U.S. patent application number 11/116256 was filed with the patent office on 2005-08-25 for drug coating with topcoat.
This patent application is currently assigned to Boston Scientific Scimed, Inc.. Invention is credited to Ding, Ni, Helmus, Michael N..
Application Number | 20050187611 11/116256 |
Document ID | / |
Family ID | 25542881 |
Filed Date | 2005-08-25 |
United States Patent
Application |
20050187611 |
Kind Code |
A1 |
Ding, Ni ; et al. |
August 25, 2005 |
Drug coating with topcoat
Abstract
A coating and method for a coating an implantable device or
prostheses are disclosed. The coating includes an undercoat of
polymeric material containing an amount of biologically active
material, particularly heparin, dispersed herein. The coating
further includes a topcoat which covers less than the entire
surface of the undercoat and wherein the topcoat comprises a
polymeric material substantially free of pores and porosigens. The
polymeric material of the topcoat can be a biostable, biocompatible
material which provides long term non-thrombogenicity to the device
portion during and after release of the biologically active
material.
Inventors: |
Ding, Ni; (Plymouth, 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: |
25542881 |
Appl. No.: |
11/116256 |
Filed: |
April 27, 2005 |
Related U.S. Patent Documents
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Application
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Filing Date |
Patent Number |
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11116256 |
Apr 27, 2005 |
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10603115 |
Jun 24, 2003 |
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10603115 |
Jun 24, 2003 |
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09942716 |
Aug 30, 2001 |
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6620194 |
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09942716 |
Aug 30, 2001 |
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09573506 |
May 18, 2000 |
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6284305 |
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09573506 |
May 18, 2000 |
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08996410 |
Dec 22, 1997 |
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6099562 |
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08996410 |
Dec 22, 1997 |
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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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08526273 |
Sep 11, 1995 |
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08424884 |
Apr 19, 1995 |
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Current U.S.
Class: |
623/1.15 ;
623/1.42 |
Current CPC
Class: |
A61F 2250/0067 20130101;
A61L 27/227 20130101; A61L 2300/606 20130101; A61L 31/10 20130101;
A61L 2300/608 20130101; A61L 2300/236 20130101; A61L 31/10
20130101; A61L 31/10 20130101; A61L 31/10 20130101; A61L 2300/406
20130101; A61L 2300/416 20130101; C08L 71/02 20130101; C08L 5/00
20130101; C08L 83/08 20130101; C08L 5/00 20130101; C08L 83/08
20130101; C08L 83/00 20130101; C08L 83/04 20130101; C08L 71/02
20130101; A61F 2/82 20130101; A61L 31/10 20130101; A61L 31/08
20130101; A61L 31/16 20130101; A61L 31/10 20130101; A61L 31/10
20130101; A61L 31/141 20130101; A61L 33/0011 20130101; A61F
2210/0014 20130101; A61F 2/90 20130101; A61L 31/16 20130101; A61L
2300/42 20130101; A61F 2/86 20130101; A61L 2420/08 20130101; A61L
2300/602 20130101 |
Class at
Publication: |
623/001.15 ;
623/001.42 |
International
Class: |
A61F 002/06 |
Claims
1-50. (canceled)
51. A medical device comprising: a metallic intravascular stent
comprising an open lattice sidewall structure and designed for
permanent implantation into a blood vessel of a patient; a first
quantity of a first polymer composition applied to at least a
portion of said stent and conforming to said open lattice sidewall
structure so as to preserve the open lattice sidewall structure of
said stent, wherein said first polymer composition comprises a
first biostable polymer and a biologically active material; and a
second quantity of a second polymer composition applied to at least
a portion of said first quantity and conforming to said stent so as
to preserve the open lattice sidewall structure of said stent,
wherein said second polymer composition comprises a second
biostable polymer that is different from said first biostable
polymer, and wherein said second polymer composition is
non-thrombogenic and substantially free of the biologically active
material or other elutable material, wherein when in use, the
biologically active material is released from the stent to the
blood vessel at a first rate that is different from a second rate,
wherein the second rate is the rate of release of the same
biologically active material from the stent had the second quantity
of the second polymer composition not been applied to the first
quantity of the first polymer composition.
52. The medical device of claim 51, wherein the metallic
intravascular stent is a stainless steel intravascular stent.
53. The medical device of claim 51, wherein the first quantity of
first polymer composition is different from the second quantity of
second polymer composition.
54. The medical device of claim 51, wherein the first biostable
polymer comprises a hydrophobic elastomeric material, an ethylene
vinyl acetate copolymer material, or a mixture thereof.
