U.S. patent application number 09/780513 was filed with the patent office on 2001-07-26 for medical device for delivering a therapeutic substance and method therefor.
Invention is credited to Berg, Eric P., Dinh, Thomas Q., Wolff, Rodney G..
Application Number | 20010009688 09/780513 |
Document ID | / |
Family ID | 22090415 |
Filed Date | 2001-07-26 |
United States Patent
Application |
20010009688 |
Kind Code |
A1 |
Dinh, Thomas Q. ; et
al. |
July 26, 2001 |
Medical device for delivering a therapeutic substance and method
therefor
Abstract
A device useful for localized delivery of a therapeutic material
is provided. The device includes a structure including a porous
material; and a water-insoluble salt of a therapeutic material
dispersed in the porous material. The water-insoluble salt is
formed by contacting an aqueous solution of a therapeutic salt with
a heavy metal water-soluble salt dispersed throughout a substantial
portion of the porous material. The porous material can be made of
a polymer other than fibrin with fibrin incorporated into the
pores, which can be the only layer of polymeric material on the
medical device (e.g., stent). A new method for preparing a porous
polymer material on a medical device.
Inventors: |
Dinh, Thomas Q.;
(Minnetonka, MN) ; Wolff, Rodney G.; (Minnetonka
Beach, MN) ; Berg, Eric P.; (Plymouth, MN) |
Correspondence
Address: |
Medtronic AVE, Inc.
Legal Department
3576 Unocal Place
Santa Rosa
CA
95403
US
|
Family ID: |
22090415 |
Appl. No.: |
09/780513 |
Filed: |
February 12, 2001 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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09780513 |
Feb 12, 2001 |
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09397678 |
Sep 16, 1999 |
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6187370 |
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09397678 |
Sep 16, 1999 |
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09069659 |
Apr 29, 1998 |
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6013099 |
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Current U.S.
Class: |
427/2.1 ;
427/430.1; 427/5 |
Current CPC
Class: |
A61F 2/86 20130101; A61N
5/1002 20130101; A61F 2250/0067 20130101; A61F 2/88 20130101 |
Class at
Publication: |
427/2.1 ;
427/430.1; 427/5 |
International
Class: |
B05D 001/18; A61L
002/00; A61M 036/14 |
Claims
We claim:
1. A medical device having at least one blood-contacting surface
comprising: a non-fibrin porous material having fibrin incorporated
therein; and a water-insoluble therapeutic salt dispersed in at
least a portion of the porous material.
2. The medical device of claim 1 wherein the non-fibrin porous
material comprises a film.
3. The medical device of claim 2 wherein the non-fibrin porous
material is selected from the group consisting of silicone,
polyurethane, polysulfone, cellulose, polyethylene, polypropylene,
polyamide, polyester, polytetrafluoroethylene, and a combination of
two or more of these materials.
4. The medical device of claim 2 wherein the water-insoluble salt
comprises a salt of an antithrombotic material.
5. The medical device of claim 4 wherein the antithrombotic
material is heparin.
6. The medical device of claim 5 wherein the water-insoluble salt
is selected from the group consisting of barium, calcium and silver
salts of the antithrombotic material.
7. An intravascular stent comprising a generally cylindrical stent
body and a single layer of a polymeric film; wherein the single
layer of a polymeric film comprises a non-fibrin porous polymer
with fibrin incorporated therein.
8. The stent of claim 7 wherein the non-fibrin porous polymer film
comprises a material selected from the group of silicone,
polyurethane, polysulfone, cellulose, polyethylene, polypropylene,
polyamide, polyester, polytetrafluoroethylene, and a combination of
two or more of these materials.
9. The stent of claim 8 wherein the non-fibrin porous polymer film
comprises a polyurethane.
10. The stent of claim 7 further comprising a water-insoluble
therapeutic salt dispersed in at least a portion of the porous
polymer film with fibrin incorporated therein.
11. The stent of claim 10 wherein the antithrombotic salt is a salt
of heparin.
12. The stent of claim 11 wherein the water-insoluble heparin salt
is selected from the group consisting of silver, barium and calcium
salts.
13. A method for coating a medical device with a porous polymer,
the method comprising: placing the medical device in a mold;
placing a solution of a polymer in the mold with the medical
device; wherein the solution of the polymer includes a solvent
capable of phase separating from the polymer at a temperature below
the freezing point of the solvent; cooling the solution of the
polymer in the mold to a temperature below the freezing point of
the solvent until a first fraction of particulate material is
formed by solidification and phase separation of the solvent from
the polymer and is dispersed within solidified polymer; cooling the
solution further and at a faster rate than in the first cooling
step to form a second fraction of particulate material dispersed
within the solidified polymer, wherein the second fraction of
particulate material has a smaller particle size than the first
fraction, and removing the particulate material from the polymer to
form pores therein.
14. The method of claim 13 wherein the medical device is a
stent.
15. The method of claim 13 wherein the solution comprises ag
polyurethane and dioxane.
16. The method of claim 13 wherein the porous polymer is in the
form of a film.
17. The method of claim 13 wherein the porous polymer is in the
form of a coating.
Description
BACKGROUND OF THE INVENTION
[0001] This invention relates to a medical device employing a
therapeutic substance as a component thereof. For example, in an
arterial site treated with percutaneous transluminal coronary
angioplasty therapy for obstructive coronary artery disease a
therapeutic antithrombogenic substance such as heparin may be
included with a device and delivered locally in the coronary
artery. Also provided is a method for making a medical device
capable of localized application of therapeutic substances. This
invention also relates to a medical device, particularly a stent,
having a porous polymeric film with fibrin incorporated therein for
enhanced biocompatibility, with or without a therapeutic substance
as a component thereof.
