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|United States Patent Application
;   et al.
May 19, 2005
Medical device with drug
A method of coating implantable open lattice metallic stent prosthesis is
disclosed which includes sequentially applying a plurality of relatively
thin outer layers of a coating composition comprising a solvent mixture
of uncured polymeric silicone material and crosslinker and finely divided
biologically active species, possibly of controlled average particle
size, to form a coating on each stent surface. The coatings are cured in
situ and the coated, cured prosthesis are sterilized in a step that
includes preferred pretreatment with argon gas plasma and exposure to
gamma radiation electron beam, ethylene oxide, steam.
Ding, Ni; (Plymounth, MN)
; Helmus, Michael N.; (Long Beach, CA)
222 EAST 41ST ST
Boston Scientific Scimed, Inc.
November 16, 2004|
|Current U.S. Class:
||424/423; 427/2.1; 623/1.42 |
|Class at Publication:
||424/423; 427/002.1; 623/001.42 |
||A61F 002/02; B05D 003/00; A61F 002/06|
24. An implantable medical device having an outer surface, covered at
least in part by a conformal coating comprising an undercoat of a
hydrophobic biostable elastomeric material which does not degrade, and
incorporates an amount of a biologically active material therein for
timed delivery therefrom; and a topcoat comprising a polymeric material,
which at least partially covers the undercoat, wherein the undercoat and
the topcoat are of different formulations.
25. The medical device of claim 24, wherein the medical device is a stent
for vascular implantation.
26. The medical device of claim 25, wherein the stent is expandable.
27. The medical device of claim 25, wherein the stent comprises a tubular
body having open ends and an open lattice sidewall structure and wherein
said coating conforms to said sidewall structure in a manner that
preserves said open lattice.
28. The medical device of claim 25, wherein the stent comprises stainless
29. The medical device of claim 24, wherein the undercoat comprises an
ethylene vinyl acetate copolymer.
30. The medical device of claim 24, wherein most or all of the
biologically active material is contained in the undercoat.
31. The medical device of claim 24, wherein the topcoat consists of the
32. The medical device of claim 24, wherein the biologically active
material is an antibiotic.
33. A method of making the medical device of claim 24, wherein said method
comprises the steps of: (a) applying an undercoat of a formulation
containing uncured hydrophobic elastomeric material in solvent mixture
and an amount of the biologically active material that is finely divided;
(b) curing said hydrophobic elastomeric material; and (c) applying a
topcoat of a formulation comprising the polymeric material to form the
34. A method of coating an implantable prosthesis with a first polymeric
material incorporating an amount of biologically active material therein
for timed delivery therefrom comprising the steps of: (a) applying a
first formulation comprising the first polymeric material and the
biologically active material to a surface of the prosthesis to form an
under layer when the biologically active material is particulate; (b)
applying a second formulation comprising a second polymeric material over
the under layer to form a top layer; and (c) curing the first and second
polymeric materials, wherein the first polymeric material is a
hydrophobic elastomeric material, the average particle size of the
biologically active material in the first formulation is less than or
equal to about 15 .mu.m and at least some of the biologically active
material is particulate after curing, whereby the first and second
formulations are applied to the prosthesis in a manner to adheringly
35. The method of claim 34 wherein the elastomeric material is selected
from the group consisting of silicones, polyurethanes, polyamide
elastomers, ethylene vinyl acetate copolymers, polyolefin elastomers,
EPDM rubbers and combinations thereof.
36. The method of claim 34, wherein the first formulation comprises about
25-45 weight percent of the biologically active material includes
37. The method of claim 34, wherein the biologically active material has
an average particle size less than or equal to about 10 .mu.m before
38. The method of claim 34, wherein the biologically active material
39. The method of any one of claim 34, wherein the implantable prosthesis
is an expandable stent having a tubular metal body with a surface of an
open lattice nature having openings therein, and wherein the first and
second formulations are applied with the stent expanded.
