Note: Descriptions are shown in the official language in which they were submitted.
CA 02470709 2005-09-29
60950-255D
-1-
STENT WITH DRUG RELEASE COATING
BACKGROUND OF THE INVENTION
Field of the Invention
The present divisional application is divided out
of parent Application No. 2,231,727, filed September 5,
1996.
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 structure.
The invention of the parent application relates to
a method of coating an implantable prosthesis, a method of
controlling the delivery kinetics of an eluting biologically
active particulate matter incorporated in an elastomeric
coating having at least one opening therein, and a coated
implantable prosthesis having at least one opening therein.
The invention of the present divisional
application relates to 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, incorporating
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. The invention of the present
CA 02470709 2004-07-05
60950-255D
-la-
divisional application also relates to a method of making
the medical device.
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
CA 02470709 2004-07-05
60950-255D
-2-
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
implantation.
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
CA 02470709 2004-07-05
60950-255D
-3-
U.S..Patents 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.
5. Patent 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. Patent 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-deg'radable or bio-sorbable polymers
include Tang et al, U.S. Patent 4,9i6,193, and,
MacGregor, U.S. Patent 4,994,071. Sahatjian in U.S.
Patent 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. Patent 5,464,650, issued November 5, 1995
and corresponding to European patent application 0 623
354 Al published'09 November 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
CA 02470709 2004-07-05
60950-255D
-4-
and Process Engineering Technical Conference (1990)
relates to polymer/drug/membrane $yst,ems for releasing
heparin. Those polymer/drug/membrane systems require
two distinct layers to function.
-5 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 techniquesthat 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 geziera'lly
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
CA 02470709 2004-07-05
60950-255D
-5-
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
5.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.
CA 02470709 2004-07-05
60950-255D
-6-
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
According to one aspect of the invention of the
parent application, there is provided a method of coating an
implantable prosthesis, having at least one opening therein,
with at least one layer comprising a hydrophobic elastomeric
material incorporating an amount of biologically active
material therein for timed delivery therefrom comprising the
steps of: (a) applying a formulation containing the
elastomeric material in solvent mixture and an amount of a
biologically active material to a surface of the prosthesis;
wherein when the biologically active material is
particulate, the average particle size of the biologically
active material in said formulation is less than or equal to
about 15 m; and wherein the coating is applied to the
prosthesis in a manner to adheringly conform thereto to
preserve the opening; and (b) curing the elastomeric
material such that at least some of the biologically active
material is particulate after curing.
According to another aspect of the invention of
the parent application, there is provided a method of
controlling the delivery kinetics of an eluting biologically
active particulate material incorporated in an elastomeric
coating having at least one layer on a surface of an
implantable prosthesis, having at least one opening therein,
the method comprising incorporating a biologically active
particulate material having an average particle size of less
than or equal to about 15 m into at least one layer of the
coating;
CA 02470709 2004-07-05
60950-255D
-6a-
and applying the coating to the prosthesis in a manner to
adheringly conform thereto to preserve the opening; and
curing the coating such that at least some of the
biologically active material is particulate after curing.
According to still another aspect of the invention
of the parent application, there is provided a coated
implantable prosthesis, having at least one opening therein,
the prosthesis having an external surface covered with at
least one layer comprising a hydrophobic elastomeric
material incorporating an amount of biologically active
material in particulate form dispersed therein for timed
delivery therefrom wherein the average particle size of the
biologically active material is less than or equal to about
m; and wherein the coating adheringly conforms to the
15 prosthesis to preserve the opening.
According to one aspect of the divisional
application, there is provided 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,
incorporating 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.
According to another aspect of the divisional
application, there is provided a method of making the
medical device as described herein, wherein said method
comprises the steps of: (a) applying an undercoat of a
formulation containing uncured hydropho:bic elastomeric
material in solvent mixture and an amount of biologically
active material that is finely divided; (b) curing said
CA 02470709 2005-09-29
60950-255D
-6b-
hydrophobic elastomeric material; and (c) applying a topcoat
of a formulation comprising polymeric material to form the
topcoat.
According to one aspect of the invention of the
present divisional application, there is provided a balloon-
expandable stent comprising: a metallic intravascular
balloon-expandable open lattice sidewall stent structure
designed for permanent implantation into a blood vessel of a
patient; a first polymer composition conforming to the open
lattice sidewall stent structure so as to preserve the open
lattice sidewall stent structure, wherein the first polymer
composition comprises a first biostable polymer and an agent
that inhibits restenosis; and a second polymer composition
applied to the first polymer composition so as to preserve
the open lattice sidewall stent structure, wherein the
second polymer composition comprises a second biostable
polymer, the second biostable polymer being different from
the first biostable polymer, wherein the second polymer
composition is substantially free of the agent that inhibits
restenosis when applied to the first polymer composition,
and wherein the second biostable polymer is non-
thrombogenic.
