U.S. patent application number 12/282715 was filed with the patent office on 2009-04-23 for methods of minimizing stent contraction following deployment.
This patent application is currently assigned to ARTERIAL REMODELING TECHNOLOGIES S.A.. Invention is credited to Patrick Sabaria.
Application Number | 20090105800 12/282715 |
Document ID | / |
Family ID | 38308646 |
Filed Date | 2009-04-23 |
United States Patent
Application |
20090105800 |
Kind Code |
A1 |
Sabaria; Patrick |
April 23, 2009 |
METHODS OF MINIMIZING STENT CONTRACTION FOLLOWING DEPLOYMENT
Abstract
The present invention provides methods for fabricating a stent
using a preheating stage. The inventors have found a fabrication
methods that result in the same and/or better product quality stent
using a single step process performed at a temperature of below
65.degree. C., more preferably below 60.degree. C., most preferably
below 55.degree. C. Stent fabrication under such reduced
temperature conditions reduces the exposure of the stent to adverse
temperature conditions, thereby enabling the greater retention of
the polymer's memory. Further, upon expansion, the stent does not
contract to a smaller diameter but instead remains at a constant
diameter or increases to a larger diameter.
Inventors: |
Sabaria; Patrick; (Saint Nom
La Breteche, FR) |
Correspondence
Address: |
TAROLLI, SUNDHEIM, COVELL & TUMMINO L.L.P.
1300 EAST NINTH STREET, SUITE 1700
CLEVEVLAND
OH
44114
US
|
Assignee: |
ARTERIAL REMODELING TECHNOLOGIES
S.A.
Noisy Le Roi
FR
|
Family ID: |
38308646 |
Appl. No.: |
12/282715 |
Filed: |
March 13, 2007 |
PCT Filed: |
March 13, 2007 |
PCT NO: |
PCT/IB2007/000589 |
371 Date: |
September 12, 2008 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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60781748 |
Mar 14, 2006 |
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60781747 |
Mar 14, 2006 |
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60781741 |
Mar 14, 2006 |
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60791220 |
Apr 12, 2006 |
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60814533 |
Jun 19, 2006 |
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60854075 |
Oct 25, 2006 |
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Current U.S.
Class: |
623/1.11 ;
264/239 |
Current CPC
Class: |
A61F 2210/0004 20130101;
A61F 2/91 20130101; A61F 2250/003 20130101; A61F 2/915 20130101;
A61B 90/39 20160201; A61F 2002/91541 20130101; A61F 2002/91558
20130101; A61F 2/82 20130101; A61F 2250/0098 20130101; A61L 31/18
20130101; A61L 31/148 20130101 |
Class at
Publication: |
623/1.11 ;
264/239 |
International
Class: |
A61F 2/06 20060101
A61F002/06; B29C 43/02 20060101 B29C043/02 |
Claims
1. A method of educating a device comprising a cylindrical
structure, wherein said device comprises at least one polymer, said
method comprising: deforming said device to a diameter of 1.0 mm to
4.0 mm; and heating said device to a temperature above the Tg of
the polymer from which said device is formed.
2. The method of claim 1, whereby said device is formed from PLA75
and said device is heated to 80.degree. C. for 30 minutes.
3. A method of fabricating a device comprising a cylindrical
structure, wherein said device comprises at least one polymer, said
method comprising: educating said device to a final predetermined
diameter size; crimping at least part of said device to a diameter
that is smaller than the final predetermined diameter size, wherein
the crimping comprises heating said device to a temperature that is
below the glass transition state of said at least one polymer for a
time sufficient for said device to temporarily maintain a diameter
that is smaller than the final predetermined diameter size.
4. The method of claim 3, whereby said device comprising a
cylindrical structure is crimped onto an inflatable device.
5. The method of claim 3, whereby the temperature that is below the
glass transition state is high enough to allow reduction in
diameter of said device comprising a cylindrical structure but not
low enough to erase the memory of the final predetermined shape and
diameter of said device comprising a cylindrical structure.
6. The method of claim 3, whereby said temperature that is below
the glass transition state is about 50.degree. C.
7. A method of deploying a stent, said method comprising:
fabricating a device comprising a cylindrical structure wherein
said device comprises at least one polymer; educating said device
to a final predetermined diameter size and crimping said device to
a diameter that is smaller than the final predetermined diameter
size, wherein the crimping comprises heating said device to a
temperature that is below the glass transition state of said at
least one polymer for a time sufficient for said device to
temporarily maintain the diameter that is smaller than the final
predetermined diameter size; placing said crimped device within a
lumen of the patient's body; and expanding said device.
