a type of stent device of one class designed to be utilized in combination with coatings in the present invention is shown diagrammatically in a side view and an end view , respectively contained in fig1 a and 1b . fig1 a shows a section of a generally cylindrical tubular body 10 having a mantle surface formed by a number of individual thread elements 12 , 14 and 13 , 15 , etc . of these elements , elements 12 , 14 , etc . extend generally in an helix configuration axially displaced in relation to each other but having center line 16 of the body 10 as a common axis . the other elements 13 , 15 , likewise axially displaced , extend in helix configuration in the opposite direction , the elements extending in the two directions crossing each other in the manner indicated in fig1 a . a tubular member so concerned and so constructed can be designed to be any convenient diameter , it being remembered that the larger the desired diameter , the larger the number of filaments of a given wire diameter ( gauge ) having common composition and prior treatment required to produce a given radial compliance . the braided structure further characteristically readily elongates upon application of tension to the ends axially displacing them relative to each other along center line 16 and correspondingly reducing the diameter of the device . this is illustrated in fig2 a and 2b in which a segment of the device 10 of fig1 a and 1b has been elongated by moving the ends 18 and 20 away from each other in the direction of the arrows . upon the release of the tension on the ends , the structure 10 , if otherwise unrestricted , will reassume the relaxed or unloaded configuration of fig1 a and 1b . the elongation / resumption characteristic flexibility of the stent device enables it to be slipped or threaded over a carrying device while elongated for transportation through the vascular or other relevant internal luminal system of a patient to the site of interest where it can be axially compressed and thereby released from the carrying mechanism , often a vascular catheter device . at the site of interest , it assumes an expanded condition held in place by mechanical / frictional pressure between the stent and the lumen wall against which it expands . the elongation , loading , transport and deployment of such stents is well known and need not be further detailed here . it is important , however , to note that when one contemplates coatings for such a stent in the manner of the present invention , an important consideration resides in the need to utilize a coating material having elastic properties compatible with the elastic deforming properties residing in the stent that it coats . the material of the stent should be rigid and elastic but not plastically deformable as used . as stated above , the preferred materials for fabricating the metallic braided stent include stainless steel , tantalum , titanium alloys including nitinol and certain cobalt - chromium alloys . the diameter of the filaments may vary but for vascular devices , up to about 10 mm in diameter is preferable with the range 0 . 01 to 0 . 05 mm . drug release surface coatings on stents in accordance with the present invention can release drugs over a period of time from days to months and can be used , for example , to inhibit thrombus formation , inhibit smooth muscle cell migration and proliferation , inhibit hyperplasia and restenosis , and encourage the formation of health neointimal tissue including endothelial cell regeneration . as such , they can be used for chronic patency after an angioplasty or stent placement . it is further anticipated that the need for a second angioplasty procedure may be obviated in a significant percentage of patients in which a repeat procedure would otherwise be necessary . a major obstacle to the success of the implant of such stents , of course , has been the occurrence of thrombosis in certain arterial applications such as in coronary stenting . of course , antiproliferative applications would include not only cardiovascular but any tubular vessel that stents are placed including urologic , pulmonary and gastro - intestinal . various combinations of polymer coating materials can be coordinated with the braided stent and the biologically active agent of interest to produce a combination which is compatible at the implant site of interest and controls the release of the biologically active species over a desired time period . preferred coating polymers include silicones ( poly siloxanes ), polyurethanes , thermoplastic elastomers in general , ethylene vinyl acetate copolymers , polyolefin rubbers , epdm rubbers , and combinations thereof . specific embodiments of the present invention include those designed to elute heparin to prevent thrombosis over a period of weeks or months or to allow the diffusion or transport of dexamethasone to inhibit fibromuscular proliferation over a like period of time . of course , other therapeutic substances and combinations of substances are also contemplated . the invention may be implanted in a mammalian system , such as in a human body . the heparin elution system is preferably fabricated by taking finely ground heparin crystal , preferably ground to an average particle size of less than 10 microns , and blending it into a liquid , uncured poly siloxane / solvent material in which the blend ( poly siloxane plus heparin ) contains from less than 10 % to as high as 80 % heparin by weight with respect to the total weight of the material and typically the layer is between 10 % and 45 % heparin . this material is diluted with a solvent and utilized to coat a metallic braided stent , which may be