[0002] Medical devices which serve as substitute blood vessels,
synthetic and intraocular lenses, electrodes, catheters and the
like in and on the body or as extracorporeal devices intended to be
connected to the body to assist in surgery or dialysis are well
known. For example, intravascular procedures can bring medical
devices into contact with the patient's vasculature. In treating a
narrowing or constriction of a duct or canal percutaneous
transluminal coronary angioplasty (PTCA) is often used with the
insertion and inflation of a balloon catheter into a stenotic
vessel. Other intravascular invasive therapies include atherectomy
(mechanical systems to remove plaque residing inside an artery),
laser ablative therapy and the like. However, this use of
mechanical repairs can have adverse consequences for the patient.
For example, restenosis at the site of a prior invasive coronary
artery disease therapy can occur. Restenosis, defined
angiographically, is the recurrence of a 50% or greater narrowing
of a luminal diameter at the site of a prior coronary artery
disease therapy, such as a balloon dilatation in the case of PTCA
therapy. In particular, an intra-luminal component of restenosis
develops near the end of the healing process initiated by vascular
injury, which then contributes to the narrowing of the luminal
diameter. This phenomenon is sometimes referred to as "intimal
hyperplasia." It is believed that a variety of biologic factors are
involved in restenosis, such as the extent of the injury,
platelets, inflammatory cells, growth factors, cytokines,
endothelial cells, smooth muscle cells, and extracellular matrix
production, to name a few.
[0003] Attempts to inhibit or diminish restenosis often include
additional interventions such as the use of intravascular stents
and the intravascular administration of pharmacological therapeutic
agents. Examples of stents which have been successfully applied
over a PTCA balloon and radially expanded at the same time as the
balloon expansion of an affected artery include the stents
disclosed in U.S. Pat. No. 4,733,665 issued to Palmaz, U.S. Pat.
No. 4,800,882 issued to Gianturco and U.S. Pat. No. 4,886,062
issued to Wiktor. Also, such stents employing therapeutic
substances such as glucocorticoids (e.g. dexamethasone,
beclamethasone), heparin, hirudin, tocopherol, angiopeptin,
aspirin, ACE inhibitors, growth factors, oligonucleotides, and,
more generally, antiplatelet agents, anticoagulant agents,
antimitotic agents, antioxidants, antimetabolite agents, and
anti-inflammatory agents have been considered for their potential
to solve the problem of restenosis. Such substances have been
incorporated into a solid composite with a polymer in an adherent
layer on a stent body with fibrin in a separate adherent layer on
the composite to form a two layer system. The fibrin is optionally
incorporated into a porous polymer layer in this two layer
system.
[0004] Another concern with intravascular and extracorporeal
procedures is the contact of biomaterials with blood which can
trigger the body's hemostatic process. The hemostatic process is
normally initiated as the body's response to injury. When a vessel
wall is injured, platelets adhere to damaged endothelium or exposed
subendothelium. Following adhesion of the platelets, these cells
cohere to each other preparatory to aggregation and secretion of
their intracellular contents. Simultaneously there is activation,
probably by electrostatic charge of the contact factors, of the
coagulation cascade. The sequential step-wise interaction of these
procoagulant proteins results in the transformation of soluble
glycoproteins into insoluble polymers, which after transamidation
results in the irreversible solid thrombus.
[0005] Immobilization of polysaccharides such as heparin to
biomaterials has been used to improve bio- and hemocompatibility of
implantable and extracorporeal devices. The mechanism responsible
for reduced thrombogenicity of heparinized materials is believed to
reside in the ability of heparn to speed up the inactivation of
serine proteases (blood coagulation enzymes) by AT-III. In the
process, AT-III forms a complex with a well defined pentasaccharide
sequence in heparin, undergoing a conformational change and thus
enhancing the ability of AT-III to form a covalent bond with the
active sites of serine proteases such as thrombin. The formed
TAT-complex then releases from the polysaccharide, leaving the
heparin molecule behind for a second round of inactivation.
[0006] Usually, immobilization of heparin to a biomaterial surface
consists of activating the material in such a way that coupling
between the biomaterial and functional groups on the heparin
(--COOH, --OH, --NH.sub.2) can be achieved. For example, Larm
presented (in U.S. Pat. No. 4,613,665) a method to activate heparin
via a controlled nitrous acid degradation step, resulting in
degraded heparin molecules of which a part contains a free terminal
aldehyde group. Heparin in this form can be covalently bound to an
aminated surface in a reductive amination process. Although the
molecule is degraded and as a result shows less catalytic activity
in solution, the end point attachment of this type of heparin to a
surface results in true anti-thrombogenicity due to the proper
presentation of the biomolecule to the surface. In this fashion,
the molecule is freely interacting with AT-III and the coagulation
enzymes, preventing the generation of thrombi and microemboli.
[0007] However, the attachment and delivery of therapeutic
substances such as heparin can involve complicated and expensive
chemistry. It is therefore an object of the present invention to
provide a medical device having a biocompatible, blood-contacting
surface with an active therapeutic substance at the surface and a
simple, inexpensive method for producing such a surface. It is also
an object of the present invention to provide a medical device
having a porous material with fibrin incorporated therein,
optionally with an active therepeutic substance at the
blood-contacting surface. It is also a further object of the
present invention to provide a medical device, such as an
intravascular stent, having a porous polymeric film adhered to the
medical device body with fibrin incorporated therein for enhanced
biocompatibility.
SUMMARY OF THE INVENTION
[0008] This invention relates to a medical device having a
blood-contacting surface with a therapeutic substance thereon.
Preferably, the device according to the invention is capable of
applying a highly localized therapeutic material into a body lumen
to treat or prevent injury. The term "injury" means a trauma, that
may be incidental to surgery or other treatment methods including
deployment of a stent, or a biologic disease, such as an immune
response or cell proliferation caused by the administration of
growth factors. In addition, the methods of the invention may be
performed in anticipation of "injury" as a prophylactic. A
prophylactic treatment is one that is provided in advance of any
symptom of injury in order to prevent injury, prevent progression
of injury or attenuate any subsequent onset of a symptom of such
injury.
[0009] In accordance with the invention, a device for delivery of
localized therapeutic material includes a structure including a
porous material and a plurality of discrete particles of a
water-insoluble salt of the therapeutic material dispersed
throughout a substantial portion of the porous material.
