ V_THF= 0.06. Example: W_silicone= 5g; V_THF= 5x0.35 / 0.06 = 29ml. Dexamethasone was weighed into a beaker on an analytical scale and 1/2 of the total amount of THF was added. The solution was stirred well to ensure complete dissolution of dexamethasone. Next, the stirred DEX-THF solution was transferred to a silicone container. The beaker was washed with the remaining THF and transferred to a silicone container. The cross-linking agent solution was added using a Pasteur P-pipette. Typically, 5 drops of crosslinker solution was used for each 5 g of silicone.

The application of the coating material to the stent was the same for all coating materials, as was for the heparin suspension and dexamethasone suspension prepared as in the above example. The suspension to be coated was transferred to the coating device at 16 in FIG. Typically, a paint jar attached to an airbrush, such as the Badger Model 150, supplied with a compressed air source from a regulator (Norgren, 0-160 psi) was used. The brush hose was attached to a compressed air source downstream of the regulator before air was supplied. The pressure was adjusted to about 15-25 psi and the nozzle condition was checked by pushing down on the trigger.

Any suitable method could be used to fix the stent for spraying, and in the laboratory it was successful with a rotating fixture. Both ends of the relevant stent were secured to the fixture with two elastic retainers, usually alligator clips, and the distance between the clips was adjusted so that the stent remained in a relaxed, non-extended state. The rotor was then powered to adjust the rotational speed to give the desired coating speed, i.e. about 40 rpm.

Adjust the spray nozzle so that the distance from the stent to the stent is approximately 2-4 inches, rotating the stent in a substantially horizontal plane, and brush the brush along the stent from the distal end of the stent. Move to the proximal end, then from the proximal end to the distal end, with a sweeping motion at such a rate that one spray cycle occurs during approximately three stent rotations. It was. Typically, less than 1 minute, usually 1/2 minute, was provided between the layers. Of course, the number of coating layers has changed and will continue to change depending on the particular application. For example, in a typical tie-layer as shown in 18 of FIG. 1, the protruding surface is 1 cm.²For an application level of 3-4 mg of heparin per application, 20 cycles of coating application is required and about 30 ml of solution is consumed for a 3.5 mm diameter x 14.5 cm long stent.

Since the viscosity of the composition and the flow rate of the spray nozzle can be adjusted as needed to modify the layered structure, of course, the rotational speed of the motor can also be adjusted. In general, for the above mix, best results are obtained with a rotation speed in the range of 30-50 rpm and a spray nozzle flow rate in the range of 4-10 ml of coating composition per minute, depending on the size of the stent. Has been done. It is conceivable that more sophisticated computer-controlled coating devices will successfully automate processes that have proven to be feasible in the laboratory.

Several coating layers form what is called an undercoat, as in 18. In one process, an additional top undercoat, which can be of the same composition or different composition with respect to the bioactive material, the matrix polymer material and the cross-linking agent, can be applied as an top layer, eg, in 20. The application of the top layer follows the same application method as the undercoat, but the number and thickness of the layers are arbitrary. Of course, the thickness of any layer can be adjusted by adjusting the rotational speed of the stent and the spraying conditions. In general, the total coating thickness is controlled by the number of spray cycles or the number of thin coats that make up the entire coat.

As shown in 22 of FIG. 1, the applied stent is subsequently subjected to a curing step, in which the prepolymer and the cross-linking agent are coordinated and the cured polymer matrix contains a biologically active substance. To generate. The curing process involves evaporation of the solvent xylene, THF, etc., and curing and cross-linking of the polymer. Certain silicone materials can be cured at relatively low temperatures (ie, room temperature to 50 ° C) in a process called the room temperature vulcanization (RTV) process. However, more typically, the curing process involves a hot curing substance and the applied stent is placed in an oven at about 90 ° C or higher for about 16 hours. For dexamethasone-containing coated stents, the temperature can be raised to as high as 150 ° C. Of course, the time and temperature can be varied depending on the particular silicone, cross-linking agent and biologically active substance.