braided cobalt chromium alloy wire , in a manner which applies a thin , uniform coating ( typically between 20 and 200 microns in thickness ) of the heparin / polymer mixture on the surfaces of the stent . the polymer is then heat cured , or cured using low temperature thermal initiators (& lt ; 100 ° c .) in a room temperature vulcanization ( rtv ) process in situ on the stent to evaporate the solvent , typically tetrahydrofuran ( tef ). the heparin forms interparticle paths in the silicone sufficiently interconnected to allow slow but substantially complete subsequent elution . the ultrafine particle size utilized allows the average pore size to be very small such that elution may take place over weeks or even months . a coating containing dexamethasone is produced in a somewhat different manner . a poly siloxane material is also the preferred polymeric material . nominally an amount equal to 0 . 4 % to about 45 % of the total weight of the layer of dexamethasone is used . the dexamethasone drug is dissolved in a solvent , e . g ., thf first . the solution is then blended into liquid uncured poly siloxane / solvent ( xylene , thf , etc .) vehicle precursor material . since the dexamethasone is also soluble in the solvent for the polysiloxane , it dissolves into the mixture . the coating is then applied to the stent and upon application , curing and drying , including evaporation of the solvent , the dexamethasone remains dispersed in the coating layer . it is believed that the coating is somewhat in the nature of a solid solution of recrystallized particles of dexamethasone in silicone rubber . dexamethasone , as a rather small molecule , however , does not need gross pores to elute and may be transported or diffused outward through the silicone material over time to deliver its anti - inflammatory medicinal effects . the coatings can be applied by dip coating or spray coating or even , in some cases , by the melting of a powdered form in situ or any other technique to which the particular polymer / biologically active agent combination is well suited . it will be understood that a particularly important aspect of the present invention resides in the technology directed to the incorporation of very fine microparticles or colloidal suspensions of the drug into the polymer matrix . in the case of a crystalline drug , such as heparin , the drug release is controlled by the network the drug forms in the polymer matrix , the average particulate size controlling the porosity and so the ultimate elution rate . fig4 a depicts a stent which has been spray coated with a solvent containing a cured polysilicone material including an amount of heparin crystals to provide a thin , uniform coating on all surfaces of the stent . the coated stent was cured at 150 ° c . for 18 minutes ; the sample was eluted in pbs for 49 days at 37 ° c . and the stent was rinsed in ethanol prior to taking the scanning electron microscope picture of fig4 a . fig4 b shows a greatly enlarged ( 600 ×) scanning electron microscope photograph ( sem ) of a portion of the coating of fig4 a in which the microporosity is evident . the coating thickness may vary but is typically from about 75 to about 200 microns . fig5 a and 5b show scanning electron microscope photographs of a heparin containing polysiloxane stent . the figure shows the coating prior to elution of the heparin . the coating was cured at 150 for 18 minutes . fig5 b is greatly enlarged photograph ( sex ) of a fragment of the coated surface of fig5 a showing the substantially non - porous surface prior to elution . fig6 a and 6b show the posture of a stent in accordance with the invention as implanted in a swine coronary . the blemish shown in fig6 a represents a histological artifact of unknown origin . as can be seen in fig6 b , a large number of heparin particles are contained in the silicone material . the substantially non - porous surface of fig7 a typically occurs with an incorporation of an amount of non - particulate material such as dexamethasone which partially or entirely dissolves in the solvent for the poly siloxane prior to coating and cure . upon curing of the polymer and evaporation of the solvent , depending on the loading of dexamethasone , the dexamethasone reprecipitates in a hydrophobic crystalline form containing dendrite or even elongated hexagonal crystals approximately 5 microns in size . as can be seen in fig7 b , even after release of the incorporated material or three months , the coating surface remains substantially non - porous indicating the transport or diffusion of the drug outward through the silicone material neither requires nor produces gross pores . the dexamethasone is incorporated in its more hydrophobic form rather than in one of the relatively more hydrophilic salt forms such as in a phosphate salt , for example . fig8 and 9 depict plots of total percent drug release related to long - term drug release stent coating layers . fig8 depicts the release of heparin from a 50 % heparin loading in silicone . the silicone was cured at 90 ° c . for 16 hours . the heparin release took place in a phosphoric buffer ( ph = 7 . 