Preferably, the device is capable of being implanted in a body so
that the localized therapeutic agent can be delivered in vivo,
typically at a site of vascular injury or trauma. More preferably,
the porous material is also biocompatible, sufficiently
tear-resistant and nonthrombogenic.
[0010] The porous material may be a film on at least a portion of
the structure or the porous material may be an integral portion of
the structure. Preferably, the porous material is selected from the
group of a natural hydrogel, a synthetic hydrogel, TEFLON
(polytetrafluoroethylene- ), silicone, polyurethane, polysulfone,
cellulose, polyethylene, polypropylene, polyamide, polyester, and a
combination of two or more of these materials. Examples of natural
hydrogels include fibrin, collagen, elastin, and the like.
[0011] Alternatively, the porous material may have fibrin
incorporated therein. Although this material preferably has a
therapeutic agent also incorporated therein, this is not necessary
for enhanced biocompatibility. Thus, in one embodiment, the present
invention provides a medical device, preferably, an intravascular
stent, that includes a porous polymer film with fibrin incorporated
within the pores, optionally with a therapeutic substance also
incorporated within the pores.
[0012] The therapeutic agent preferably includes an antithrombotic
material. More preferably, the antithrombotic material is a heparin
or heparin derivative or analog. Also preferably, the insoluble
salt of the therapeutic material is one of the silver, barium or
calcium salts of the material.
[0013] The structure of the device can be adapted for its intended
extracorporeal or intravascular purpose in an internal human body
site, such as an artery, vein, urethra, other body lumens,
cavities, and the like or in an extracorporeal blood pump, blood
filter, blood oxygenator or tubing. In one aspect of the invention,
the shape is preferably generally cylindrical, and more preferably,
the shape is that of a catheter, a stent, or a guide wire.
[0014] In another aspect of the invention, an implantable device
capable of delivery of a therapeutic material includes a structure
comprising a porous material; and a plurality of discrete particles
comprising a heavy metal water-soluble salt dispersed throughout a
substantial portion of the porous material. Preferably, the heavy
metal water-soluble salt is selected from the group of AgNO.sub.3,
Ba(NO.sub.3).sub.2, BaCl.sub.2, and CaCl.sub.2. The amount of
water-soluble salt dispersed throughout a portion of the porous
material determines the total amount of therapeutic material that
can be delivered once the device is implanted.
[0015] The invention provides methods for manufacturing medical
devices. Specifically, the invention provides a method for coating
a medical device with a porous polymer (film or coating). The
method includes: placing the medical device in a mold; placing a
solution of a polymer in the mold with the medical device; wherein
the solution of the polymer includes a solvent capable of phase
separating from the polymer at a temperature below the freezing
point of the solvent; cooling the solution of the polymer in the
mold to a temperature below the freezing point of the solvent until
a first fraction of particulate material is formed by
solidification and phase separation of the solvent from the polymer
and is dispersed within solidified polymer; cooling the solution
further and at a faster rate than in the first cooling step to form
a second fraction of particulate material dispersed within the
solidified polymer, wherein the second fraction of particulate
material has a smaller particle size than the first fraction; and
removing the particulate material from the polymer to form pores
therein. Preferably, the medical device is a stent and the solution
includes a polyurethane dissolved in dioxane.
[0016] The invention also provides methods for making an
implantable device which includes therapeutic materials. In one
embodiment, a method of the invention includes loading a structure
comprising a porous material with a heavy metal water-soluble salt
dispersed throughout a substantial portion of the porous material,
sterilizing the loaded structure, and packaging for storage and,
optionally, delivery of the sterilized loaded structure.
Preferably, the method of the invention further includes
substantially contemporaneously loading of a water soluble
therapeutic material, wherein a water insoluble salt of the
therapeutic material is produced throughout a substantial portion
of the porous material of the structure. "Substantially
contemporaneously," means that the step of loading a water soluble
therapeutic material occurs at or near a step of positioning the
device proximate to a desired area, i.e., at or near the surgical
arena prior to administration to or implantation in, a patient.
More preferably, the water insoluble salt of the therapeutic
material is dispersed throughout a substantial portion of the
porous material.
[0017] In another aspect of the invention, a method includes
loading a structure comprising a porous material with a heavy metal
water-soluble salt dispersed throughout a substantial portion of
the porous material; loading a water soluble therapeutic material,
wherein a water insoluble salt of the therapeutic material is
produced in a substantial portion of the porous material of the
structure; and packaging for delivery of the loaded structure.
[0018] Thus, the methods for making an implantable device to
deliver a therapeutic material and device in vivo, or in an
extracorporeal circuit in accordance with the invention, are
versatile. A therapeutic material may be loaded onto a structure
including a porous material at any number of points between, and
including, the point of manufacture and the point of use. As a
result of one method, the device can be stored and transported
prior to incorporation of the therapeutic material. Thus, the end
user can select the therapeutic material to be used from a wider
range of therapeutic agents.
BRIEF DESCRIPTION OF THE DRAWINGS
[0019] FIG. 1 is an elevational view of one embodiment of a device
according to the invention with a balloon catheter as a mode of
delivery of the device;
[0020] FIG. 2 is an elevational view of another embodiment of a
device according to the invention with a balloon catheter as a mode
of delivery of the device;
[0021] FIG. 3 is a flow diagram schematically illustrating methods
according to the invention;
[0022] FIG. 4 is a photograph taken from a scanning electron
microscope of a surface showing the insoluble therapeutic material
according to the invention;
[0023] FIGS. 5a and 5b are photographs showing the histological
comparison between a stent heparinized according to the present
invention (5a) and a control stent (5b); and
[0024] FIGS. 6 and 7 are photographs showing different
magnifications of an artery wall with an expanded stent therein
having a porous polymer film with fibrin incorporated therein.
DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS
[0025] One of the more preferred configurations for a device
according to the invention is a stent for use in artery/vascular
therapies. The term "stent" refers to any device capable of being
delivered by a catheter and which, when placed into contact with a
portion of a wall of a lumen to be treated, will also deliver
localized therapeutic material at a luminal or blood-contacting
portion of the device. A stent typically includes a lumen
wall-contacting surface and a lumen-exposed surface. Where the
stent is shaped generally cylindrical or tube-like, including a
discontinuous tube or ring-like structure, the lumen-wall
contacting surface is the surface in close proximity to the lumen
wall whereas the lumen-exposed surface is the inner surface of the
cylindrical stent. The stent can include polymeric or metallic
elements, or combinations thereof, onto which a porous material is
applied. For example, a deformable metal wire stent is useful as a
stent framework of this invention, such as that described in U.S.
Pat. No. 4,886,062 to Wiktor, which discloses preferred methods for
making a wire stent. Other metallic stents useful in this invention
include those of U.S. Pat. No. 4,733,655 to Palmaz and U.S. Pat.
No. 4,800,882 to Gianturco.
[0026] Referring now to FIG. 1, the stent 20 comprises a stent
framework 22 and a porous material coating 24. The stent framework
22 is deformable and can be formed from a polymeric material, a
metal or a combination thereof. A balloon 15 is positioned in FIG.
1 adjacent the lumen-exposed surface of the stent to facilitate
delivery of the stent. The stent 20 can be modified to increase or
to decrease the number of wires provided per centimeter in the
stent framework 22. Similarly, the number of wire turns per
centimeter can also be modified to produce a stiffer or a more
flexible stent framework.
[0027] Polymeric stents can also be used in this invention. The
polymers can be nonbioabsorbable or bioabsorbable in part, or
total. Stents of this invention can be completely nonbioabsorbable,
totally bioabsorbable or a composite of bioabsorbable polymer and
nonabsorbable metal or polymer. For example, another stent suitable
for this invention includes the self-expanding stent of resilient
polymeric material as disclosed in International Publication No. WO
91/12779.
[0028] Nonbioabsorbable polymers can be used as alternatives to
metallic stents. The stents of this invention should not
substantially induce inflammatory and neointimal responses.
Examples of biostable nonabsorbable polymers that have been used
for stent construction with or without metallic elements include
polyethylene terephthalate (PET), polyurethane urea and silicone
(for example, see van Beusekom et al., Circulation, 86(supp.
I):I-731, 1992 and Lincoff et al., J. Am. Coll Cardiol., 21(supp.
1):335A, 1994. Although the porous material is shown as a coating
24, it is to be understood that, for the purposes of this
invention, the porous material can be incorporated into the
material of the stent.
[0029] Referring to FIG. 2, an alternative stent 30 is shown. The
stent framework 34 is affixed with a film of a porous material 32.
This can be accomplished by wrapping the film 32 around the stent
framework 34 and securing the film 32 to the framework 34 (i.e.,
the film is usually sufficiently tacky to adhere itself to the
framework but a medical grade adhesive could also be used if
needed) so that the film 32 will stay on the balloon 36 and
framework 34 until it is delivered to the site of treatment. The
film 32 is preferably wrapped over the framework with folds or
wrinkles that will allow the stent 30 to be readily expanded into
contact with the wall of the lumen to be treated. Alternatively,
the film 32 can be molded to the stent framework 34 such that the
framework 34 is embedded within the film 32. Preferably, the film
32 is located on a lumen-wall contacting surface 33 of the stent
framework 34 such that therapeutic material is substantially
locally delivered to a lumen wall, for example, an arterial wall
membrane (not shown).
[0030] Porous Material
[0031] As mentioned above, the device according to the invention is
generally a structure including a porous material. In one
embodiment, the porous material is a film on at least a portion of
the structure. In another embodiment, the porous material is an
integral portion of the structure. Preferably, the porous material
is biocompatible, and sufficiently tear-resistant and
nonthrombogenic. More preferably, the porous material is selected
from the group of a natural hydrogel, a synthetic hydrogel, TEFLON
(polytetrafluoroethylene), silicone, polyurethane, polysulfone,
cellulose, polyethylene, polypropylene, polyamide, polyester, and a
combination of two or more of these materials. Examples of natural
hydrogels include fibrin, collagen, elastin, and the like. In
materials which do not include pores in their usual structural
configurations, pores between one micrometer in diameter or as
large as 1000 micrometers in diameter can be introduced by
conventional means such as by introducing a solvent soluble
particulate material into the desired structure and dissolving the
particulate material with a solvent. However, no particular pore
size is critical to this invention.
[0032] In a preferred embodiment, the porous material is made of a
polymer other than fibrin (i.e., a non-fibrin porous material) and
has fibrin incorporated within the pores. Typically, and
preferably, the porous material is in the form of a sheet material
of a synthetic biostable polymer. Such synthetic biostable polymers
include silicone, polyurethane, polysulfone, cellulose,
polyethylene, polypropylene, polyamide, polyester,
polytetrafluoroethylene, and combinations thereof.
[0033] If the porous material is in the form of a porous sheet or
film, it can be made by a variety of methods. These methods can
include, for example, forming the films using a solid particulate
material that can be substantially removed after the film is
formed, thereby forming pores. By using a solid particulate
material during film formation the size of the pores can, to some
extent, be controlled by the size of the solid particulate material
being used. The particulate material can range from less than about
1 micrometer in diameter to about 1000 micrometers, preferably
about 1 micrometer to about 100 micrometers, more preferably about
25 micrometers to about 60 micrometers. For uniformity of pores,
the particulate material can be screened through successively finer
mesh sieves, e.g., through 100, 170, 270, 325, 400, and 500 mesh
analytical grade stainless steel mesh sieves, to produce a desired
range of particle sizes.
[0034] In one method according to the present invention, a porous
polyurethane sheet material (i.e., film) can be made by dissolving
a polyether urethane in an organic solvent such as
1-methyl-2-pyrrolidone; mixing into the resulting polyurethane
solution a crystalline, particulate material like a salt or sugar
that is not soluble in the solvent; casting the solution with
particulate material into a thin film; and then applying a second
solvent, such as water, to dissolve and remove the particulate
material, thereby leaving a porous sheet. Such a method is
disclosed, for example, in U.S. Pat. Nos. 5,599,352 and 5,591,227,
both issued to Dinh et al. A portion of the particulate material
may remain within the film. As a result, it is preferred that the
solid particulate material be biocompatible.