55. The medical device of claim 51, wherein the biologically active
material is an antithrobotic agent, anticoagulant, antibiotic,
antiplatelet agent, thrombolytic agent, antiproliferative agent,
steroidal antiinflammatory agent, nonsteroidal antiinflammatory
agent, agent that inhibits hyperplasia, smooth muscle cell
inhibitor, growth factor, growth factor inhibitor, cell adhesion
inhibitor, cell adhesion promoter, drug that enhances the formation
of healthy neointimal tissue, or a mixture thereof.
56. The medical device of claim 51, wherein the biologically active
material is an antibiotic.
57. The medical device of claim 51, wherein the biologically active
material inhibits smooth muscle cell.
58. The medical device of claim 51, wherein the biologically active
material inhibits restenosis.
59. The medical device of claim 51, wherein the second biostable
polymer remains non-thrombogenic during and after release of the
biologically active material.
60. The medical device of claim 51, wherein the stent releases the
biologically active material over a period of time.
61. A method of treating restenosis comprising implanting the
medical device of claim 51 into the blood vessel of the
patient.
62. A medical device comprising: a metallic intravascular stent
comprising an open lattice sidewall structure and designed for
permanent implantation into a blood vessel of a patient; a first
quantity of a first polymer composition applied to at least a
portion of said stent and conforming to said open lattice sidewall
structure so as to preserve the open lattice sidewall structure of
said stent, wherein said first polymer composition comprises an
ethylene vinyl acetate copolymer material and an antibiotic; and a
second quantity of a second polymer composition applied to at least
a portion of said first quantity and conforming to said stent so as
to preserve the open lattice sidewall structure of said stent,
wherein said second polymer composition comprises a second
biostable polymer that is different from the ethylene vinyl acetate
copolymer material of said first polymer composition, and wherein
said second polymer composition is non-thrombogenic and
substantially free of the antibiotic of the first polymer
composition or other elutable material, wherein when in use, the
antibiotic is released from the stent to the blood vessel at a
first rate that is different from a second rate, wherein the second
rate is the rate of release of the same antibiotic from the stent
had the second quantity of the second polymer composition not been
applied to the first quantity of the first polymer composition.
63. The medical device of claim 62, wherein the metallic
intravascular stent is a stainless steel intravascular stent.
64. The medical device of claim 62, wherein the first quantity of
first polymer composition is different from the second quantity of
second polymer composition.
65. The medical device of claim 62, wherein the antibiotic inhibits
smooth muscle cell.
66. The medical device of claim 62, wherein the antibiotic inhibits
restenosis.
67. The medical device of claim 62, wherein the second biostable
polymer remains non-thrombogenic during and after release of the
antibiotic.
68. The medical device of claim 62, wherein the stent releases the
antibiotic over a period of time.
69. A method of treating restenosis comprising implanting the
medical device of claim 62 into the blood vessel of the patient.
Description
CROSS-REFERENCE TO RELATED APPLICATIONS
[0001] The present application is a Continuation-In-Part of
co-pending application Ser. No. 08/633,518, filed Jun. 13,
1996.
BACKGROUND OF THE INVENTION
[0002] 1. Field of the Invention
[0003] The present invention relates generally to providing
biostable elastomeric coatings on the surfaces of implants which
incorporate biologically active species having controlled release
characteristics in the coating and relates particularly to
providing a non-thrombogenic surface during and after timed release
of the biologically active species. The invention is particularly
described in terms of coatings on therapeutic expandable stent
prostheses for implantation in body lumens, e.g., vascular
implantation.
[0004] 2. Related Art
[0005] In surgical or other related invasive procedures, the
insertion and expansion of stent devices in blood vessels, urinary
tracts or other locations difficult to otherwise access 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 prostheses 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 the same
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. Prostheses 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 U.S. Pat. No.
5,061,275 to Wallsten et al.
[0007] Implanted stents have been used to carry medicinal agents,
such as thrombolytic agents. U.S. Pat. No. 5,163,952 to Froix
discloses a thermalmemoried expanding plastic stent device
formulated to carry a medicinal agent in the material of the stent
itself. Pinchuk, in U.S. Pat. No. 5,092,877, discloses a stent of a
polymeric material which may have 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.
[0008] A patent to Sahatjian, U.S. Pat. No. 5,304,121, discloses a
coating applied to a stent consisting of a hydrogel polymer and a
preselected drug such as a cell growth inhibitors or heparin. A
further method of making a coated intravascular stent carrying a
therapeutic material is described in Berg et al., U.S. Pat. No.