4 ) for 90 days at 37 ° c . the heparin concentration in the phosphoric buffer was analyzed by azure a assay . fig9 depicts a graphical analysis , similar to that depicted for heparin in fig8 , for the release of dexamethasone at two different concentrations , i . e ., 5 % and 10 % in silicone polymer . the coated stents were cured at 150 ° c . for 20 minutes and the release took place in a polyethylene glycol ( peg ), mw = 400 / water solution at 37 ° c . (( peg 400 / h 2 o ) ( 40 / 60 , vol / vol )). the dexamethasone concentrations were analyzed photometrically at 241 μm . fig8 and 9 illustrate possible stent coating layers of polymer / bioactive species combinations for long - term release . as stated above , the release rate profile can be altered by varying the amount of active material , the coating thickness , the radial distribution of bioactive materials , the mixing method , and the crosslink density of the polymer matrix . sufficient variation is possible such that almost any reasonable desired profile can be simulated . according to the present invention , the stent coatings incorporating biologically active materials for timed delivery in situ in a body lumen of interest are preferably sprayed in many thin layers from prepared coating solutions or suspensions . the steps of the process are illustrated generally in fig1 . the coating solutions or suspensions are prepared at 10 as will be described later . the desired amount of crosslinking agent is added to the suspension / solution as at 12 and material is then agitated or stirred to produce a homogenous coating composition at 14 which is thereafter transferred to an application container or device which may be a container for spray painting at 16 . typical exemplary preparations of coating solutions that were used for heparin and dexamethasone appear next . silicone was obtained as a polymer precursor in solvent ( xylene ) mixture . for example , a 35 % solid silicone weight content in xylene was procured from applied silicone , part # 40 , 000 . first , the silicone - xylene mixture was weighed . the solid silicone content was determined according to the vendor &# 39 ; s analysis . precalculated amounts of finely divided heparin ( 2 - 6 microns ) were added into the silicone , then tetrahydrofuron ( thf ) hpcl grade ( aldrich or em ) was added . for a 37 . 5 % heparin coating , for example : w silicone = 5 g ; solid percent = 35 %; w hep = 5 × 0 . 35 × 0 . 375 /( 0 . 625 )= 1 . 05 g . the amount of thf needed ( 44 ml ) in the coating solution was calculated by using the equation w silicone solid / va thf = 0 . 04 for a 37 . 5 % heparin coating solution ). finally , the manufacturer crosslinker solution was added by using pasteur p - pipet . the amount of crosslinker added was formed to effect the release rate profile . typically , five drops of crosslinker solution were added for each five grams of silicone - xylene mixture . the crosslinker may be any suitable and compatible agent including platinum and peroxide based materials . the solution was stirred by using the stirring rod until the suspension was homogenous and milk - like . the coating solution was then transferred into a paint jar in condition for application by air brush . silicone ( 35 % solution as above ) was weighed into a beaker on a metler balance . the weight of dexamethasone free alcohol or acetate form was calculated by silicone weight multiplied by 0 . 35 and the desired percentage of dexamethasone ( 1 to 40 %) and the required amount was then weighed . example : w silicone = 5 g ; for a 10 % dexamethasone coating , w dex = 5 × 0 . 35 × 0 . 1 / 0 . 9 = 0 . 194 g and thf needed in the coating solution calculated . w silicone solid / v thf = 0 . 06 for a 10 % dexamethasone coating solution . example : w silicone = 5 g ; v thf = 5 × 0 . 35 / 0 . 06 = 29 ml . the dexamethasone was weighed in a beaker on an analytical balance and half the total amount of thf was added . the solution was stirred well to ensure full dissolution of the dexamethasone . the stirred dex - thf solution was then transferred to the silicone container . the beaker was washed with the remaining thf and this was transferred to the silicone container . the crosslinker was added by using a pasteur pipet . typically , five drops of crosslinker were used for five grams of silicone . the application of the coating material to the stent was quite similar for all of the materials and the same for the heparin and dexamethasone suspensions prepared as in the above examples . the suspension to be applied was transferred to an application device , typically a paint jar attached to an air brush , such as a badger model 150 , supplied with a source of pressurized air through a regulator ( norgren , 0 - 160 psi ). once the brush hose was attached to the source of compressed air downstream of the regulator , the air was applied . the pressure was adjusted to approximately 15 - 25 psi and the nozzle condition checked by depressing the trigger . while any appropriate method can be used to secure the stent for spraying , rotating fixtures were utilized successfully in the laboratory . both ends of the relaxed stent were fastened to the fixture by two resilient retainers , commonly alligator clips , with the distance between the clips adjusted so that the stent remained in a relaxed , unstretched condition . the rotor was then energized and the spin speed adjusted to the desired coating speed , nominally about 40 rpm . with the stent rotating in a substantially horizontal plane , the spray nozzle was adjusted so that the distance from the nozzle to the stent was about 2 - 4 inches and the composition was sprayed substantially horizontally with the brush being directed along the stent from the distal end of the stent to the proximal end and then from the proximal end to the distal end in a sweeping motion at a speed such that one spray cycle occurred in about three stent rotations . typically a pause of less than one minute , normally about one - half minute , elapsed between layers . of course , the number of coating layers did and will vary with the particular application . for example , for a coating level of 3 - 4 mg of heparin per cm 2 of projected area , 20 cycles of coating application are required and about 30 ml of solution will be consumed for a 3 . 