[0035] Although films for stent bodies according to the present
invention can be manufactured separately from the support structure
of the stent and attached to the support structure after formation,
preferred methods include forming the films directly on the support
structure such that the support structure is at least partially,
preferably completely, encapsulated by the film. In one such method
disclosed in International Publication No. WO 97/07973, a stent is
placed on a mandrel. A particulate material is then applied to the
mandrel and stent such that it is lightly adhered to the mandrel.
The particulate material should be readily soluble in a solvent
which will not also dissolve the polymer chosen for the film. For
example, crystalline sodium bicarbonate is a water soluble material
that can be used as the particulate material. A nonaqueous liquid,
preferably a solvent for the polymer film material, can be applied
to the mandrel before applying the particulate material in order to
retain more of the particulate material on the mandrel. For
example, when a polyurethane is to be used for the film material,
the solvent 1-methyl-2-pyrrolidinone (NMP) can be used to wet the
surface of the mandrel before the application of particulate
material. Preferably, the mandrel is completely dusted with the
particulate in the portions of the mandrel to be coated with the
polymer film. This can be accomplished by dipping the mandrel in
NMP, allowing it to drain vertically for a few seconds and then
dusting the sodium bicarbonate onto the mandrel while rotating it
horizontally until no further bicarbonate particles adhere. Excess
particulate material can be removed by gently tapping the
mandrel.
[0036] Coating with polymer may proceed immediately following
application of the particulate material. A polymer is provided in a
dilute solution and is applied to the particle-coated stent and
mandrel. For example, polyurethane can be dissolved in NMP to make
a 10% solution. Gel particles and particulate impurities can be
removed from the solution by use of a clinical centrifuge. The
polymer solution can be applied by dipping the mandrel into the
solution and letting the solvent evaporate. With the solution of
polyurethane and NMP, a single dip in the solution can provide a
film of adequate thickness. To assist in the formation of
communicating passageways through the polymer between the
blood-contacting surface and the lumen-contacting surface,
additional sodium bicarbonate particles are preferably dusted onto
the polymer solution immediately after the dipping operation and
before the polymer solution has dried. Excess particulate material
can be removed by gently tapping the mandrel. To precipitate and
consolidate the polyurethane film on the stent, it can be dipped
briefly (about 5 minutes) in water and then rolled gently against a
wetted surface, such as a wet paper towel. The stent assembly can
then be placed into one or more water baths over an extended period
(e.g., 8 hours) to dissolve and remove the sodium bicarbonate.
After drying in air at temperatures from about 20.degree. C. to
about 50.degree. C., the film then can be trimmed to match the
contour of the wire.
[0037] In yet another method, a solvent in which the polymer (i.e.,
a film-forming polymer) is soluble that is capable of phase
separating from the polymer at a reduced temperature can be used to
prepare a porous polymer film. In this method, the stent or other
medical device is placed in a cavity of a mold designed for forming
a film around the stent, similar to that disclosed in U.S. Pat. No.
5,510,077 to Dinh et al. A solution of the desired polymer, such as
polyurethane, dissolved in a solvent, such as dioxane, is added to
the mold. The temperature of the solution is then reduced to a
temperature at which the solvent freezes and phase separates from
the polymer as it forms a film, thereby forming particulate
material (i.e., frozen solvent particles) in situ. Typically, for
polyurethane in dioxane, this is a temperature of about -70.degree.
C. to about 3.degree. C. The composition is then immersed in an ice
cold water bath (at about 3.degree. C.) for a few days to allow the
dioxane to dissolve into the ice cold water, thereby forming pores.
The number and size of the pores can be controlled by the
concentration of the polymer and the freezing temperature. A method
similar to this is disclosed in Liu et al., J. Biomed. Mater. Res.,
26, 1489 (1992). This method can be improved on by using a two-step
freezing process described herein. In a first step, the mixture is
cooled slowly to create a first fraction of particulate material
(i.e., frozen solvent particles) dispersed within solidified
polymer. In a second step, the mixture is cooled further (and more
quickly) to create a second fraction of particulate material of
smaller size dispersed within solidified polymer. In this way, a
wider range of pore sizes can be formed with greater control. This
modified method is further described in Example 6.
[0038] Therapeutic Material
[0039] The therapeutic material used in the present invention could
be virtually any therapeutic substance which possesses desirable
therapeutic characteristics and which can be provided in both water
soluble and water insoluble salts and which have bioactivity as an
insoluble salt. For example, antithrombotics, antiplatelet agents,
antimitotic agents, antioxidants, antimetabolite agents,
anti-inflammatory agents and radioisotopes could be used.
"Insoluble salt" or "water insoluble salt" of the therapeutic
substance as set forth herein, means that the salt formed has a
relatively poor solubility in water such that it will not readily
disperse from the pores of the device. In particular, anticoagulant
agents such as heparin, heparin derivatives and heparin analogs
could be used to prevent the formation of blood clots on the
device. Also, water-insoluble radioactive salts such as
Agl.sup.135, BaS.sup.35O.sub.4, and
(Ca).sub.3(P.sup.32O.sub.4).sub.2 could be used for application of
radiotherapy to a body lumen or blood.
[0040] Preferably, the water-insoluble salt of the therapeutic
material is formed by a heavy metal water-soluble salt interacting
with an aqueous solution of the therapeutic material. In the
present invention, the heavy metal water-soluble salt is dispersed
throughout a substantial portion of the porous material.
Preferably, the heavy metal water-soluble salt is selected from the
group of AgNO.sub.3, Ba(NO.sub.3).sub.2, BaCl.sub.2, CaCl.sub.2,
and a mixture thereof. The amount of water-soluble salt dispersed
throughout a portion of the porous material determines the ultimate
amount of therapeutic material capable of being administered once
the device is implanted.