40. The method of claim 39 wherein the expandable stent is a
41. The method of claim 34 wherein the second formulation is substantially
free of any biologically active material.
42. The method of claim 34 wherein the second formulation is substantially
free of the biologically active material in the first formulation.
43. The method of claim 34 wherein the under layer and the top layer each
have a thickness and the ratio of the thickness of the top layer to the
thickness of the under layer is from about 1:10 to 1:2.
I. CROSS-REFERENCE TO RELATED APPLICATIONS
 The present application is a Continuation-In-Part of copending
application Ser. No. 08/526,273, filed Sep. 11, 1995, and a
Continuation-In-Part of copending application Ser. No. 08/424,884, filed
Apr. 19, 1995, all portions of the parent applications not contained in
this application being deemed incorporated by reference for any purpose.
Cross-reference is also made to application Ser. No. 08/______, entitled
"DRUG RELEASE STENT COATING AND PROCESS", filed of even date and of
common inventorship and assignee, that is also a Continuation-In-Part of
both above-referenced patent applications. Any portion of that
application that is not contained herein is also deemed to be
incorporated by reference for any purpose.
BACKGROUND OF THE INVENTION
 II. Field of the Invention
 The present invention relates generally to therapeutic expandable
stent prosthesis for implantation in body lumens, e.g., vascular
implantation and, more particularly, to a process for providing biostable
elastomeric coatings on such stents which incorporate biologically active
species having controlled release characteristics directly in the coating
 II. Related Art
 In surgical or other related invasive medicinal procedures, the
insertion and expansion of stent devices in blood vessels, urinary tracts
or other difficult to access places for the purpose of preventing
restenosis, providing vessel or lumen wall support or reinforcement and
for other therapeutic or restorative functions has become a common form
of long-term treatment. Typically, such prosthesis are applied to a
location of interest utilizing a vascular catheter, or similar
transluminal device, to carry the stent to the location of interest where
it is thereafter released to expand or be expanded in situ. These devices
are generally designed as permanent implants which may become
incorporated in the vascular or other tissue which they contact at
 One type of self-expanding stent has a flexible tubular body formed
of several individual flexible thread elements each of which extends in a
helix configuration with the centerline of the body serving as a common
axis. The elements are wound in a common direction, but are displaced
axially relative to each other and meet, under crossing a like number of
elements also so axially displaced, but having the opposite direction of
winding. This configuration provides a resilient braided tubular
structure which assumes stable dimensions upon relaxation. Axial tension
produces elongation and corresponding diameter contraction that allows
the stent to be mounted on a catheter device and conveyed through the
vascular system as a narrow elongated device. Once tension is relaxed in
situ, the device at least substantially reverts to its original shape.
Prosthesis of the class including a braided flexible tubular body are
illustrated and described in U.S. Pat. Nos. 4,655,771 and 4,954,126 to
Wallsten and 5,061,275 to Wallsten et al.
 Implanted stents have also been used to carry medicinal agents,
such as thrombolytic agents. U.S. Pat. No. 5,163,952 to Froix discloses a
thermal memoried expanding plastic stent device which can be formulated
to carry a medicinal agent by utilizing the material of the stent itself
as an inert polymeric drug carrier. Pinchuk, in U.S. Pat. No. 5,092,877,
discloses a stent of a polymeric material which may be employed with a
coating associated with the delivery of drugs. Other patents which are
directed to devices of the class utilizing bio-degradable or bio-sorbable
polymers include Tang et al, U.S. Pat. No. 4,916,193, and MacGregor, U.S.
Pat. No. 4,994,071. Sahatjian in U.S. Pat. No. 5,304,121, discloses a
coating applied to a stent consisting of a hydrogel polymer and a
preselected drug; possible drugs include cell growth inhibitors and
heparin. A further method of making a coated intravascular stent carrying
a therapeutic material in which a polymer coating is dissolved in a
solvent and the therapeutic material dispersed in the solvent and the
solvent thereafter evaporated is described in Berg et al, U.S. Pat. No.