Stents applied and cured as described above must be sterilized before being packaged for future transplantation. For sterilization, gamma rays are the preferred method, especially for heparin-containing coatings; however, when γ-ray sterilization is applied to a stent that has been applied and cured according to the methods of the invention, these are applied and cured. Unless the stents are first pretreated as in 24, these stents are very likely to restore their original posture when catheterized into blood vessels or other lumens. It can be slow.

The pretreatment step involves argon plasma treatment of the stent, which has been applied and cured in an unconstrained form. According to this method, the stent is placed in a room of a plasma surface treatment system such as Plasma Science 350 (Himont / Plasma Science, Foster City, CA). The system includes a reaction chamber, a radio frequency (RF) solid state generator operating at 13.56 mHz and an output of 0-500 watts with a microprocessor control system, and a complete vacuum pump package. .. The reaction chamber has an undisturbed work volume of 16.75 inches (42.55 cm) x 13.5 inches (34.3 cm) x 17.5 inches (44.45 cm) depth.

In the plasma process, an unconstrained applied stent is placed in the reaction chamber, the system is purged with nitrogen and a vacuum is applied up to 20-50 mTorr. The inert gas (argon, helium, or a mixture thereof) is then placed in the reaction chamber for plasma treatment. A highly preferred method of operation consists of the use of argon gas and an operation with an output range of 200-400 watts, a flow rate of 150-650 standard ml / min equivalent to 100-450 mTorr and an exposure time of 30 seconds to about 5 minutes. .. The stent can be removed immediately after plasma treatment or kept in an argon atmosphere for an additional period of time, typically 5 minutes.

After this, the stent can be exposed to gamma sterilization at 2.5-3.5 Mrad, as shown in 26. Radiation can be performed on a stent that is not radially constrained or that is radially constrained.

However, it is preferable to modify the surface by one of several additional treatment methods prior to plasma treatment or just prior to the sterilization method, some of which are related to the following examples. explain.

Example 1.<u style="single">Fluorosilicone surface treatment that elutes the heparin coating</u> The stent undercoat was applied as a multi-coating layer as described above and then cured as described in 22. The heparin content of this undercoat was 37.5% and the coating thickness was about 30-40μ. Weigh a certain amount of fluorosilicone suspension from the fluorosilicone suspension (Applied Silicone # 40032) at 30 and the relational expression: V_THFA fluorosilicone (FSi) spray solution was prepared by adding tetrahydrofuran (THF) according to the weight of the = 1.2x fluorosilicone suspension. The solution was stirred well enough and sprayed onto the stent at 32 using the method of application of the undercoat process at 18, and the applied stent was cured at 90 ° C for 16 hours. The applied stent is treated with argon plasma according to steps 22 to 26 by the method described above before the above-mentioned γ-ray sterilization.

FIG. 7 is a plot of heparin release rates in a phosphate buffer system with and without a fluorosilicone topcoat. The thickness of the top coat is about 10-15μ. Although not shown in the graph of Figure 7, it should be noted that the release rate of the coating without FSi is initially about 25 times greater than that of the coating with FSi. This, of course, clearly goes beyond the scale of the graph. However, it is noteworthy that coatings with an FSi top layer or diffusion barrier show a decrease in initial release rate and an increase in elution rate from day 1 to day 10 throughout the first week. In addition, fluorosilicone (FSi) topcoats maintain non-thrombogenic surface quality during and after elution of biologically active substances due to the high electronegativity of fluorination. Moreover, due to the negative charge of heparin itself, the electronegativity of fluorosilicone topcoats is at least a partial cause of changes in the heparin release rate profile.