[0041] Fibrin and Methods of Incorporation into Porous Polymer
[0042] The term "fibrin" herein means naturally occurring polymer
of fibrinogen that arises during blood coagulation. It is an
insoluble, crosslinked polymer generated by the action of thrombin
on fibrinogen. Fibrinogen has three pairs of polypeptide chains
(ALPHA 2-BETA 2-GAMMA 2) covalently linked by disulfide bonds with
a total molecular weight of about 340,000. Fibrinogen is converted
to fibrin through proteolysis by a fibrinogen-coagulating protein,
such as thrombin, reptilase, or ancrod.
[0043] Methods of making fibrin and forming it into implantable
devices are well known in the art. See, for example, U.S. Pat. Nos.
4,548,736 (Muller et al.) and 3,523,807 (Gerendas), and European
Patent Application 0 366 564. In one method, fibrin is formed by
contacting fibrinogen with a fibrinogen-coagulating protein, such
as thrombin, reptilase, or ancrod. Preferably, the fibrinogen and
fibrinogen-coagulating protein (e.g., thrombin) used to make fibrin
is from the same animal or human species as that in which the
medical device (e.g., stent) of the present invention will be
implanted in order to avoid cross-species immune reactions. The
resulting fibrin can also be subjected to heat treatment at about
150.degree. C. for 2 hours in order to reduce or eliminate
antigenicity.
[0044] Preferably, the fibrinogen used to make the fibrin is a
bacteria-free and virus-free fibrinogen such as that described in
U.S. Pat. No. 4,540,573 (Neurath et al.). The fibrinogen is
preferably used in solution at a concentration of at least about 10
mg/ml, more preferably at least about 26 mg/ml, and no greater than
about 50 mg/ml. The pH of the solution is preferably about 5.8 to
about 9.0 with an ionic strength of about 0.05 to about 0.45. The
fibrinogen solution can include pure fibrinogen, although
preferably the solution also includes proteins and enzymes such as
albumin, fibronectin, Factor XIII, plasminogen, antiplasmin,
Antithrombin HI, and the like. Most preferably, the fibrinogen is
cryoprecipitated fibrinogen, which can include hundreds of proteins
and enzymes, as disclosed in Spotmitz et al., The American Surgeon,
53, 460-462 (1987).
[0045] The fibrinogen solution also preferably includes a
fibrinogen-coagulating protein, such as thrombin. Alternatively,
however, the fibrinogen-coagulating protein solution can be applied
after the fibrinogen has been applied. This fibrinogen-coagulating
protein solution may or may not include fibrinogen. The thrombin is
preferably used in solution at a concentration of at least about 1
NIH unit/ml, and no greater than about 120 NIH units/ml. Calcium
ions may also be present in the thrombin solution to enhance
mechanical properties and biostability of the device. If used, they
are preferably present in a concentration of about 0.02 M to about
0.2 M.
[0046] Preferably, the coagulating effect of any residual
coagulation protein in the fibrin should be neutralized before
employing it in a medical device, such as a stent in order to
prevent clotting at the fibrin interface with blood after
implantation. This can be accomplished, for example, by treating
the fibrin with irreversible coagulation inhibitor compounds or
heat after polymerization. For example, hirudin or
D-phenylalanyl-propylarginine chloromethyl ketone (PPACK) can be
used for this purpose. Anticoagulants, such as heparin, can also be
added to reduce the possibility of further coagulation.
[0047] The porous sheet can then be placed into a fibrinogen
solution in order to fill the pores with fibrinogen, followed by
application of a solution of thrombin and fibrinogen to the surface
of the sheet material to establish a fibrin matrix that occupies
both the surface of the sheet and the pores of the sheet.
Alternatively, the thrombin can be included within the first
fibrinogen solution. If desired, ultrasonics, vacuum, and/or
pressure can be used to ensure that the fibrinogen applied to the
sheet is received into the pores.
[0048] Methods of Making an Implantable Device Having a Therapeutic
Substance
[0049] Referring now to FIG. 3, a structure having a porous
material is loaded with a heavy metal water-soluble salt.
Preferably, this step includes contacting, more preferably
immersing, the structure with an aqueous solution of the heavy
metal water-soluble salt, as described above. Preferably, the heavy
metal water-soluble salt is dispersed throughout a substantial
portion of the porous material. This may be assisted by degassing
the pores of the structure by such techniques as ultrasound or
vacuum degassing. The resulting stature can now be sterilized,
packaged and, optionally, stored until use.
[0050] In one embodiment of the invention, a sterilized structure
is shipped or delivered to the relevant consumer. The structure is
substantially contemporaneously loaded with a water soluble
therapeutic material. Preferably, the loading of the therapeutic
material includes contacting, more preferably immersing, the porous
material in an aqueous solution comprising a salt of the
therapeutic material, as described above. Again, degassing of the
device can help to bring the therapeutic material into the pores. A
water-insoluble therapeutic salt is thereby formed within the
porous material. Examples of aqueous radioactive salt solutions for
radiotherapy include Nal.sup.125, K.sub.2S.sup.35O.sub.4,
Na.sub.2S.sup.35O.sub.4, and Na.sub.3P.sup.32O.sub.4, to name a
few.
[0051] This method is advantageous in that the structure can be
loaded with the therapeutic material in situ, i.e., at or near the
point of therapeutic use, typically before administration,
preferably implantation, to a patient. This is particularly useful
because the device can be stored and transported prior to
incorporation of the therapeutic material. This feature has several
advantages. For example, the relevant consumer can select the
therapeutic material to be used from a wider range of therapeutic
materials, e.g., a radioisotope with a certain half-life with
certain particle emitting characteristics can be selected. Thus,
the therapeutic material selected is not limited to only those
supplied with the device but can instead be applied according to
the therapy required.
[0052] In another aspect of the invention, a sterilized structure
is loaded with a therapeutic material. Preferably, the loading of
the therapeutic material includes contacting, more preferably
immersing, the porous material in an aqueous solution comprising a
salt of the therapeutic material, wherein a water-insoluble salt of
the therapeutic material is formed within the porous material.