5,464,650, issued on Nov. 7, 1995 and corresponding to European
Patent Application No. 0 623 354 A1 published 9 Nov. 1994. In that
disclosure, a polymer coating material is dissolved in a solvent
and the therapeutic material dispersed in the solvent; the solvent
evaporated after application.
[0009] 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 types of layers to function.
[0010] It has been recognized that contacting blood with the
surface of a foreign body in vivo has a tendency to induce
thrombogenic responses and that as the surface area of a foreign
device in contact with host blood increases, the tendency for
coagulation and clot forming at these surfaces also increases. This
has led to the use of immobilized systemic anti-coagulant or
thrombolytic agents such as heparin on blood contacting surfaces
such as oxygen uptake devices to reduce this phenomenon. Such an
approach is described by Winters, et al., in U.S. Pat. Nos.
5,182,317; 5,262,451 and 5,338,770 in which the amine functional
groups of the active material are covalently bonded using
polyethylene oxide (PEO) on a siloxane surface.
[0011] Another approach is described in U.S. Pat. No. 4,613,665 to
Larm in which heparin is chemically covalently bound to plastic
surface materials containing primary amino groups to impart a
non-thrombogenic surface to the material. Other approaches for
bonding heparin are described in Barbucci, et al., "Coating of
commercially available materials with a new heparinizable
material", Journal of Biomedical Materials Research, Vol. 25, pp.
1259-1274 (1991); Hubbell, J. A., "Pharmacologic Modification of
Materials", Cardiovascular Pathology, Vol. 2, No. 3 (Suppl.),
121S-127S (1993); Gravlee, G. P., "Heparin-Coated Cardiopulmonary
Bypass Circuits", Journal of Cardiothoracic and Vascular
Anesthesia, Vol. 8, No. 2, pp. 213-222 (1994).
[0012] Moreover, drug elution rates for a coating containing a
hydrophilic or a lipophobic drug is usually very fast initially
when the coated device contacts body fluid or blood. One of the
methods to reduce the so called "burst effect" is to add a membrane
containing porosigen over the coating layer containing the
biologically active material. See e.g., U.S. Pat. No. 5,605,696 to
Eury et al. and U.S. Pat. No. 5,447,724 to Helmus et al. When the
porosigen elutes, a porous membrane is formed and the drug in the
undercoat will release. Even though the method might be quite
successful to control the drug release, it increases the coating
thickness, reduces the effective drug loading and introduces
undesirable additional foreign materials into the patient. Hence,
there is a need for a coating which reduces the burst effect but is
not too thick and does not require the release of porosigens into
the body.
[0013] With regard to stents, polymeric stents, although effective,
may have mechanical properties that are inferior to those of metal
stents of like thickness and weave. Metallic vascular stents
braided of even relatively fine metal can provide a large amount of
strength to resist inwardly directed circumferential pressure. A
polymer material of comparable strength requires a much
thicker-walled structure or heavier, denser filament weave, which
in turn, reduces the cross-sectional area available for flow
through the stent and/or reduces the relative amount of open space
in the weave. Also, it is usually more difficult to load and
deliver polymeric stents using catheter delivery systems.
[0014] While certain types of stents such as braided metal stents
may be preferred for some applications, the coating and coating
modification process of the present invention is not so limited and
can be used on a wide variety of prosthetic devices. Thus, in the
case of stents, 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; as well as
polymeric stents of all kinds. Other medical devices that can
benefit from the present invention include blood exchanging
devices, vascular access ports, central venus catheters,
cardiovascular catheters, extracorpeal circuits, vascular grafts,
pumps, heart valves, and cardiovascular sutures, to name a few.
Regardless of detailed embodiments, applicability of the invention
should not be considered limited with respect to implant design,
implant location or materials of construction. Further, the present
invention may be used with other types of implantable
prostheses.
[0015] Accordingly, it is a primary object of the present invention
to provide a coating and process for coating a stent to be used as
a deployed stent prostheses, the coating being capable of effective
controlled long-term delivery of biologically active materials.
[0016] Another object of the invention is to provide a coating and
process for coating a stent prostheses using a biostable
hydrophobic elastomer in which biologically active species are
incorporated within a coating.
[0017] Still another object of the present invention is to provide
a multi-layer coating and process for the delivery of biologically
active species in which the percentage of active material can vary
from layer to layer.
[0018] Yet another object of the present invention is to provide a
multi-layer coating and process for the delivery of biologically
active species from a coating with a non-thrombogenic surface.
[0019] Another object of the invention is to provide a coating for
the delivery of biologically active species having a top layer or
topcoat which reduces the initial release of the species, in which
the topcoat is substantially free of pores or porosigens and covers
less than the entire surface of the undercoat. The topcoat can
cover less than the entire surface of the undercoat before and/or
while the device is implanted.