5,464,650, issued Nov. 5, 1995 and corresponding to European patent
application 0 623 354 A1 published 09 Nov. 1994.
 An article by Michael N. Helmus (a co-inventor of the present
invention) entitled "Medical Device Design--A Systems Approach: Central
Venous Catheters", 22nd International Society for the Advancement of
Material and Process Engineering Technical Conference (1990) relates to
polymer/drug/membrane systems for releasing heparin. Those polymer/
drug/membrane systems require two distinct layers to function.
 The above cross-referenced grandparent application supplies an
approach that provides long-term drug release, i.e., over a period of
days or even months, incorporated in a controlled-release system. The
parent application and present invention provide a process for coating
such stents including techniques that enable the initial burst effect of
drug elation to be controlled and the drug release kinetic profile
associated with long-term therapeutic effect to be modified.
 Metal stents of like thickness and weave generally have better
mechanical properties than polymeric stents. Metallic vascular stents
braided of even relatively fine metal filament can provide a large amount
of strength to resist inwardly directed circumferential pressure in blood
vessels. In order for a polymer material to provide comparable strength
characteristics, a much thicker-walled structure or heavier, denser
filament weave is required. This, in turn, reduces the cross-sectional
area available for flow through the stent and/or reduces the relative
amount of open space available in the structure. In addition, when
applicable, it is usually more difficult to load and deliver polymeric
stents using vascular catheter delivery systems.
 It will be noted, however, that while certain types of stents such
as braided metal stents may be superior to others for some applications,
the process of the present invention is not limited in that respect and
may be used to coat a wide variety of devices. The present invention also
applies, for example, to the class of stents that are not self-expanding
including those which can be expanded, for instance, with a balloon.
Polymeric stents, of all kinds can be coated using the process. Thus,
regardless of particular detailed embodiments the use of the invention is
not considered or intended to be limited with respect either to stent
design or materials of construction. Further, the present invention may
be utilized with other types of implant prostheses.
 Accordingly, it is a primary object of the present invention to
provide a coating process for coating a stent to be used as a deployed
stent prosthesis, the coating being capable of long-term delivery of
biologically active materials.
 Another object of the invention is to provide a process for coating
a stent prosthesis using a biostable hydrophobic elastomer in which
biologically active species are incorporated within a cured coating.
 Still another object of the present invention is to provide a
multi-layer coating in which the percentage of active material can vary
from layer to layer.
 A further object of the present invention is to control or modify
aspects of the timed or time variable drug delivery from a stent coating
by controlling average particle size in the biologically active species.
 Other objects and advantages of the present invention will become
apparent to those skilled in the art upon familiarization with the
specification and appended claims.
SUMMARY OF THE INVENTION
 The present invention provides processes for producing a relatively
thin layer of biostable elastomeric material in which an amount of
biologically active material is dispersed as a coating on the surfaces of
a deployable stent prosthesis. The preferred stent to be coated is a
self-expanding, open-ended tubular stent prosthesis. Although other
materials, including polymer materials, can be used, in the preferred
embodiment, the tubular body is formed of an open braid of fine single or
polyfilament metal wire which flexes without collapsing and readily
axially deforms to an elongate shape for transluminal insertion via a
vascular catheter. The stent resiliently attempts to resume predetermined
stable dimensions upon relaxation in situ.
 The coating is preferably applied as a mixture, solution or
suspension of polymeric material and finely divided biologically active
species dispersed in an organic vehicle or a solution or partial solution
of such species in a solvent or vehicle for the polymer and/or
biologically active species. For the purpose of this application, the
term "finally divided" means any type or size of included material from
dissolved molecules through suspensions, colloids and particulate
mixtures. The active material is dispersed in a carrier material which
may be the polymer, a solvent, or both. The coating is preferably applied
as a plurality of relatively thin layers sequentially applied in
relatively rapid sequence and is preferably applied with the stent in a
radially expanded state. In some applications the coating may further be
characterized as a composite initial tie coat or undercoat and a
composite topcoat. The coating thickness ratio of the topcoat to the
undercoat may vary with the desired effect and/or the elution system.