Examples of therapeutic salt solutions may be those previously
mentioned above. The structure is preferably packaged and can be
shipped to the relevant consumer. The structure can now be
administered to, preferably implanted into, a patient. Thus, in
this embodiment, the structure is loaded with the therapeutic
material prior to reaching the point of use, which may be more
convenient depending upon the facilities available to the relevant
consumer.
EXAMPLES
[0053] The following nonlimiting examples will further illustrate
the invention. All parts, percentages, ratios, etc. are by weight
unless otherwise indicated.
Example 1
[0054] The following solutions were used in the procedure:
[0055] Solution A: 1-10% aqueous solution of BaCl.sub.2
[0056] Solution B: 1-10% aqueous solution of Ba(NO.sub.3).sub.2
[0057] Solution C: 1-10% aqueous solution of
Na.sub.2S.sup.35O.sub.4
[0058] A porcine fibrin stent made according to U.S. Pat. No.
5,510,077 was treated by rehydration in Solution A by immersion for
about 5 to about 10 minutes. The stent was removed and excess
solution was blotted with absorbent paper. The stent was then
dehydrated and sterilized by gamma radiation. Alternatively,
Solution B can be used in place of Solution A.
[0059] This treated stent was then rehydrated in an aqueous
solution of Na.sub.2S.sup.35O.sub.4 radioisotope having a specific
activity of about 10 .mu.Ci/ml to about 500 .mu.Ci/ml. A white
precipitate of BaS.sup.35O.sub.4 was observed within the pores of
the stent surface. The stent can now be implanted into an artery
for localized delivery of .beta.-radiation or packaged for delivery
to the consumer.
Example 2
[0060] Fibrin stents made according to U.S. Pat. No. 5,510,077 were
soaked in a 20% by weight solution of BaCl.sub.2 (preferably
soaking for about 10 to 30 minutes). The stents were then subjected
to degassing by vacuum to remove air from the pores of the fibrin
matrix, thus allowing the BaCl.sub.2 solution to fill the pores.
The stents were dried overnight. The dried stents were placed into
a solution of sodium heparinate (preferably soaking in a solution
of 1000 U/ml to 20,000 U/ml for 10-20 minutes--most preferably a
solution of at least 10,000 U/ml) to allow the BaCl.sub.2 in the
fibrin matrix to react with the sodium heparinate to form barium
heparinate which was precipitated within the fibrin matrix.
Scanning electron microscopy (SEM) showed that particulates of
barium heparinate on the order of 10 microns (i.e., micrometers)
and smaller were trapped within the fibrin matrix (FIG. 4). In vivo
evaluation of the barium heparinate stents were carried out using a
carotid crush model in pigs with standard fibrin stents as
controls. After 24 hours, the stents were compared for flow and
were then examined histologically. While flow did not differ in a
statistically significant manner between the control stent and the
barium heparinate stent, the histological study showed
substantially reduced clot formation on the lumenal surface of the
barium heparinate stent (FIG. 5a) when compared with the control
stent (FIG. 5b).
Example 3
[0061] (A) Preparation of porous polyurethane coated stents. Wiktor
type stents were placed over 3.0 mm diameter smooth glass rods and
rolled by hand to assure a snug fit. The stent and rod assemblies
were dipped in 1-methy-2-pyrrolidinone (NMP) alone at room
temperature, allowed to drain vertically for a few seconds, then
rotated horizontally while dusting with 400-500 mesh sodium
bicarbonate until no further bicarbonate would adhere. After gently
tapping the rod assemblies to dislodge lightly adherent
bicarbonate, the assemblies were dipped once in a solution of 10
wt.% polyurethane in NMP. After draining vertically for a few
seconds, the rod and stent assemblies were rotated horizontally
while dusting with 400-500 mesh sodium bicarbonate until no further
sodium bicarbonate adhered, then gently tapped to dislodge lightly
adherent sodium bicarbonate. The assemblies were immersed in water
for about 5 minutes, then removed and the coating lightly compacted
by gently rolling the coated stent on the mandrel against a wet
paper towel. After immersing the stent assemblies in fresh water
for at least 8 hours at room temperature the coated stents were
removed from their mandrels and immersed in fresh water for 4-8
hours at room temperature. The coated stents were subsequently
dried in a forced air oven at 50.degree. C. for about 8 hours and
then trimmed of excess coating beyond the stent wires. After
passing a visual inspection the porous polyurethane stents were
ready for subsequent fibrin impregnation.
[0062] (B) Preparation of the composite porous polyurethane-fibrin
stent. The porous polyurethane stents prepared in (A) were
suspended in a fibrinogen-thrombin solution. The ratio of
fibrinogen to thrombin was predetermined so that the clotting time
of fibrinogen was between 5-10 minutes. The fibrinogen-thrombin
solution containing the porous polyurethane stents was subjected to
vacuum or ultrasonic degassing for about 34 minutes. Using either
method, the air in the pores of the porous polyurethane stents was
driven out and the pores filled with the fibrinogen-thrombin
mixture. The stents were then removed from the fibrinogen-thrombin
solution and clotting of the fibrinogen in the pores was
further-carried out for another 10-20 minutes at room temperature
after which the stents were incubated in sterile water overnight.
After incubation, the stents were dehydrated at room temperature
for at least 2 hours. Alternatively, the porous polyurethane fibrin
composite stents can be further compressed in glass molds to
densify the fibrin prior to dehydration.
Example 4
[0063] In Vivo Assessment. A total of eight prototype stents from
example 3 were implanted into 4 pigs for an in vivo assessment. In
particular, a single stent was implanted into 2 coronary arteries
in each of the 4 pigs. The purpose of this pilot animal study was
to asess the deliverability of the composite (fibrin filled porous
polyurethane) stents, the acute clinical performance, and to
determine the 28-day biocompatibility in terms of tissue response.