[0020] A further object of the invention is to provide a multilayer
coating for the delivery of biologically active species such as
heparin having a fluorosilicone top layer.
[0021] A still further object of the invention is to provide a
multi-layer coating for the delivery of biologically active species
such as heparin having a surface containing immobilized
polyethylene glycol (PEG).
[0022] 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
[0023] The present invention provides a relatively thin layered
coating of biostable elastomeric material containing an amount of
biologically active material dispersed therein in combination with
a non-thrombogenic surface that is useful for coating the surfaces
of prostheses such as deployable stents.
[0024] The preferred stent to be coated is a self-expanding,
open-ended tubular stent prostheses. Although other materials,
including polymer materials, can be used, in the preferred
embodiment, the tubular body is formed of a self expanding open
braid of fine single or polyfilament metal wire which flexes
without collapsing, readily axially deforms to an elongate shape
for transluminal insertion via a vascular catheter and resiliently
expands toward predetermined stable dimensions upon removal in
situ.
[0025] In the process, the initial 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 biologically
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.
[0026] In many applications the layered coating is referred to or
characterized as including an undercoat and topcoat. The coating
thickness ratio of the topcoat to undercoat may vary with the
desired effect and/or the elution system. Typically these are of
different formulations with most or all of the biologically active
material being contained in the undercoat and a non-thrombogenic or
biocompatible non-porous surface found in the topcoat.
[0027] It is desirable that the topcoat be substantially free of
connected pores or porosigens (materials which can elute during
implantation and form pores). The addition of a porous membrane as
a top coat will increase the coating thickness and reduce the
overall drug loading. Also, the release of porosigens into the body
can be undesirable since it introduces additional foreign materials
into the body, which can cause the patient to have adverse
reactions.
[0028] Since in some embodiments the topcoat should be
substantially free of pores, the topcoat should cover less than the
entire surface of the undercoat. Preferably, the topcoat should
cover only about 10% to about 95% of the surface of the
undercoat.
[0029] By partially covering the surface during manufacture or
inducing "breaks" in the topcoat during mounting/implanting of the
coated device, the biologically active material or drug of the
undercoat is permitted to be released from the undercoat through
those parts of the undercoat which are not covered by the
topcoat.
[0030] This mechanism is illustrated by FIG. 9, which shows a
surface of a prosthesis 101 covered by a coating 102 comprising an
undercoat 103 and a topcoat 104. The topcoat 104, which only
partially covers the undercoat 103, leaves certain areas 106 of the
undercoat, including drug particles 105, exposed to body fluids at
the implantation site. It is through these "uncovered" areas 106 of
the undercoat that the drug particles 105 of the undercoat 103 are
allowed to be released into the implantation site.
[0031] Additionally, it is preferred that the topcoats have an
average thickness equivalent to the average particle size of the
drug in the undercoat. Preferably the average thickness is about 1
to 7 microns and more preferable that the topcoat average thickness
be about 1 to 5 microns. Also, the polymer of the topcoat may be
the same as or different from the polymer of the undercoat.
[0032] 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.
[0033] The coating is exposed to room temperature ventilation for a
predetermined time (possibly one hour or more) for solvent vehicle
evaporation. In the case of certain undercoat materials, thereafter
the polymer material is cured at room temperature or elevated
temperatures. 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 induce
physico-chemical changes. Where, for example, polyurethane
thermoplastic elastomers are used as an undercoat material, solvent
evaporation can occur at room temperature rendering the undercoat
useful for controlled drug release without further curing.
[0034] The applicable 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. Urethane
prepolymers can also be utilized. Unlike the polyurethane
thermoplastic elastomers, some of these materials are applied as
prepolymers in the coating composition and must thereafter be heat
cured. The preferred silicone 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 the undercoat or a certain number of lower layers
and the top layers or a single curing step used after coating is
completed.
[0035] The coated stents may thereafter be subjected to a postcure
process which includes an inert gas plasma treatment, and
sterilization which may include gamma radiation, ETO treatment,
electron beam or steam treatment.
[0036] 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 20-50 mTorr. Thereafter, inert gas (argon,
helium or mixture of them) is admitted to the reaction chamber for
the plasma treatment. One method uses argon (Ar) 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.
[0037] In accordance with the invention, the topcoat or surface
coating may be applied in any of several ways to further control
thrombolytic effects and optionally, control the release profile
especially the initial very high release rate associated with the
elution of heparin.