FIG. 8 is a plot of a fluorosilicone (FSi) topcoat containing 16.7% embedded heparin and a topcoat containing only fluorosilicone (FSi). The undercoat is the same as the undercoat used in FIG. 7, contains about 37.5% heparin and is about 37-40 μ thick. These elution rates are quite comparable to the heparin-free FSi upper layer, which significantly reduces the initial burst of heparin release, or else heparin in the FSi upper layer produces a slightly higher release over the test period. Shown. Example 2.<u style="single">Solid polyethylene glycol (PEG) on drug elution undercoat</u><u style="single">Samadhi</u> An undercoat was applied onto the stent and cured in 22 as in Example 1. The stent was then treated with argon gas plasma as in 24 and with ammonium gas plasma at 40. The apparatus and process for the argon gas plasma treatment were as described above. Ammonium plasma treatment was performed immediately after argon gas plasma treatment to aminate the surface of the coating. Ammonium flow rate is 100-700 cm³It was in the range of / minute (ccM), preferably in the range of 500 to 600 ccM. The output of the radio frequency plasma was in the range of 50 to 500 watts, preferably ~ 200 watts. The process time is in the range of 30 seconds to 10 minutes, preferably ~ 5 minutes.

Immediately after amination, the stent was immersed in electrophilically activated polyethylene glycol (PEG) solution at 42. PEG is known to be an inhibitor of protein absorption. Examples of electrophilically activated polyethylene glycol are PEG nitrophenyl carbonates, PEG trichlorophenyl carbonates, PEG tresylates, PEG glycidyl ethers, PEG isocyanates, etc., and one end is optionally terminated with a methoxy group. The molecular weight of PEG ranges from about 1000 to 6000, preferably about 3000. Simple ammonium amination does not produce large amounts of primary and secondary amines on the surface of the elastomeric polymer (eg silicone). Instead, imines (&gt; C = NH) and other more oxidizing nitro-containing groups dominate the surface. For example, NaBH in the reaction medium so that the functional groups on the PEG can react with the imine and possibly other nitro-containing species on the surface to immobilize the PEG on the surface.₃It is generally necessary to add a reducing agent such as CN. NaBH₃A typical concentration of CN is about 2 mg / ml. Since PEG and its derivatives dissolve in water and many polar aromatic solvents, the solvent used for coating must be a PEG solvent, but the drug in the undercoat to prevent possible loss of the drug due to leaching. Must not be the solvent for. In the case of an eluting-heparin coating, a mixed solvent of formamide and methyl ethyl ketone or a mixed solvent of form amide and acetone (preferably in a volume ratio of formamide 30: methyl ethyl ketone or acetone 70) dissolves heparin. It is a preferred solvent because it does not. The PEG concentration, reaction time, reaction temperature and pH value depend on the type of PEG used. In the case of heparin-eluting coatings, 5% PEG tresylate in (30/70) formamide / methyl ethyl ketone has been used successfully. The reaction time was 3 hours at room temperature. Next, PEG was covalently bonded to the surface. Next, gamma rays were used for sterilization of this embodiment as described above.

With respect to the anticoagulant heparin, the proportion in the undercoat is nominally about 30-50% and the proportion in the topcoat is about 0-30% active substance. The coating thickness ratio of topcoat to undercoat varies from about 1:10 to 1: 2, preferably in the range of about 1: 6 to 1: 3.

During the stending process, the doctor administers a bolus injection of antiplatelet / anticoagulant drug to the patient, so suppressing the burst effect can also reduce drug loading, in other words, reduce the thickness of the coating layer. To. As a result, the drug encapsulated in the stent can be completely used without waste. Regulating day 1 release and maximizing day 2 and day 3 release in the thinnest possible coating shape reduces acute or subacute thrombi.

Figure 4 shows the general effect of drug loading on coatings of the same thickness. The initial elution rate increases with the amount of drug carried, as shown in FIG. The release rate also increases with the coating thickness at the same loading, but tends to be inversely proportional to the topcoat thickness, as indicated by the same undercoat thickness as in FIG.

However, what is clear from the data collected to date is that the processes of the present invention allow the drug elution rate to be controlled in a desirable manner to meet the requirements of a particular stent application. In a similar manner, it is possible to produce a stent coating using a combination of two or more drugs, controlling the drug release sequence and rate. For example, multiple antiproliferative agents can be combined in the undercoat and multiple antiplatelet agents can be combined in the topcoat. In this way, antiplatelet agents, such as heparin, can be eluted first, then antiproliferative agents, to better enable safe encapsulation of the implanted stent.

The heparin concentration was measured using a standard curve created by complexing the Azure A dye with a dilute solution of heparin. 16 standards were used to compile the standard curve in a well-known manner.