The pigs were given 325 mg (ASA) per day throughout their course
and heparin during the procedure. Stents were crimped onto
commercially available PTCA catheters and implanted using standard
stent delivery techniques. All eight stents were easily crimped
tightly onto the balloons and successfully delivered to the target
site through 8 Fr or 9 Fr guide catheters. Implants were
successfully performed in all 3 major coronary arteries (RCA, LAD,
LCX). Full expansion was acheived with 6-8 atmospheres pressure.
Post-implant angiograms indicated good flow, with wide open lumens
in all vessels. There were no acute events and the animals
recovered without incident. Three of the animals reached their
28-day planned sacrifice date while the fourth pig was humanely
sacrificed early at 21 days due to a chronic lung infection
(non-cardiac related). All eight stented artery segments were
pressure perfusion fixed and sectioned using standard
histopathology processing techniques, and subsequently examined
histologically. Upon exam all eight stented artery segments were
patent (i.e., open)--there were no total occlusions. In 3 of the
pigs (6 stents) the lumens were widely patent with mostly thin to
very thin amounts of maturing neointima deposited on top of the
porous film. In addition, there was absence of or minor amounts of
inflammation near the stent material. In the 4th pig (2 stents) a
thick neointima compromised the lumen with >60% stenosis and
there was significant inflammatory response in the tissue
surrounding the stent material. In this particular animal it was
noted that the arteries were small and there may have been an
overinjury component contributing to the gross vessel response. The
pores of the thin film in all these stents were typically filled
with native tissue as the exogenous fibrin appears to have been
replaced with a cellular infiltrate typical of maturing neointima.
There were also some areas of the porous film structure which were
still containing the yet-to-be-absorbed exogenous fibrin.
[0064] FIGS. 6 and 7 are 28-day examples of elastic van Gieson
(EVG) and haematoxylin and eosin (H&E) stains respectively.
FIG. 6 indicates the artery wall structure and expanded stent (wire
holes plus porous film) with the thin film of neointima on top of
the porous film. FIG. 7 is a higher magnification photo of the
cellular infiltrate into the porous film with neovascularization
taking place at the porous film/neointima interface. Original slide
magnifications are 7x and 67x, respectively, for FIGS. 6 and 7.
This data indicates that porous polyurethane stents having fibrin
incorporated therein can be advantageously used for preventing
restenosis without any therapeutic material if so desired.
Example 5
[0065] (1) Porous polyurethane impregnated with barium heparinate.
A porous polymer stent prepared according to Example 3(A) was
soaked in a 20 wt. % solution of barium chloride (BaCl.sub.2) for
about 10-20 minutes. The stent was then removed and dehydrated for
at least 2 hours (preferably overnight). After dehydration, the
barium chloride impregnated porous polyurethane stent was placed in
a sodium heparin solution with a concentration of at least 10,000
U/ml. The stent was immersed in the heparin solution for about
10-15 minutes during which time the barium chloride reacted with
sodium heparin to form barium heparinate. The reaction took place
within the pores containing barium chloride and also at the surface
of the stent where some of the barium chloride leached out The
barium heparinate, which is not soluble in water, was trapped
within the pores of the porous polyurethane stent. The stent was
rinsed with sterile water to remove excess, unreacted barium
chloride and then dehydrated.
[0066] (2) Composite porous polyurethane-fibrin stent impregnated
with barium heparinate. The procedure used was similar to that in
(1) above except that the stent described in Example 3(B) was used.
After the reaction between barium chloride and sodium heparin was
completed, the barium chloride was replaced by barium heparinate
that was trapped within the fibrin matrix.
Example 6
[0067] Preparation of porous polyurethane coated stent by using a
freeze-immersion-precipitation method. Bare stents were cleaned
with a mixture of alcohol/water (50:50), and then Freon TE/TF
(DuPont) and then dried. After cleaning, the stents were expanded
to about 3.1 mm in diameter under clean room conditions. The stents
were then individually placed in glass mold cavities. The glass
mold cavities have similar configurations as those described in
U.S. Pat. No. 5,697,967. A polyurethane solution (6.5 wt. %
polyurethane in 1,4-dioxan) was injected into the mold cavities
using a 3 ml sterile syringe and a 18 Ga sterile needle. After
injecting the polymer solution, the glass mold (containing the
stents and polymer solution) was place in a refrigerator at
3.degree. C. for 2 hours. Since the freezing temperature of dioxan
is at 12.degree. C., the polymer solution frozen slowly thus
creating a coarse structure of solvent/polymer. After 2 hours at
3.degree. C., the mold was removed from the refrigerator and
transferred to a freezer at -15.degree. C. to -18.degree. C. and
kept at this temperature for an additional 1-1.5 hours. The mold
was then immersed in an ice cold water bath at 3.degree. C. for 34
days to allow the solvent to leach out into the ice cold water.
Porous, uniform polyurethane stents were formed after the dioxan
completely dissolved into the ice cold water. The mold (and the ice
cold water bath) were allowed to warm up slowly to room
temperature. The stents were then removed from the glass mold
cavities and continued to immerse in fresh water at room
temperature for at least one day to remove traces of solvent. The
stents were then air dried at room temperature under a clean room
flow hood.
[0068] The above two-step freezing method was designed to create
porous polyurethane stents with desired pore structures and pore
sizes. By freezing the polymer solution first at 3.degree. C.,
large pores (in the order of 50 micrometers or greater) were
formed. This was reinforced with smaller pores formed due to
additional freezing and crystallization of solvent and polymer at
-15.degree. C. to -18.degree. C. The two-step freezing method can
create porous polyurethane stents with pore sizes range from 5 to
120 micrometers (with average pore sizes of 50 to 60 micrometers).
Pore sizes at 50 to 60 micrometers are desirable for tissue
ingrowth.
[0069] Fibrin and barium heparinate can also be incorporated in
this type of stent using the procedures described in Examples 3 and
5.
[0070] The complete disclosures of all patents, patent
applications, and publications referenced herein are incorporated
herein by reference as if individually incorporated. Various
modifications and alterations of this invention will become
apparent to those skilled in the art without departing from the
scope and spirit of this invention, and it should be understood
that this invention is not to be unduly limited to illustrative
embodiments set forth herein.
* * * * *