[0038] In one embodiment, an outer layer of fluorosilicone (FSi) is
applied to the undercoat as a topcoat. The outer layer can also
contain heparin. In another embodiment, polyethylene glycol (PEG)
is immobilized on the surface of the coating. In this process, the
underlayer is subjected to inert gas plasma treatment and
immediately thereafter is treated by ammonia (NH3) plasma to
aminate the surface. Amination, as used in this application, means
creating mostly imino groups and other nitro containing species on
the surface. This is followed by immediate immersion into
electrophilically activated polyethylene glycol (PEG) solution with
a reductive agent, i.e., sodium cyanoborohydride.
[0039] To form a topcoat which is substantially free of pores,
porosigens or materials capable of eluting from the topcoat during
implantation, should not be included in the composition used to
form the topcoat. For example, a substantially non-porous topcoat
can be produced from a topcoat composition which comprises a
substantially pure polymeric material. The material preferably
provides biocompatibility to the implanted device during and after
release of the biologically active material.
[0040] To prepare a topcoat which covers less than the entire
surface of the undercoat, a number of methods can be used. For
instance, by controlling the thickness of the topcoat so that it
has an average thickness less than that of the diameter of certain
drug particles, the undercoat will be only partly covered by the
topcoat since some of drug particles will not be covered by the
topcoat.
[0041] Also, a partial topcoat can be formed by using a topcoat
polymer which is incompatible with the undercoat polymer to
generate a microphase separation in the topcoat. Furthermore, to
make a topcoat which covers less than the entire surface of the
undercoat or which only partially covers the undercoat, a poor
solvent wash can be applied to the topcoat, to force the topcoat
polymer to shrink so that the undercoat is not entirely
covered.
[0042] In other embodiments the topcoat can partially or fully
cover the undercoat prior to delivery or implantation of the
device. The topcoat materials can be selected so they have certain
water permeability. When water penetrates the topcoat and into the
drug particles of the undercoat, the water will swell the particles
or dissolve the particles. In either situation, it creates osmotic
pressure in the surrounding coating material of the undercoat. The
pressure then breaks the thinnest part of the topcoat, and leave
the void space in the topcoat for the drug to elute out.
[0043] In another embodiment, the topcoat material has a different
Young's modulus (either while it is wet or dry) than that of the
undercoat material. More specifically, the Young's modulus can be
higher for the topcoat material. During the mounting of the coated
devices onto the delivery device or during deployment of the coated
device, the coating undergoes compression or stretching. Topcoat
materials with higher Young's modulus tend to crack and form void
spaces for the drug to elute from undercoat.
[0044] Another way to form a topcoat is to cover the undercoat with
a bioabsorbable material. In this embodiment, the topcoat can cover
either the entire undercoat or only part of the undercoat before or
after implantation. Upon contact with body fluids, the
bioabsorbable material of the topcoat will degrade. The rate of
degradation depends upon the bioabsorbable material used. When the
topcoat is partially absorbed, the undercoat is exposed to the body
fluid and the drug is released, however the burst effect will be
reduced.
[0045] The coated and cured stents having the modified outer layer
or surface optionally are subjected to a final gamma radiation
sterilization nominally at 2.5-3.5 Mrad. Argon (Ar) plasma treated
stents enjoy full resiliency after radiation whether exposed in a
constrained or non-constrained status, while constrained stents
subjected to gamma sterilization without Ar plasma pretreatment
lose resiliency and do not recover at a sufficient or appropriate
rate where the undercoat is silicone.
[0046] The elastomeric materials that form the stent coating
underlayers should possess certain properties. Preferably the
layers should be of suitable hydrophobic biostable elastomeric
materials which do not degrade. Surface layer or topcoat materials
should minimize tissue rejection and tissue inflammation and permit
encapsulation by tissue adjacent the stent implantation site.
Exposed material is designed to reduce clotting tendencies in blood
contacted and the surface is preferably modified accordingly. Thus,
underlayers of the above materials are preferably provided with a
silicone or a fluorosilicone outer coating layer which should not
contain imbedded bioactive material, such as heparin in order to
avoid the formation of pores in the topcoat. Alternatively, the
outer coating may consist essentially of polyethylene glycol (PEG),
polysaccharides, phospholipids, or combinations of the
foregoing.
[0047] Polymers generally suitable for the undercoats or
underlayers include silicones (e.g., polysiloxanes and substituted
polysiloxanes), polyurethanes, thermoplastic elastomers in general,
ethylene vinyl acetate copolymers, polyolefin elastomers, polyamide
elastomers, and EPDM rubbers. The above-referenced materials are
considered hydrophobic with respect to the contemplated environment
of the invention. Surface layer or topcoat materials can include
the same polymer as that of the undercoat. Examples of suitable
polymers include without limitation fluorosilicones and
polyethylene glycol (PEG), polysaccharides, phospholipids, and
combinations of the foregoing.