5 mm diameter by 14 . 5 cm long stent . the rotation speed of the motor , of course , can be adjusted as can the viscosity of the composition and the flow rate of the spray nozzle as desired to modify the layered structure . generally , with the above mixes , the best results have been obtained at rotational speeds in the range of 30 - 50 rpm and with a spray nozzle flow rate in the range of 4 - 10 ml of coating composition per minute , depending on the stent size . it is contemplated that a more sophisticated , computer - controlled coating apparatus will successfully automate the process demonstrated as feasible in the laboratory . several applied layers make up what is called the tie layer as at 18 and thereafter additional upper layers , which may be of a different composition with respect to bioactive material , the matrix polymeric materials and crosslinking agent , for example , are applied as the top layer as at 20 . the application of the top layer follows the same coating procedure as the tie layer with the number and thickness of layers being optional . of course , the thickness of each layer can be adjusted by adjusting the speed of rotation of the stent and the spraying conditions . generally , the total coating thickness is controlled by the number of spraying cycles or thin coats which make up the total coat . as shown at 22 in fig1 , the coated stent is thereafter subjected to a curing step in which the polymer precursor and crosslinking agents cooperate to produce a cured polymer matrix containing the biologically active species . the curing process involves evaporation of the solvent xylene , thf , etc . and the curing and crosslinking of the polymer . certain silicone materials can be cured at relatively low temperatures , ( i . e . rt - 50 ° c .) in what is known as a room temperature vulcanization ( rtv ) process . more typically , however , the curing process involves higher temperature curing materials and the coated stents are put into an oven at approximately 90 ° c . or higher for approximately 16 hours . the temperature may be raised to as high as 150 ° c . for dexamethasone containing coated stents . of course , the time and temperature may vary with particular silicones , crosslinkers biologically active species and coating thicknesses . stents coated and cured in the manner described need to be sterilized prior to packaging for future implantation . for sterilization , gamma radiation is a preferred method particularly for heparin containing coatings ; however , it has been found that stents coated and cured according to the process of the invention subjected to gamma sterilization may be too slow to recover their original posture when delivered to a vascular or other lumen site using a catheter unless a pretreatment step as at 24 is first applied to the coated , cured stent . the pretreatment step involves an argon plasma treatment of the coated , cured stents in the unconstrained configuration . in accordance with this procedure , the stents are placed in — a chamber of a plasma surface treatment system such as a plasma science 350 ( himont / plasma science , foster city , calif .). the system is equipped with a reactor chamber and ri solid - state generator operating at 13 . 56 mhz and from 0 - 500 watts power output and being equipped with a microprocessor controlled system and a complete vacuum pump package . the reaction chamber contains an unimpeded work volume of 16 . 75 inches ( 42 . 55 cit ) by 13 . 5 inches ( 34 . 3 cm ) by 17 . 5 inches ( 44 . 45 cm ) in depth . in the plasma process , unconstrained coated stents are placed in a reactor chamber and the system is purged with nitrogen and a vacuum applied to 20 mtorr . thereafter , inert gas ( argon , helium or mixture of them ) is admitted to the reaction chamber for the plasma treatment . a highly preferred method of operation consists of using argon gas , operating at a power range from 200 to 400 watts , a flow rate of 150 - 650 standard ml per minute , which is equivalent to 100 - 450 mtorr , and an exposure time from 30 seconds to about 5 minutes . the stents can be removed immediately after the plasma treatment or remain in the argon atmosphere for an additional period of time , typically five minutes . after this , as shown at 26 , the stents are exposed to gamma sterilization at 2 . 5 - 3 . 5 mrad . the radiation may be carried out with the stent in either the radially non - constrained status or in the radially constrained status . with respect to the anticoagulant material , heparin , the percentage in the tie layer is nominally from about 30 - 50 % and that of the top layer from about 0 - 30 % active material . the coating thickness ratio of the top layer to the tie layer varies from about 1 : 6 to 1 : 2 and is preferably in the range of from about 1 : 5 to 1 : 3 . suppressing the burst effect also enables a reduction in the drug loading or in other words , allows a reduction in the coating thickness , since the physician will give a bolus injection of antiplatelet / anticoagulation drugs to the patient during the stenting process . as a result , the drug imbedded in the stent can be fully used without waste . tailoring the first day release , but maximizing second day and third day release at the thinnest possible coating configuration will reduce the acute or subcute thrombosis . fig1 depicts the general effect of drug loading for coatings of similar thickness . the initial elution rate increases with the drug loading as shown in fig1 . the release rate also increases with the thickness