For elution tests, the stent was immersed in a pH 7.4 phosphate buffer in an incubator at about 37 ° C. The heparin elution amount was measured by periodically sampling this solution. After each sampling, each stent was placed in a heparin-free buffer solution.

As mentioned above, the acceptable heparin support of the elastomeric material varies, but in the case of silicone materials, heparin can exceed 60% of the total weight of the layer. However, the most commonly used support is generally in the range of about 10% to 45% of the total weight of the layer. In the case of dexamethasone, the support can be as much as 50% or more of the total weight of the layer, but preferably in the range of about 0.4% to 45%.

It will be appreciated that the mechanism of incorporation of the biologically active substance into the thin surface coating structure applicable to metal stents is an important aspect of the present invention. The need for relatively thick polymer elution stents or optional membrane overlays associated with prior art drug elution devices is the need for the use of biodegradable or reabsible vehicles for the transport of biologically active substances. Not as well as. This method clearly allows long-term administration and minimizes interference with the independent mechanical or therapeutic benefits of the stent itself.

The coating material is designed by a specific application method, coating / drug combination and drug infusion mechanism. Consideration of the specific shape and release mechanism of the biologically active substance in the coating yields very good results for this method. In this way, the administration of the biologically active substance from the coating structure can be adjusted to suit a variety of uses.

The above example shows two different drug carriers or proportions of biologically active material to be released, but this is by no means limited to the present invention and with any number of layers to obtain the desired release profile. It is intended that a combination of supports can be used. For example, gradual grading and variation of layer support can be utilized, for example, high support can be used for thin layers. A layer having no drug support can also be used. For example, a pulsatile heparin-releasing system can be obtained by coating with alternating layers containing heparin between non-supporting layers of silicone, or by using other substances as part of the coating. In other words, the present invention allows for countless combinations, resulting in tremendous flexibility in controlling the release of biologically active substances with respect to the implanted stent. Each coating layer is typically about 0.5 to 15 microns thick. The total number of spray-coated layers can, of course, vary widely, from less than 10 layers to more than 50 layers, generally including 20-40 layers. The overall thickness of the coating can also vary widely, but can generally be about 10-200 microns.

The polymer of the coating is any compatible biostable elastomeric material that can adhere to the stent material as a thin layer, but the release of biologically active material is generally predictable by hydrophobic materials. Hydrophobic substances are preferred as they have been found to be controllable. Preferred substances are particularly preferably silicone rubber elastomers and biologically stable polyurethanes.

The present invention has been described in considerable detail in accordance with patent law, applying novel principles to those skilled in the art, and providing the information necessary to construct and use embodiments of the Examples as needed. However, it should be understood that the present invention can be carried out specifically by different devices and that various modifications can be made without departing from the scope of the present invention itself.