[0048] While heparin is preferred as the incorporated active
material, agents possibly suitable for incorporation include
antithrobotics, anticoagulants, antibiotics antiplatelet agents,
thrombolytics, antiproliferatives, steroidal and nonsteroidal
antiinflammatories, agents that inhibit hyperplasia and in
particular restenosis, smooth muscle cell inhibitors, 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).
[0049] Suitable 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 U.S. Pat. No. 5,061,275 to Wallsten et al., which are
incorporated by reference herein.
[0050] 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.
[0051] 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 the material released traverses through another
material.
[0052] 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
denser crosslink structure will result in a longer release time and
reduced burst effect.
[0053] It will also be appreciated that an unmedicated silicone
thin top layer provides some advantage and additional control over
drug elution.
BRIEF DESCRIPTION OF THE DRAWINGS
[0054] In the drawings, wherein like numerals designate like parts
throughout the same:
[0055] FIG. 1 is a schematic flow diagram illustrating the steps of
the process of the invention;
[0056] FIG. 2 represents a release profile for a multi-layer system
showing the percentage of heparin released over a two-week
period;
[0057] FIG. 3 represents a release profile for a multi-layer system
showing the relative release rate of heparin over a two-week
period;
[0058] 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 without associated
means to provide a long term non-thrombogenic surface
thereafter;
[0059] FIG. 5 illustrates drug elution kinetics at a given loading
of heparin over a two-week period at different coating thicknesses
without associated means to provide a long term non-thrombogenic
surface thereafter;
[0060] FIG. 6 illustrates the release kinetics for a given
undercoat and topcoat material varied according to thickness in
which the percentage heparin in the undercoat and topcoats are kept
constant;
[0061] FIG. 7 is a plot of heparin release kinetics in phosphate
buffer system at pH 7.4 with and without fluorosilicone (FSi)
topcoat; and
[0062] FIG. 8 is another plot of heparin release kinetics in
phosphate buffer system in which a topcoat containing
fluorosilicone (FSi) only is compared with an FSi topcoat
containing 16.7% imbedded heparin.
[0063] FIG. 9 illustrates a surface of an implantable prosthesis
covered with an undercoat, containing a biologically active
material, which is partly covered by a topcoat.
DETAILED DESCRIPTION
[0064] 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 (if any)
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.
[0065] General Preparation of Heparin Undercoating Composition
[0066] 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 tetrahydrofuran (THF) HPCL grade (Aldrich or EM) was
added. For a 37.5% heparin coating, having 35% solids,
W.sub.silicone=5 g and W.sub.hep=5.times.0.35.times.0.375/(0.-
625)=1.05 g were used. 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 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.
[0067] General Preparation of Dexamethasone Undercoating
Composition
[0068] 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 and 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.
[0069] 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, at 16 in FIG. 1. 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) was used.
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.
[0070] 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.
[0071] 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, typical tie-layers as at 18 in FIG. 1,
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.
[0072] 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.
[0073] Several applied layers make up what is called the undercoat
as at 18. In one process, additional upper undercoat layers, which
may be of the same or different composition with respect to
bioactive material, the matrix. polymeric materials and
crosslinking agent, for example, may be applied as the top layer as
at 20. The application of the top layer follows the same coating
procedure as the undercoat with the number and thickness of layers
being optional. of course, the thickness of any layer can be
adjusted by adjusting the speed of rotation of the stent and the
spraying conditions. Generally, the total coating thickness is
controlled by the number of spraying cycles or thin coats which
make up the total coat.
[0074] As shown at 22 in FIG. 1, the coated stent is thereafter
subjected to a curing step in which the prepolymer 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.
[0075] 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.
[0076] 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.
[0077] 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.
[0078] After this, as shown at 26, the stents may be exposed to
gamma sterilization at 2.5-3.5 Mrad. The radiation may be carried
out with the stent in either the radially unconstrained status or
in the radially constrained status.
[0079] Preferably, however, the surface is modified prior to plasma
treatment or just prior to sterilization by one of several
additional processing methods of which some are described in
relation to the following examples.
EXAMPLE 1
Fluorosilicone Surface Treatment of Eluting Heparin Coating
[0080] The undercoat of a stent was coated as multiple applied
layers as described above thereafter and cured as described at 22.