Typically these are of different formulations.
 The coating may be applied by dipping or spraying using evaporative
solvent materials of relatively high vapor pressure to produce the
desired viscosity and quickly establish coating layer thicknesses. The
preferred process is predicated on reciprocally spray coating a rotating
radially expanded stent employing an air brush device. The coating
process enables the material to adherently conform to and cover the
entire surface of the filaments of the open structure of the stent but in
a manner such that the open lattice nature of the structure of the braid
or other pattern is preserved in the coated device.
 The coating is exposed to room temperature ventilation for a
predetermined time (possibly one hour or more) for solvent vehicle
evaporation. Thereafter the polymeric precuser material is cured at room
temperature or elevated temperatures or the solvent evaporated away from
the dissolved polymer as the case may be. Curing is defined as the
process of converting the elastomeric or polymeric material into the
finished or useful state by the application of heat and/or chemical
agents which include physical-chemical charges. Where, for example,
polyurethane thermoplastic elastomers are used, solvent evaporation can
occur at room temperature rendering the polymeric material useful for
controlled drug release without further curing. Non-limiting examples of
curing according to this definition include the application of heat
and/or chemical agents and the evaporation of solvent which may induce
physical and/or chemical changes.
 The ventilation time and temperature for cure are determined by the
particular polymer involved and particular drugs used. For example,
silicone or polysiloxane materials (such as polydimethylsiloxane) have
been used successfully. These materials are applied as pre-polymer in the
coating composition and must thereafter be cured. The preferred species
have a relatively low cure temperatures and are known as a room
temperature vulcanizable (RTV) materials. Some polydimethylsiloxane
materials can be cured, for example, by exposure to air at about
90.degree. C. for a period of time such as 16 hours. A curing step may be
implemented both after application of a certain number of lower undercoat
layers and the topcoat layers or a single curing step used after coating
 The coated stents may thereafter be subjected to a postcure
sterilization process which includes an inert gas plasma treatment, and
then exposure to gamma radiation, electron beam, ethylene oxide (ETO) or
steam sterilization may also be employed.
 In the plasma treatment, unconstrained coated stents are placed in
a reactor chamber and the system is purged with nitrogen and a vacuum
applied to about 20-50 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 about 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
 After the argon plasma pretreatment, the coated and cured stents
are subjected to gamma radiation sterilization nominally at 2.5-3.5 Mrad.
The stents enjoy full resiliency after radiation whether exposed in a
constrained or non-constrained status. It has been found that constrained
stents subjected to gamma sterilization without utilizing the argon
plasma pretreatment lose resiliency and do not recover at a sufficient or
 The elastomeric material that forms a major constituent of the
stent coating should possess certain properties. It is preferably a
suitable hydrophobic biostable elastomeric material which does not
degrade and which minimizes tissue rejection and tissue inflammation and
one which will undergo encapsulation by tissue adjacent to the stent
implantation site. Polymers suitable for such coatings include silicones
(e.g., polysiloxanes and substituted polysiloxanes), polyurethanes
(including polycarbonate urethanes), thermoplastic elastomers in general,
ethylene vinyl acetate copolymers, polyolefin elastomers, EPDM rubbers
and polyamide elastomers. The above-referenced materials are considered
hydrophobic with respect to the contemplated environment of the
 Agents suitable for incorporation-include antithrobotics,
anticoagulants, antiplatelet agents, thrombolytics, antiproliferatives,
antinflammatories, agents that inhibit hyperplasia and in particular
restenosis, smooth muscle cell inhibitors, antibiotics growth factors,
growth factor inhibitors, cell adhesion inhibitors, cell adhesion
promoters and drugs that may enhance the formation of healthy neointimal
tissue, including endothelial cell regeneration. The positive action may
come from inhibiting particular cells (e.g., smooth muscle cells) or
tissue formation (e.g., fibromuscular tissue) while encouraging different
cell migration (e.g., endothelium) and tissue formation (neointimal