Embodiments of the present invention are as follows. 1. A implantable medical device having an outer surface, wherein at least a portion of the outer surface is coated with a shape-matched coating of a hydrophobic elastomeric material, wherein the hydrophobic elastomeric material is said to be hydrophobic. It contains a predetermined amount of biologically active material within it for time-controlled delivery from the Elastomer material. A implantable medical device having a means for providing a non-thrombogenic surface after time-controlled delivery of a biologically active substance, which is associated with a shape-matched coating. 2. The medical device according to 1 above, wherein the shape-matched coating comprises a predetermined amount of a biologically active substance subdivided into a hydrophobic elastomeric material. 3. The medical device according to 2 above, wherein the finely divided biologically active substance has an average particle size of less than about 15 microns. 4. The above 2 above, wherein the finely divided biologically active material has an average particle size of less than about 10 microns and carries about 25-60% by weight of the drug in a shape-matched coating. Medical equipment. 5. The medical device according to 1 above, wherein the shape-matched coating comprises a predetermined amount of a biologically active substance dispersed in an elastomeric material at the molecular level. 6. Biologically active substances are antithrombotic agents, anticoagulants, antiplatelet agents, thrombolytic agents, antiproliferative agents, steroidal anti-inflammatory agents, non-steroidal anti-inflammatory agents, hyperplasia, especially , Re-stenosis inhibitor, smooth muscle cell inhibitor, growth factor, growth factor inhibitor, cell adhesion inhibitor, cell adhesion promoter, healthy neointimal tissue formation promoter, and a combination thereof. The medical device described in 1 above to be selected. 7. The medical device according to 1 above, wherein the hydrophobic elastomer material is selected from the group consisting of silicone, polyurethane, ethylene-vinyl acetate copolymer, polyolefin elastomer, polyamide elastomer, EPDM rubber, and a combination thereof. 8. The medical device according to 1 above, wherein the shape-matched coating comprises a plurality of layers of elastomeric material, wherein the elastomeric material comprises a predetermined amount of a biologically active substance therein. 9. The means associated with the conforming coating have an outer layer, the outer layer at least partially covering the conforming coating and comprising a non-thrombogenic polymer material in 1 above. The listed medical device. Ten. 9. The medical device according to 9 above, wherein the non-thrombotic polymer material is selected from the group consisting of fluorosilicone, polyethylene glycol, polysaccharides, phospholipids, and combinations thereof. 11. The medical device according to 10 above, wherein the non-thrombogenic polymer material comprises a fluorosilicone coating adhered to a shape-matched coating. 12. The medical device according to 10 above, wherein the non-thrombotic polymer material comprises polyethylene glycol covalently attached to a coating to which the aminated shape fits. 13. The medical device according to 1 above, in which the transplantable medical device is made of metal. 14. Transplantable to a human body having a tubular metal body with an open edge and a side wall structure consisting of a grid with gaps; and a continuous shape-matched coating provided on the surface of the side wall structure. An expandable stent in which the shape-fitting coating comprises a hydrophobic elastomeric material so that the hydrophobic elastomeric material provides a time-controlled delivery from the hydrophobic elastomeric material. Containing a predetermined amount of active material within it, the coating adapts to the sidewall structure so as to hold a grid with gaps, the coating is a stent with a non-thrombotic outer surface. .. 15. The stent according to 14 above, wherein the biologically active substance is heparin. 16. The stent according to 15 above, wherein the outer layer of the coating comprises a material selected from the group consisting of fluorosilicone and polyethylene glycol (PEG). 17. The stent according to 16 above, wherein the outer layer is fluorosilicone. 18. The stent according to 17 above, wherein the outer layer further comprises a predetermined amount of finely divided heparin, if desired. 19. The stent according to 16 above, wherein the outer layer contains polyethylene glycol. 20. With a tubular metal body having an open end and a side wall; with a continuous shape-fitting coating on the surface of the side wall; The coating is further provided with an undercoat made of a hydrophobic elastomeric material, the hydrophobic elastomeric material being time controlled from the hydrophobic elastomeric material. A stent in which a predetermined amount of finely divided heparin is contained therein for delivery, the coating further comprising a topcoat containing a predetermined amount of fluorosilicone. 21. The stent according to 20 above, wherein the topcoat further contains a predetermined amount of finely divided heparin, if desired. 22. A stent for implantation into the lumen of a blood vessel having a tubular metal body with an open end and a side wall; and a continuous shape-matched coating provided on the surface of the side wall; The coating has an undercoat made of a hydrophobic elastomeric material, wherein the hydrophobic elastomeric material contains a predetermined amount of finely divided heparin for time-controlled delivery from the hydrophobic elastomeric material. The coating is a stent having a topcoat containing a predetermined amount of polyethylene glycol. 23. A method of coating a implantable stent prosthesis with a layer, wherein the layer comprises a hydrophobic elastomeric material so that the hydrophobic elastomeric material provides time-controlled delivery from the hydrophobic elastomeric material. In addition, a predetermined amount of a biologically active substance is contained therein, and (a) a step of applying a topcoat, wherein the topcoat is an uncured polymeric material in a solvent mixture. And a composition containing a predetermined amount of finely divided biologically active substances; (b) a step of curing the polymeric material; (c) a step of applying a topcoat, and here. In, the topcoat is composed of a composition having properties, the properties of which are to impart a non-thrombotic surface and allow modified delivery of biologically active substances; Method. twenty four. The method according to 23 above, wherein the elastomeric material is silicone and the biologically active substance is heparin. 25. The method of 24 above, wherein the topcoat comprises fluorosilicone. 26. The method of 25 above, wherein the topcoat comprises a predetermined amount of heparin. 27. The method of 24 above, wherein the topcoat comprises polyethylene glycol. 28. (d) The cured topcoat is treated with an inert gas plasma and then with ammonia plasma; (e) From a solution of polyethylene glycol (PEG), an outer coating of polyethylene glycol is applied. 27. The method of 27 above, further comprising a step of coating and; 29. The method according to 28 above, wherein the polyethylene glycol (PEG) in the solution was selected from PEG nitrophenyl carbonates, PEG trichlorophenyl carbonates, PEG tresylate, PEG glycidyl ether, PEG isocyanate and combinations thereof. 30. The method according to 28 above, wherein the polyethylene glycol is electrophilically active. 31. The method according to 28 above, wherein the polyethylene glycol has a methoxy group terminal. 28. The method according to 28 above, wherein the polyethylene glycol (PEG) in the solution is selected from PEG nitrophenyl carbonates, PEG trichlorophenyl carbonates, PEG tresylate, PEG glycidyl ether, PEG isocyanate and combinations thereof. 30. The method according to 28 above, wherein the polyethylene glycol is electrophilically active. 31. The method according to 28 above, wherein the polyethylene glycol has a methoxy group terminal. 28. The method according to 28 above, wherein the polyethylene glycol (PEG) in the solution is selected from PEG nitrophenyl carbonates, PEG trichlorophenyl carbonates, PEG tresylate, PEG glycidyl ether, PEG isocyanate and combinations thereof. 30. The method according to 28 above, wherein the polyethylene glycol is electrophilically active. 31. The method according to 28 above, wherein the polyethylene glycol has a methoxy group terminal.