The heparin content of the undercoat was 37.5% and the coating
thickness was about 30-40.mu.. Fluorosilicone (FSi) spray solution
was prepared at 30 from a fluorosilicone suspension (Applied
Silicone #40032) by weighing an amount of fluorosilicone suspension
and adding tetrahydrofuran (THF) according to the relation equation
of V.sub.THF=1.2.times. the weight of fluorosilicone suspension.
The solution was stirred very well and spray-coated on the stent at
32 using the technique of the application of the undercoat process
at 18 and the coated stents were cured at 90.degree. C. for 16
hours. The coated stents are argon plasma treated prior to gamma
sterilization according to the procedures described above in
accordance with steps 22-26.
[0081] FIG. 7 is a plot of heparin release kinetics in phosphate
buffer system with fluorosilicone topcoat and without any topcoat.
The thickness of the topcoat is about 10-15 A. While it does not
appear on the graph of FIG. 7 it should be noted that the release
rate for the coating without FSi is initially about 25 times higher
than that with FSi, i.e., during the first 2 hours. This is, of
course, clearly off the scale of the graph. It is noteworthy,
however, that the coating with the FSi top layer or diffusion
barrier does show a depressed initial release rate combined with an
enhanced elution rate after the first day and through the first
week up until about the tenth day. In addition, the fluorosilicone
(FSi) topcoat, by virtue of the high electronegativity of
fluorination maintains non-thrombogenic surface qualities during
and after the elusion of the biologically active heparin species.
In addition, because of the negative charges off the heparin
itself, the electro-negativity of the fluorosilicone topcoat may
be, at least in part, responsible for the modified heparin release
kinetic profile.
[0082] FIG. 8 compares a plot of fluorosilicone (FSi) top coating
containing 16.7% imbedded heparin with one containing
fluorosilicone (FSi) only. An undercoating is identical to that
utilized in FIG. 7 containing about 37.5% heparin to a thickness of
about 30-40 microns. These elution kinetics are quite comparable
with the heparin-free FSi top layer greatly reducing the initial
burst of heparin release and otherwise the heparin in the FSi top
layer imparts a slightly greater release over the period of the
test.
EXAMPLE 2
Immobilization of Polyethylene Glycol (PEG) on Drug Eluting
Undercoat
[0083] An undercoat was coated on a stent and cured at 22 as in
Example 1. The stent was then treated by argon gas plasma as at 24
and ammonium gas plasma at 40. The equipment and the process of
argon gas plasma treatment was as has been described above. The
ammonium plasma treatment was implemented immediately after the
argon gas plasma treatment, to aminate the surface of the coating.
The ammonium flow rate was in the range of 100-700 cubic centimeter
per minute (ccM) in preferably in the range of 500-600 ccM. The
power output of radio frequency plasma was in the range of 50-500
watts, preferably in .about.200 watts. The process time was in the
range of 30 sec-10 min, preferably .about.5 min.
[0084] Immediately after amination, the stents were immersed into
electrophilically activated polyethylene glycol (PEG) solution it
42. PEG is known to be an inhibitor of protein absorption. Examples
of electrophilically activated PEG are PEG nitrophenyl carbonates,
PEG trichlorophenyl carbonates, PEG tresylate, PEG glycidyl ether,
PEG isocyanate, etc., optionally with one end terminated with
methoxyl group. Molecular weight of PEG ranged from about
1000-6000, and is preferable about 3000. It has been observed that
simple ammonium amination will not generate large quantities of
primary and secondary amines on the elastomeric polymer surface
(for example silicone). Instead, imine (>C.dbd.N--H), and other
more oxidative nitro containing groups will dominate the surface.
It is generally necessary to add reductive agent, such as
NaBH.sub.3CN into the reaction media so that the functional group
on PEG can react with imine and possibly other nitro-containing
species on the surface, and therefore immobilize PEG onto the
surface. The typical concentration of NaBH.sub.3CN is about 2
mg/ml. Since PEG and its derivatives dissolve in water and many
polar and aromatic solvents, the solvent used in the coating must
be a solvent for PEG but not for the drug in the undercoat to
prevent the possible loss of the drug through leaching. In the case
of eluting-heparin coating, a mixed solvent of formamide and methyl
ethyl ketone (MEK) or a mixed solvent of formamide and acetone are
preferred solvents, (preferably at ratios of 30 formamide: 70 MEK
or acetone by volume), since they will not dissolve heparin. The
concentration of PEG, the reaction time, the reaction temperature
and the pH value depend on the kind of PEG employed. In the case of
eluting heparin coating, 5% PEG tresylate in (30-70) Formamide/MEK
was used successfully. The reaction time was 3 hours at room
temperature. PEG was then covalently bound to the surface. Gamma
radiation was then used for sterilization of this embodiment as
previously described.