 The preferred materials for fabricating the braided stent include
stainless steel, tantalum, titanium alloys including nitinol (a nickel
titanium, thermomemoried alloy material), and certain cobalt alloys
including cobalt-chromium-nickel alloys such as Elgiloy.RTM. and
Phynox.RTM.. Further details concerning the fabrication and details of
other aspects of the stents themselves, may be gleaned from the above
referenced U.S. Pat. Nos. 4,655,771 and 4,954,126 to Wallsten and
5,061,275 to Wallsten et al. To the extent additional information
contained in the above-referenced patents is necessary for an
understanding of the present invention, they are deemed incorporated by
 Various combinations of polymer coating materials can be
coordinated with biologically active species of interest to produce
desired effects when coated on stents to be implanted in accordance with
the invention. Loadings of therapeutic materials may vary. The mechanism
of incorporation of the biologically active species into the surface
coating, and egress mechanism depend both on the nature of the surface
coating polymer and the material to be incorporated. The mechanism of
release also depends on the mode of incorporation. The material may elute
via interparticle paths or be administered via transport or diffusion
through the encapsulating material itself.
 For the purposes of this specification, "elution" is defined as
any-process of release that involves extraction or release by direct
contact of the material with bodily fluids through the interparticle
paths connected with the exterior of the coating. "Transport" or
"diffusion" are defined to include a mechanism of release in which a
material released traverses through another material.
 The desired release rate profile can be tailored by varying the
coating thickness, the radial distribution (layer to layer) of bioactive
materials, the mixing method, the amount of bioactive material, the
combination of different matrix polymer materials at different layers,
and the crosslink density of the polymeric material. The crosslink
density is related to the amount of crosslinking which takes place and
also the relative tightness of the matrix created by the particular
crosslinking agent used. This, during the curing process, determines the
amount of crosslinking and so the crosslink density of the polymer
material. For bioactive materials released from the crosslinked matrix,
such as heparin, a crosslink structure of greater density will increase
release time and reduce burst effect.
 Additionally, with eluting materials such as heparin, release
kinetics, particularly initial drug release rate, can be affected by
varying the average dispersed particle size. The observed initial release
rate or burst effect may be substantially reduced by using smaller
particles, particularly if the particle size is controlled to be less
than about 15 microns and the effect is even more significant in the
particle size range of .ltoreq.10 microns, especially when the coating
thickness is not more than about 50 .mu.m and drug loading is about 25-45
 It will also be appreciated that an unmedicated silicone thin top
layer provides an advantage over drug containing top coat. Its surface
has a limited porosity and is generally smooth, which may be less
thrombogeneous and may reduce the chance to develop calcification, which
occurs most often on the porous surface.
BRIEF DESCRIPTION OF THE DRAWINGS
 In the drawings, wherein like numerals designate like parts
throughout the same:
 FIG. 1 is a schematic flow diagram illustrating the steps of the
process of the invention;
 FIG. 2 represents a release profile for a multi-layer system
showing the percentage of heparin released over a two-week period;
 FIG. 3 represents a release profile for a multi-layer system
showing the relative release rate of heparin over a two-week period;
 FIG. 4 illustrates a profile of release kinetics for different drug
loadings at similar coating thicknesses illustrating the release of
heparin over a two-week period;
 FIG. 5 illustrates drug elution kinetics at a given loading of
heparin over a two-week period at different coating thicknesses;
 FIG. 6 illustrates the release kinetics in a coating having a given
tie-layer thickness for different top coat thicknesses in which the
percentage heparin in the tie coat and top coats are kept constant;
 FIG. 7 illustrates the release kinetics of several coatings having
an average coating thickness of 25 microns and a heparin loading of 37.5%
but using four different average particle sizes;
 FIGS. 8-11 are photomicrographs of coated stent fragments for the
coatings of FIG. 7 having a corresponding average particle size of 4
microns, 17 microns, 22 microns and 30 microns, respectively.