of the coating at the same loading but tends to be inversely proportional to the thickness of the top layer as shown by the same drug loading and similar tie - coat thickness in fig1 . what is apparent from the data gathered to date , however , is that the process of the present invention enables the drug elution kinetics to be controlled in a manner desired to meet the needs of the particular stent application . in a similar manner , stent coatings can be prepared using a combination of two or more drugs and the drug release sequence and rate controlled . for example , antiproliferation drugs may be combined in the tie layer and antiplatelet drugs in the top layer . in this manner , the antiplatelet drugs , for example , heparin , will elute first followed by antiproliferation drugs to better enable safe encapsulation of the implanted stent . the heparin concentration measurement were made utilizing a standard curve prepared by complexing azure a dye with dilute solutions of heparin . sixteen standards were used to compile the standard curve in a well - known manner . for the elution test , the stents were immersed in a phosphate buffer solution at ph 7 . 4 in an incubator at approximately 37 ° c . periodic samplings of the solution were processed to determine the amount of heparin eluted . after each sampling , each stent was placed in heparin - free buffer solution . as stated above , while the allowable loading of the elastomeric material with heparin may vary , in the case of silicone materials heparin may exceed 60 % of the total weight of the layer . however , the loading generally most advantageously used is in the range from about 10 % to 45 % of the total weight of the layer . in the case of dexamethasone , the loading may be as high as 50 % or more of the total weight of the layer but is preferably in the range of about 0 . 4 % to 45 %. it will be appreciated that the mechanism of incorporation of the biologically active species into a thin surface coating structure applicable to a metal stent is an important aspect of the present invention . the need for relatively thick - walled polymer elution stents or any membrane overlayers associated with many prior drug elution devices is obviated , as is the need for utilizing biodegradable or reabsorbable vehicles for carrying the biologically active species . the technique clearly enables long - term delivery and minimizes interference with the independent mechanical or therapeutic benefits of the stent itself . coating materials are designed with a particular coating technique , coating / drug combination and drug infusion mechanism in mind . consideration of the particular form and mechanism of release of the biologically active species in the coating allow the technique to produce superior results . in this manner , delivery of the biologically active species from the coating structure can be tailored to accommodate a variety of applications . whereas the above examples depict coatings having two different drug loadings or percentages of biologically active material to be released , this is by no means limiting with respect to the invention and it is contemplated that any number of layers and combinations of loadings can be employed to achieve a desired release profile . for example , gradual grading and change in the loading of the layers can be utilized in which , for example , higher loadings are used in the inner layers . also layers can be used which have elutable compounds but no drug loadings at all . for example , a pulsatile heparin release system may be achieved by a coating in which alternate layers containing heparin are sandwiched between unloaded layers of silicone or other materials for a portion of the coating . in other words , the invention allows untold numbers of combinations which result in a great deal of flexibility with respect to controlling the release of biologically active materials with regard to an implanted stent . each applied layer is typically from approximately 0 . 5 microns to 15 microns in thickness . the total number of sprayed layers , of course , can vary widely , from less than 10 to more than 50 layers ; commonly , 20 to 40 layers are included . the total thickness of the coating can also vary widely , but can generally be from about 10 to 200 microns . whereas the polymer of the coating may be any compatible biostable elastomeric material capable of being adhered to the stent material as a thin layer , hydrophobic materials are preferred because it has been found that the release of the biologically active species can generally be more predictably controlled with such materials . preferred materials include silicone rubber elastomers and biostable polyurethanes specifically . this invention has been described herein in considerable detail in order to comply with the patent statutes and to provide those skilled in the art with the information needed to apply the novel principles and to construct and use embodiments of the example as required . however , it is to be understood that the invention can be carried out by specifically different devices and that various modifications can be accomplished without departing from the scope of the invention itself .
the present invention is directed to an expandable stent for implantation in a patient comprising a tubular metal body having open ends and a sidewall structure having openings therein and a coating disposed on a surface of said sidewall structure , said coating comprising a hydrophobic biostable elastomeric material and a biologically active material , wherein said coating continuously conforms to said structure in a manner that preserves said openings .