8. The method of claim 7, whereby said device expansion is
independent of an inflatable device.
9. The method of claim 7, wherein whereby said device comprising a
cylindrical structure is crimped onto an inflatable device.
10. The method of claim 7, whereby said inflatable device is
inflated and/or heated to initiate the expansion of said device
comprising a cylindrical structure.
11. The method of claim 7, whereby the positive recoil properties
of said device comprising a cylindrical structure expands said
device to its final predetermined diameter.
12. The method of claim 7, whereby said inflatable device is heated
to a temperature below the Tg of the at least one polymer to
initiate expansion of said device comprising a cylindrical
structure.
13. The method of claim 7, such that when said device comprising a
cylindrical structure is expanded, said device exhibits positive
recoil to create outward radial pressure.
14. The method of claim 1, whereby said device comprising a
cylindrical structure is a stent.
15. The method of claim 1, whereby said at least one polymer has at
least one in vivo lifetime.
Description
BACKGROUND OF THE INVENTION
[0001] The use of stents in various surgical, interventional
cardiology, and radiology procedures has quickly become accepted as
experience with stent devices accumulates and as the advantages of
stents become more widely recognized. Stents are often used in body
lumens to maintain open passageways such as the prostatic urethra,
the esophagus, the biliary tract, intestines, and various coronary
arteries and veins, as well as more remote cardiovascular vessels
such as the femoral artery.
[0002] Stents are often used to treat atherosclerosis, a disease in
which vascular lesions or plaques consisting of cholesterol
crystals, necrotic cells, lipid pools, excess fiber elements and
calcium deposits accumulate in the walls of an individual's
arteries. One of the most successful procedures for treating
atherosclerosis is to insert a deflated balloon within the lumen,
adjacent the site of the plaque or atherosclerotic lesion. The
balloon is then inflated to put pressure on and "crack" the plaque.
This procedure increases the cross-sectional area of the lumen of
the artery. Unfortunately, the pressure exerted also traumatizes
the artery, and in 30-40% of the cases, the vessel either gradually
renarrows or recloses at the locus of the original stenotic lesion.
This renarrowing is known as restenosis
[0003] A common approach to prevent restenosis is to deploy a
metallic stent to the site of the stenotic lesion. Although
metallic stents have the mechanical strength necessary to prevent
the retractile form of restenosis, their presence in the artery can
lead to biological problems including vasospasm, compliance
mismatch, and even occlusion. Moreover, there are inherent,
significant risks from having a metal stent permanently implanted
in the artery, including erosion of the vessel wall. The stents may
also migrate on occasion from their initial insertion location
raising the potential for stent induced blockage. Metal stents,
especially if migration occurs, cause irritation to the surrounding
tissues in a lumen. Also, since metals are typically much harder
and stiffer than the surrounding tissues in a lumen, this may
result in an anatomical or physiological compliance mismatch,
thereby damaging tissue or eliciting unwanted biologic responses.
In addition, the constant exposure of the stent to the blood can
lead to thrombus formation within the blood vessel. Stents also
allow the cellular proliferation associated with the injured
arterial wall to migrate through the stent mesh, where the cells
continue to proliferate and eventually lead to the narrowing of the
vessel. Further, metal stents typically have some degree of
negative recoil. Finally, metallic stents actually prevent or
inhibit the natural vascular remodeling that can occur in the
organism by rigidly tethering the vessel to a fixed, maximum
diameter.
[0004] Because of the problems of using a metallic stent, others
have recently explored use of bioabsorbable and biodegradable
materials stents. The conventional bioabsorbable or bioresorbable
materials from which such stents are made are selected to absorb or
degrade over time. This degradation enables subsequent
interventional procedures such as restenting or arterial surgery to
be performed. It is also known that some bioabsorbable and
biodegradable materials tend to have excellent biocompatibility
characteristics, especially in comparison to most conventionally
used biocompatible metals. Another advantage of bioabsorbable and
biodegradable stents is that the mechanical properties can be
designed to substantially eliminate or reduce the stiffness and
hardness that is often associated with metal stents. This is
beneficial because the metal stent stiffness and hardness can
contribute to the propensity of a stent to damage a vessel or
lumen. Examples of novel biodegradable stents include those found
in U.S. Pat. No. 5,957,975, which is incorporated by reference in
its entirety.
[0005] There are, however, still problems with many biodegradable
stents. For example, testing in animals has shown that
biodegradable stents still suffer from multiple complications,
including relaxation-related negative recoil, lack of sufficient
radial strength, difficulty in deployment and distal migration of
the entire stent or portions thereof and formation of an occlusive
thrombus within the lumen of the stent.
