U.S. patent application number 10/989946 was filed with the patent office on 2005-05-19 for medical device with drug.
This patent application is currently assigned to Boston Scientific Scimed, Inc.. Invention is credited to Ding, Ni, Helmus, Michael N..
Application Number | 20050106210 10/989946 |
Document ID | / |
Family ID | 27062077 |
Filed Date | 2005-05-19 |
United States Patent
Application |
20050106210 |
Kind Code |
A1 |
Ding, Ni ; et al. |
May 19, 2005 |
Medical device with drug
Abstract
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.
Inventors: |
Ding, Ni; (Plymounth,
MN) ; Helmus, Michael N.; (Long Beach, CA) |
Correspondence
Address: |
JONES DAY
222 EAST 41ST ST
NEW YORK
NY
10017
US
|
Assignee: |
Boston Scientific Scimed,
Inc.
|
Family ID: |
27062077 |
Appl. No.: |
10/989946 |
Filed: |
November 16, 2004 |
Related U.S. Patent Documents
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Application
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Filing Date |
Patent Number |
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10989946 |
Nov 16, 2004 |
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10067041 |
Feb 4, 2002 |
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10067041 |
Feb 4, 2002 |
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09012443 |
Jan 23, 1998 |
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6358556 |
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09012443 |
Jan 23, 1998 |
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08663490 |
Jun 13, 1996 |
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5837313 |
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09012443 |
Jan 23, 1998 |
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08663518 |
Jun 13, 1996 |
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6120536 |
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08663518 |
Jun 13, 1996 |
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08526273 |
Sep 11, 1995 |
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08663518 |
Jun 13, 1996 |
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08424884 |
Apr 19, 1995 |
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Current U.S.
Class: |
424/423 ;
427/2.1; 623/1.42 |
Current CPC
Class: |
A61L 2300/236 20130101;
A61L 2300/606 20130101; A61L 2300/622 20130101; A61F 2/90 20130101;
A61F 2250/0067 20130101; A61L 31/16 20130101; A61L 2420/02
20130101; A61L 2300/406 20130101; A61F 2/86 20130101; A61F 2/82
20130101; A61L 2300/222 20130101; A61F 2210/0014 20130101; A61L
31/10 20130101; A61L 31/141 20130101; A61L 33/0011 20130101; A61L
2300/608 20130101; A61L 27/227 20130101; A61L 2300/42 20130101;
A61L 2300/602 20130101; A61L 31/10 20130101; C08L 71/02 20130101;
A61L 31/10 20130101; C08L 83/04 20130101; A61L 31/10 20130101; C08L
83/00 20130101; A61L 31/10 20130101; C08L 83/08 20130101; A61L
31/10 20130101; C08L 5/00 20130101; A61L 31/10 20130101; C08L 75/04
20130101 |
Class at
Publication: |
424/423 ;
427/002.1; 623/001.42 |
International
Class: |
A61F 002/02; B05D
003/00; A61F 002/06 |
Claims
1-23. (canceled)
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 steel.
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 polymeric material.
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 topcoat.
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 conform thereto.
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 heparin.
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 curing.
38. The method of claim 34, wherein the biologically active
material includes heparin.
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
self-expanding stent.
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.
Description
I. CROSS-REFERENCE TO RELATED APPLICATIONS
[0001] 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
[0002] II. Field of the Invention
[0003] 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.
[0004] II. Related Art
[0005] 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 implantation.
[0006] 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.
[0007] 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.
[0008] 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.
[0009] 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.
[0010] 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.
[0011] 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.
[0012] 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.
[0013] 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.
[0014] 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.
[0015] 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.
[0016] 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
[0017] 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.
[0018] 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.
[0019] 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.
[0020] 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.
[0021] 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 is
completed.
[0022] 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.
[0023] 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
minutes.
[0024] 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.
[0025] 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
invention.
[0026] 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
tissue).
[0027] 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 reference herein.
[0028] 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.
[0029] 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.
[0030] 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.
[0031] 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 weight percent.
[0032] 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
[0033] In the drawings, wherein like numerals designate like parts
throughout the same:
[0034] FIG. 1 is a schematic flow diagram illustrating the steps of
the process of the invention;
[0035] FIG. 2 represents a release profile for a multi-layer system
showing the percentage of heparin released over a two-week
period;
[0036] FIG. 3 represents a release profile for a multi-layer system
showing the relative release rate of heparin over a two-week
period;
[0037] 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;
[0038] FIG. 5 illustrates drug elution kinetics at a given loading
of heparin over a two-week period at different coating
thicknesses;
[0039] 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;
[0040] 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;
[0041] 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.
DETAILED DESCRIPTION
[0042] 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.
[0043] General Preparation of Heparin Coating Composition
[0044] 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.
[0045] General Preparation of Dexamethasone Coating Composition
[0046] 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.
[0047] 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.
[0048] 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.
[0049] 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.
[0050] 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.
[0051] 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.
[0052] 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.
[0053] 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.
[0054] 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.
[0055] 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.
[0056] 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.
[0057] 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.
[0058] 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.
[0059] 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.
[0060] 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.
[0061] 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.
[0062] 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 drug
material.
[0063] 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.
[0064] 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.
[0065] 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.
[0066] 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%.
[0067] 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.
[0068] 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.
[0069] 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.
[0070] 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.
[0071] 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.
* * * * *