 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 FIG. 1. 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.
 General Preparation of Heparin Coating Composition
 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'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.sub.silicone=5 g; solid percent=35%; W.sub.hep=5.times.0.35.ti-
mes.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.sub.silicone
solid/V.sub.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.
 General Preparation of Dexamethasone Coating Composition
 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.sub.silicone=5 g; for a 10% dexamethasone coating,
W.sub.dex=5.times.0.35.times.0.1/0.9=0.194 g and THF needed in the
coating solution calculated. W.sub.silicone solid/V.sub.THF=0.06 for a
10% dexamethasone coating solution. Example: W.sub.silicone=5 g;
V.sub.THF=5.times.0.35/0.06.apprxeq.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.
 Any appropriate method can be used to secure the stent for spraying
and 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.sup.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 any layer can be adjusted by
modifying 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 FIG. 1, the coated stent is thereafter subjected
to a curing step in which the pre-polymer 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.degree. 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.degree. C. or higher for approximately 16 hours.
The temperature may be raised to as high as 150.degree. C. for
dexamethasone containing coated stents. Of course, the time and
temperature may vary with particular silicones, crosslinkers, and
biologically active species.
 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 RF 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 cm) 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 20-50 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
 With respect to the anticoagulant material heparin, the percentage
in the tie layer is nominally from about 20-50% 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 1:10 to 1:2 and is preferably in
the range of from about 1:6 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.
 FIG. 4 depicts the general effect of drug loading for coatings of
similar thickness. The initial elution rate increases with the drug
loading as shown in FIG. 5. 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 FIG. 6.
 The effect of average particle size is depicted in the FIGS. 7-11
in which coating layers with an average coating thickness of about 25
microns (.mu.m), prepared and sterilized as above, were provided with
dispersed heparin particles (to 37.5% heparin) of several different
average particle sizes. FIG. 7 shows plots of elution kinetics for four
different sizes of embedded heparin particles. The release took place in
phosphate buffer (pH 7.4) at 37.degree. C. The release rate using
smaller, particularly 4-6 .mu.m average sized particles noticeably
reduces the initial rate or burst effect and thereafter the elution rate
decreases more slowly with time. Average particle sizes above about 15
.mu.m result in initial release rates approaching bolus elution. This, of
course, is less desirable, both from the standpoint of being an
unnecessary initial excess and for prematurely depleting the coating of
deserved drug material.
 In addition, as shown in the photomicrographs of FIGS. 8-11, as the
average particle size increases, the morphology of the coating surface
also changes coatings containing larger particles (FIGS. 9-11) have very
rough and irregular surface characteristics. These surface irregularities
may be more thrombogenic or exhibit an increased tendency to cause
embolization when the corresponding stent is implanted in a blood vessel.
 Accordingly, it has been found that the average particle size
should generally be controlled below about 15 .mu.m to reduce the burst
effect and preferably should be .ltoreq. about 10 .mu.m for best results.
The 4-6 .mu.m size worked quite successfully in the laboratory. However,
it should be noted that larger particle size can also be advantageously
used, for instance, when the drug load is low, such as below 25 weight
percent. Elution kinetics can be adjusted by a combination of changing
the particle size and changing the load or concentration of the dispersed
 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 modified 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 undercoat and
anti-thrombotic drugs in the topcoat layer. In this manner, the
anti-thrombotic drugs, for example, heparin, will elute first followed by
antiproliferation drugs, e.g. dexamethasone, 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.degree. 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 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.
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