According to another aspect of the invention of
the present divisional application, there is provided a
method of making a balloon-expandable stent for delivery of
an agent that inhibits restenosis into a blood vessel of a
patient comprising: providing a metallic intravascular
balloon-expandable open lattice sidewall stent structure
designed for permanent implantation into a blood vessel of a
patient; applying a first polymer composition to the open
lattice sidewall stent structure so as to preserve the open
lattice sidewall stent structure, said first polymer
composition conforming to the open lattice sidewall stent
CA 02470709 2005-09-29
60950-255D
-6c-
structure, and wherein the first polymer composition
comprises a first biostable polymer and an agent that
inhibits restenosis; and applying a second polymer
composition to the first polymer composition so as to
preserve the open lattice sidewall stent structure, wherein
the second polymer composition comprises a second biostable
polymer, the second biostable polymer being different from
the first biostable polymer, wherein the second polymer
composition is substantially free of the agent that inhibits
restenosis when applied to the first polymer composition,
and wherein the second biostable polymer is non-
thrombogenic.
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 "finely divided" means any type or size of included
CA 02470709 2005-09-29
60950-255D
-6d-
material from dissolved molecules through suspensions,
colloids and particulate mixtures. The active material is
dispersed in a carrier
CA 02470709 2004-07-05
60950-255D
-7-
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
5, 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 materialq 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 empioying 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 precurser 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
CA 02470709 2004-07-05
60950-255D
-8-
the elastomeric or polymeric material into.the fini'shed
or useful state by the application,of heat and/or
chemical agents which include physical-chemical charges.
Where, for example, polyurethane thermoplastic
5. 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
10. 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
15 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. T~ preferred species have a
relatively low cure temperatures and are known as a room
20 temperature vulcanizable (RTV) materials. Some
polydimethylsiloxane materials can be cured, for
example, by exposure to air at about 90 C for a period
of time such as 16 hours. A curing step may be
implemented both after application of a certain number
25 of lower undercoat layers and the topcoat layers or a
single curing step used after coating is completed.
The coated stents may thereafter be subjected to a
postcure sterilization process which includes an inert
gas plasma treatment, and then exposure to gamma
CA 02470709 2004-07-05
60950-255D
-9-
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
5. purged with nitrogen and a vacuum applied to about 20-
50mTorr. Thereafter, inert gas (argon, helium or
mixture of them) is admitted to the reaction.chambe.r 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 tim4 from 3.0 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 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 appropriate rate.
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
CA 02470709 2004-07-05
60950-255D
-10-
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),
-5 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 invention.
Agents suitable for incorporation include
antithrobotics, anticoagulants, antiplatelet agents,
thrombolytics, antiproliferatives,'antinflainmatories,
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 tissue).
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 and
Phynox . Further details concerning the fabrication and
CA 02470709 2004-07-05
60950-255D
-1].-
details of other aspects of the stents themselves, may
be gleaned from the above referenced U.S. Patents
4,655,771 and 4.,954,126 to Wallsten and 5,061*,275 to
Walisten et al.
Various combinations.of polymer coating materials
can be coordinated with biologically active species of
interest to pro4uce desired effects wrhen coated on
stents to be implanted in accordance with the invention.
Loadings of therapeutic materials,inay 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
.30 varying the coating thickness, the radial distribution
CA 02470709 2004-07-05
60950-255D
-12-
(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 ef fect .
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 less than or equal to 10 microns, especially
when the coating thickness is not more than about 50 m
and drug loading is about 25-45 weight percent.
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
CA 02470709 2004-07-05
60950-255D
-13-
calcification, which occurs most often on the porous
surf ace .