[0006] Accordingly, it is desirable to have a new stent that
overcomes the disadvantages of the current stent designs. A
polymer-based stent that exhibits little to no relaxation-related
negative recoil when implanted in the blood vessel or duct of a
mammalian subject is desirable. Indeed, it is preferred that the
stent have a positive recoil. It is also desirable to have a
polymer-based stent assembly that does not require a mechanical
restraint to prevent the stent from expanding when stored at room
temperature. To achieve these goals, the stent is fabricated using
several heating steps. For instance, in a typical fabrication there
is at least one preheating stage performed prior to the cutting
procedure. Crimping step is performed as a two-step process at a
temperature of 65.degree. C. This method, however, has the drawback
in that the multiple heating steps alter the stent "memory" of the
ideal final diameter.
[0007] The inventors have found a novel method of stent fabrication
that decreases the time the stent is exposed to adverse temperature
condition, thereby enabling greater memory retention of the
polymers diameter.
SUMMARY OF THE INVENTION
[0008] The present invention provides methods for fabricating a
stent using a preheating stage. As opposed to using a two or more
step heating process, the inventors have found a preferable
embodiment using new fabrication methods that result in the same
and/or better product quality stent using a single step process
performed at a temperature of below 65.degree. C., more preferably
below 60.degree. C., most preferably below 55.degree. C. In certain
embodiments, a temperature below about 50.degree. C. is most
preferred. Stent fabrication under such reduced temperature
conditions permits technicians to avoid burning their hands but
more importantly, results in a reducing the exposure of the stent
to adverse temperature conditions, thereby enabling the greater
retention of the polymer's memory.
[0009] Maintaining the stent at a temperature of fabrication as
described also provides a beneficial result by being below the
glass transition temperature of the polymeric material. Under the
previously employed procedure(s), the stent would be expanded to
balloon nominal diameter of: coronary balloon from typically 2.0 mm
to 4.5 mm; vascular peripheral balloon (PTA) from typically from 3
mm to more than 20 mm depending on balloon diameter. For example,
for a 3 mm balloon, the stent would be expanded to 3 mm and when
the balloon was removed, it would contract to a diameter of 2.7,
followed by a slow expansion to the final desired diameter of 3.2
mm. Under the presently employed procedure at T=zero, the stent is
expanded to 3 mm when deployed. Following initial deployment the
stent does not contract, but instead remains at a diameter of 3 mm
and gradually over time expands to the desired diameter of 3.2 mm.
Furthermore, this deployment to a final desired diameter is
essentially balloon inflation independent.
DETAILED DESCRIPTION
Definitions
[0010] "Bioresorbable polymer" as used herein refers to a polymer
whose degradation by-products can be bio-assimilated or excreted
via natural pathways in a human body. "Crimping" as used herein
refers to a process that involves radial pressing on a polymeric
cylindrical device having slits, or openings in the wall thereof to
allow a decrease in the diameter of the device without
substantially affecting the thickness of the wall or struts of the
cylindrical device. Such process, typically also results in an
increase in length of the cylindrical device. "Degradable polymer"
or "biodegradable polymer" as used herein refers to a polymer that
breaks down into monomers and oligomers when placed in a human body
or in an aqueous solution and maintained under conditions of
temperature, osmolality, pH, etc., that mimic physiological media
preferably without involving enzymatic degradation to minimize the
risk of triggering the antigenantibody defense system of the human
body. "Final predetermined shape and diameter" as used herein
refers to the desired diameter, length, design and wall thickness
of a stent that has been deployed to a target site in a vessel,
particularly a blood vessel, duct, or tube in a mammalian subject,
particularly a human subject. "Negative recoil" as used herein
refers to an undesirable decrease in the size or diameter of an
expanded stent after initial deployment. "Positive recoil" as used
herein refers to an increase in the size or diameter of a stent
that has been educated to have a desired final diameter but has not
been fully expanded to the desired final diameter.
"Relaxation-related recoil" as used herein refers to the slow
change in dimensions of a polymeric device due to a time-dependent
slow rearrangement of molecule conformations according to a
well-known behavior of viscoelastic polymeric matters. Such
rearrangement is due to thermal agitation that slowly leads the
polymeric material to a thermodynamic equilibrium typical of the
storage conditions when it has been processed under different
environmental conditions. Relaxation is very slow below Tg, i.e.
when the matter is in the glassy state. "Tg" or "glass transition
temperature" as used herein refers to the temperature at which a
polymer changes from a rubbery state to a glassy state and vice
versa.