[0085] With respect to the anticoagulant material heparin, the
percentage in the undercoat is nominally from about 30-50% and that
of the topcoat from about 0-30% active material. The coating
thickness ratio of the topcoat to the undercoat varies from about
1:10 to 1:2 and is preferably in the range of from about 1:6 to
1:3.
[0086] To produce a topcoat which is substantially free of pores,
materials such as porosigens, which can be removed or leached out
of the topcoat should not be included in the composition used to
form the topcoat. One way of preparing a substantially non-porous
topcoat is to apply a topcoat composition which comprises
substantially pure polymeric materials. These materials preferably
impart biocompatibility to the implanted device during and after
the release of the biologically active material.
[0087] A topcoat which only partially covers the undercoat can be
formed in a number of ways. Such methods include controlling the
thickness of the topcoat so that it is less than the diameter of
certain drug particles in the undercoat. For example, when a drug
used in the undercoat has an average particle size of 5 .mu.m, it
is possible that 15% of the particles will be greater than or equal
to 8 .mu.m. At the molecular level, the surface unevenness is at
least more than 5 microns. By application of a topcoat of about 5
micron or less. The uneven surface will become smooth to a certain
degree. But uncovered areas will still exist, which allows water to
penetrate into the undercoat, swelling the drug particle, or
dissolving the drug. Due to the osmotic pressure, the drug will
elute out through the uncovered areas. If the osmotic pressure is
too high, cracks or voids may form in the topcoat which allows
drugs to elute from the undercoat to the body.
[0088] As illustrated in FIG. 9, in a coating 102 which covers a
surface of a prosthesis 101, a thin topcoat 104 of the coating 102
only partially covers the undercoat 103 of the coating 102. By
having the topcoat 104 cover less than the entire surface of the
undercoat 103, parts of the undercoat 103, including a number of
drug particles 105, are exposed to body fluids at the implantation
site so that the drug 105 can be released. By using a topcoat 104
having a thickness which is about the average particle size of the
drug 105, certain larger sized drug particles 106 will not be
covered by the topcoat 104.
[0089] Other methods to form a partially covered topcoat include
using a polymer which is incompatible to the undercoat elastomer as
the biocompatible material of the topcoat. Because of the
incompatibility between the materials, a microphase separation will
form in the topcoat which will leave the undercoat partially
uncovered. Persons skilled in the art are aware of suitable
combinations of such incompatible materials.
[0090] Another method involves applying a poor solvent wash to the
topcoat to force the biocompatible polymer to shrink and create
uneven surfaces voids in the topcoat which forms a topcoat which
partially covers the undercoat.
[0091] In yet another method of making a topcoat which partly or
fully covers the undercoat, topcoat and undercoat materials having
different Young's moduli (either before or after they have cured)
are used. For instance, a topcoat material having a higher Young's
modulus compared to that of the undercoat material can be employed.
When the coated device is mounted on the delivery device or during
deployment of the coated device, the topcoat undergoes compression
or stretching or other types of mechanical challenges. Since the
topcoat material has a higher Young's modulus, it will tend to
crack and form voids in the topcoat to allow the drug of the
undercoat to elute therefrom.
[0092] Another method for forming the topcoat involves using a
bioabsorbable material in the topcoat, which can cover the entire
undercoat or only a part of the undercoat. When the coated device
comes into contact with body fluid, the topcoat begins to degrade
either at the surface or throughout the bulk of the topcoat. The
rate of degradation depends upon the type of bioabsorbable material
used. Once the topcoat has been partially absorbed, the undercoat
is exposed to body tissue and the drug in the undercoat is
released, but the burst release or effect is reduced.
[0093] 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 subacute
thrombosis particularly if drugs such as heparin are
incorporated.
[0094] 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 topcoat
as shown by the same drug loading and similar undercoat thickness
in FIG. 6.
[0095] What is apparent from the data gathered to date, however, is
that the process of the present invention enables the drug elution
kinetics to be controlled in a manner desired to meet the needs of
the particular stent application. In a similar manner, stent
coatings can be prepared using a combination of two or more drugs
and the drug release sequence and rate controlled. For example,
antiproliferation drugs may be combined in the undercoat and
antiplatelet drugs in the topcoat. In this manner, the antiplatelet
drugs, for example, heparin, will elute first followed by
antiproliferation drugs to better enable safe encapsulation of the
implanted stent.
[0096] 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.
[0097] 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.
[0098] 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%.
[0099] 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.
[0100] 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.
[0101] 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.
[0102] 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.
[0103] 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.
* * * * *