BRIEF DESCRIPTION Ok' THE DRAWINGS
In-the drawings, wherein like numerals designate
5, like parts throughout the same:
FIGURE 1 is a schematic flow diagram illustrating
the steps of the process of the invention;
FIGURE 2 represents a release profile for a multi-
layer system showing the percentage of heparin released
over a two-week period;
FIGURE 3 represents a release profile for a multi-
layer system showing the relative release rate of
heparin-over a two-week period;
FIGURE 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;
FIGURE 5 illustrates drug elution.kinetics at a
given loading of heparin over a two-week period at
different coating thicknesses;
FIGURE 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;
FIGURE 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;
FIGURES 8-11 are photomicrographs,of coated stent
fragments for the coatings of FIGURE 7 having a
CA 02470709 2004-07-05
60950-255D
-14-
corresponding average particle size of 4 microns, 17
microns, 22 microns and 30 microns, respectively.
DETAILED DESCRIPTION
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'Figure 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 applicationcontainer 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 Prepa a jon of Heparin Coating Composition
Silicone was obtained as a polymer precursor in
solvent (xylene) mixture. For example, a 35 s 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 37.5n heparin coating, for example:
CA 02470709 2004-07-05
60950-255D
-15-
W(silicone) = 5 g; solid percent = 35 a; W'(,hep) = 5 x
0.35 x.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)/V(THF) = 0.04 for a 37.50
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
applicationby 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 40a) and the required amount was
then weighed. Example: W(silicone) = 5 g; for a 10%
dexamethasone coating, W(dex) = 5 x 0.35 x 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 x 0. 35/0 . 06 = 29.17 rnl. The
dexamethasone was weighed in a beaker on an analytical
balance and half the total amount of THF was added.
CA 02470709 2004-07-05
60950-255D
-].6-
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
.5: 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
1-1.7 atm. (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, nomirially about 40 rpm.
CA 02470709 2004-07-05
60950-255D
-17-
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
5-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 thre.e 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 cm2 of projected area, 20 cycles of
coating application are required.and.about 30 ml of
,,
solution vaill 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
CA 02470709 2004-07-05
60950-255D
-18-
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 thiekness 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 Figure 1, t=he 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 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, and biologically active species.
CA 02470709 2004-07-05
60950-255D
-19-
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 a.nvolves an argon plasma
treatment of the coated, cured stents in the
unconstrained configuration. In accordance withthis
procedure,,the stents are placed in a chamber of a
plasma surf'ace treatment system such as a Plasma Science
350 (Himont/Plasma Science, Foster City, CA). 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 42.55 cm (16.75 inches) by
34.3 cm (13.5 inches) by 44.45 cm (17.5 inches) 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 20-50mTorr.
Thereafter, inert gas (argon, helium or mixture of them).
is admitted to the reaction chamber for the plasma
CA 02470709 2004-07-05
60950-255D
-20-
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
20-50% and that of the top layer from about 0-30k 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 maximizirig second day and
third day release at the thinnest possible coating
configuration will reduce the acute or subcute
thrombosis.
CA 02470709 2004-07-05
60950-255D
-21-
Figure 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
Figure 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 Figure 6.
The effect of average particle size is depicted in
the FIGURES 7-11 in which coating layers with an average
coating thickness of about 25 microns ( m), prepared and
sterilized as above, were providedwith dispersed
heparin particles (to 37.5* heparin) of several
different average particle si=zes. FIGURE 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 C. The release rate using
smaller, particularly 4-6 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 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
FIGURES 8-11, as the average particle size increases,
the morphology of the coating surface also changes.
CA 02470709 2004-07-05
60950-255D
-22-
Coatings containing larger particles (FIGURES 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
5,.-the corresponding stent is implanted iri a blood vessel.
Accordingly, it has been found that the average
particle size should generally be controlled below about
15 m to reduce the burst effect and preferably should
be less than or equal to about 10 m for.best results.
The 4-6 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.be.low 25
weight percent. Elution kinetics can be adj.usted by a
combination,of changing the particle size and changing
the load or concentration of the dispersed drug
material.
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 andrate 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,
CA 02470709 2004-07-05
60950-255D
-23-
e.g. dexamethasone, to better enable safe encapsulat'ion
of the implanted stent.
The heparin concentration measurements 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 60o..of the
total weight'of the layer. However, the loading:generally most advantageously
used is in the range from
about 10% to 450 of the total weight of the layer. In
the case of dexamethasone, the loading may be as high as
50t or more of the total weight of the layer but is
preferably in t=he range of about 0.4t 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
CA 02470709 2004-07-05
60950-255D
-24-
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.
5' 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
CA 02470709 2004-07-05
60950-255D
-25-
controlling the release of biologically active materials=
with regard to an implanted stent.,,;Each applied layer
is typically from approximately 0:'S 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 biologa.cally.active..
species can generally be more predictably controlled
with such,inaterials. 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.