[0011] The present invention provides a stent fabrication method
that only requires a single step process performed at the
temperature of below 65.degree. C., more preferably below
60.degree. C., most preferably below 55.degree. C. In certain
embodiments, a temperature below about 50.degree. C. is most
preferred. The process that results in the same and/or better
quality product that methods that require more than one heating
steps. Maintaining the stent at a temperature of the invention or
below also provides a beneficial result by avoiding the glass
transition temperature of the polymeric material. This new stent
fabrication method also results in the final stent having no
negative recoil when the stent is deployed into a mammalian body.
Indeed, the final stent has positive recoil when deployed.
[0012] Under the previously employed procedure(s), the stent would
be expanded to 3 mm and when the balloon was removed, it would
contract to a diameter of 2.7 mm, followed by a slow expansion to
the final desired diameter of 3 mm. Under the presently employed
procedure at T=zero, the stent is expanded to 3 mm when deployed.
Following initial deployment the stent does not contract, but
instead remains at a diameter of 3 mm. Further, over several days,
the stent may further expand to a larger desired final diameter,
say, of approximately 3.2 mm. In one embodiment, a balloon is used
merely as a carrier for the stent through the body. Further, the
deployment of the stent into the body may use the balloon as a
carrier that also initiates the stent expansion. It is also
contemplated that the deployment of the stent into the body may be
balloon inflation independent. In addition, it is contemplated that
the stent final diameter be larger than that of the deploying
balloon diameter.
I. Stent Fabrication and Properties
[0013] The stents may be formed from any biodegradable,
biocompatible, bioresorbable polymer, preferably a thermoplastic
polymer. As used herein, a bioresorbable polymer is one whose
degradative products are metabolized in vivo or excreted from the
body via natural pathways. Preferably, the stent of the present
assembly is formed from a degradable and bioresorbable polymer
having a Tg at least 8 degrees above 37.degree. C., preferably at
least 20 degrees above 37.degree. C. The polymer of the stent can
be a homopolymer or a copolymer. Preferably, the stent is formed
from a thin layer of one or more amorphous, bioresorbable polymers,
i.e., the polymers used to form the stent preferably are not
crystalline. It is also preferred that the polymers used to form
the stent do not generate crystalline residues upon degradation in
vivo. It is also contemplated that the chains of the polymer may be
or may not be cross-linked. Light cross-linking, however, is
acceptable provided that thermal and viscoelastic characteristics
that allow education, crimping, and deployment of the device are
sufficiently maintained.
[0014] Appropriate biodegradable polymers may include, but are not
limited to, poly(L-lactide), polyglycolide, poly(D,L-lactide),
copolymers of lactide and glycolide, polycaprolactone,
polyhydroxyvalerate, polyhydroxybutyrate,
polytrimethylenecarbonate, polyorthoesters, polyanhydrides, and
polyphosphazenes. Examples of the types of polymers that are
suitable for the stent of the present invention include, but are
not limited to, lactic acid-based stereocopolymers (PLAx copolymers
composed of L and D units, where X is the percentage of L-lactyl
units) (55<Tg<60), copolymers of lactic and glycolic acids
(PLAxGAy, where X, the percentage of L-lactyl units, and Y, the
percentage of glycolyl units, are such that the Tg of the copolymer
is above 45.degree. C.), and Poly(lactic-co-glycolic-co-gluconic
acid) where the OH groups of the gluconyl units can be more or less
substituted (pLAxGayGLx, where X, the percentage of L-lactyl units,
and Y, the percentage of glycolyl units, and Z the percentage of
gluconyl units are such that the Tg of the terpolymer is above
45.degree. C.). Other suitable polymers include, but are not
limited to, polylactic acid (PLA), polyglycolic acid (PGA)
polyglactin (PLAGA copolymer), polyglyconate (copolymer of
trimethylene carbonate and glycolide, and a copolymer of
polyglycolide or lactide acid or polylactic acid with
.epsilon.-caprolactone), provided that the polymer has a glass
transition temperature, Tg, of at least 45.degree. C. or
greater.
[0015] In one preferred embodiment, the stent comprises a
polylactic acid stereocopolymer produced from L and DL lactides.