In the figure below, the approximated citation number indicates the approximated part.<figref num="1">It is a process drawing which shows the process of the method of this invention.</figref><figref num="2">The release profile in a multi-layer system is shown, and the percentage of heparin released over a 2-week period is shown. [ Bonding layer = 37.5% heparin coating, top layer = silicone] [ Bonding layer = 37.5% heparin coating, top layer = 16.7% heparin coating] [ Single layer = 37.5 heparin coating]</figref><figref num="3">It shows the release profile in a multi-layer system and shows the relative release rate of heparin over a 2-week period. [ Bonding layer = 37.5% heparin coating, upper layer = silicone] [ Bonding layer = 37.5% heparin coating, upper layer = 16.7% heparin coating]</figref><figref num="4">It shows the kinetic profiles of different drug charges at similar coating thicknesses and shows the release of heparin over a 2-week period. No related means is provided to provide a non-thrombogenic surface for a long period of time.</figref><figref num="5">We show drug elution kinetics over a two-week period when the coating thickness is different and the heparin filling is constant. No related means is provided to provide a non-thrombogenic surface for a long period of time. Here, the filling of the drug, ie heparin, is 41.1%.</figref><figref num="6">The release kinetics when the composition of the undercoat and the topcoat is constant and the thickness of the undercoat is different is shown. The proportion of heparin was kept constant in each of the undercoat and the topcoat.</figref><figref num="7">It is a plot of heparin release kinetics in a phosphate buffer system at pH 7.4. The presence or absence of a fluorosilicone (FSi) topcoat is different. [ With fluorosilicone top layer] [ Without fluorosilicone top layer] In the first 2 hours, the release rate of the coating without fluorosilicone is 25 times faster than with fluorosilicone.</figref><figref num="8">It is a plot of heparin release kinetics in a phosphate buffer system at pH 7.4. Topcoats containing only fluorosilicone (FSi) are compared with FSi topcoats with 16.7% heparin embedded. [ No heparin in the fluorosilicone top layer] [ 16.7% heparin in the fluorosilicone top layer] The thickness of the 37.5% heparin binding coat is about 40 microns.</figref>