The polymer is designated herein as "PLAX" where X represents the
percentage of the L-lactic acid units in the mixture of monomers
used to prepare the lactides. Preferably X is in the range of 10 to
90, more preferably 25 to 75. In another preferred embodiment, the
stent comprises a poly-lactic acid, glycolic acid copolymer
produced from L and DL lactides and glycolides. The polymer is
designated herein as "PLAXGAY" where Y represents the percentage of
glycolic acid units in the mixture of monomers used to prepare the
copolymers. Preferably, the copolymers do not contain glycolyl
repeating units since such units are known to be more inflammatory
than lactyl repeating units. Preferably, the polymers are prepared
using Zn metal or Zn lactate as initiator. To ensure good initial
mechanical properties of the stent, the molecular weight of the
polymer in the region having the second in vivo lifetime is
preferably above 20,000 daltons, more preferably 100,000 daltons or
larger. The polydispersity, I=Mw/Mn, is preferably below two and
should not greatly reflect the presence of low molecular weight
oligomers smaller than 2,000 daltons as determined by size
exclusion chromatography. Optionally, the polymeric layer used to
make the stent may be impregnated with an anticoagulant agent, such
as heparin, anti-oxidants, such as vitamin E, compounds that
regulate cellular proliferation, or anti-inflammatory drugs, such
as corticosteroids, to provide localized drug delivery. Such drugs
are incorporated into the polymeric layer using techniques known in
the art. Agents may also be incorporated into the base polymer that
forms the body of the stent, as long as the incorporation does not
have significant adverse effects on the desired physical
characteristics of the stent such as during radial stent deployment
and degradation time. For intravascular stents, it is preferred
that the film have a thickness of from about 0.05 mm to 0.2 mm.
[0016] It is contemplated that the stent may be made by any method.
In one preferred embodiment, the stent is a formed from a
biodegradable polymeric band comprising a head having a slot and a
tongue comprising a catch or locking mechanism proximate the
longitudinal edge thereof. The cylindrical element, which has an
inner and outer surface, is formed by inserting a portion of the
tongue through the slot to provide a cylindrical element having a
first reduced diameter configuration. Following deployment, the
cylindrical element is in a second expanded diameter configuration
wherein the distal catch mechanism engages the inner surface of the
head and prevents radial collapse or recoil of the polymeric stent.
In a second preferred embodiment, the stent is formed from a
plurality of interconnected polymeric bands each of which comprises
a head having a slot and a tongue comprising a catch mechanism
proximate the longitudinal edge thereof.
[0017] In one embodiment, the stent is formed by laser cutting of a
cylindrical tube. In another embodiment, the stent is formed by
laser cutting a flat polymeric sheet in the form of the stent, and
then rolling the pattern into the shape of the cylindrical stent
and providing a longitudinal weld to form the stent. In another
embodiment, the stents are created by chemically etching a flat
polymeric sheet and then rolling and welding it to form the stent,
or coiling a polymeric wire to form the stent.
[0018] In another preferred embodiment, the stent may also be
formed by molding or injection molding of a thermoplastic or
reaction injection molding of a thermoset polymeric material. The
flat grid is then rolled and extremities are welded or glued to
form a cylinder. Filaments of the compounded polymer may be
extruded or melt spun. These filaments can then be cut, formed into
ring elements, welded closed, corrugated to form crowns, and then
the crowns welded together by heat or solvent to form the stent.
Lastly, hoops or rings may be cut from tubing stock, the tube
elements stamped to form crowns, and the crowns connected by
welding or laser fusion to form the stent.
[0019] Generally, the struts are arranged in patterns that are
designed to contact the lumen walls of a vessel and to maintain
patency of the vessel thereby. A myriad of strut patterns are known
in the art for achieving particular design goals.
[0020] It is contemplated that a crimped stent may incorporate
slits or open spaces to allow for the temporary reduction in
diameter of the cylindrical tube without substantially altering the
wall thickness. Moreover, a stent embodying the present invention
can include teeth and corresponding catching structure that
operates to maintain an expanded form. Moreover, polymer based
stents embodying structure defined by a wire or ribbon coil or
helix or a knitted mesh configuration are possible examples of
self-expanding stents. Other important design characteristics of
stents include radial or hoop strength, expansion ratio or coverage
area, and longitudinal flexibility. One strut pattern may be
selected over another in an effort to optimize those parameters
that are of importance for a particular application.
[0021] It is also contemplated that the biodegradable stent may
have a programmed pattern of in vivo degradation. Stent polymeric
structure allows for differential speed degradation. See, for
example, U.S. Pat. No. 5,957,975, the entirety of which is
incorporated by reference. In one embodiment, the stent comprises
at least one substantially cylindrical element having two open ends
and a plurality of regions circumferentially spaced around the
cylindrical element and extending from one open end to the other
open end of the cylindrical element. Each of the regions is
configured or designed to have a desired in vivo lifetime. At least
one of the regions is designed to have a shorter in vivo lifetime
than the other region or regions. This means that the region having
the shorter in vivo lifetime degrades sooner after deployment than
the regions having a longer in vivo lifetime. Thus, when stents
designed in accordance with the present invention are deployed
within the lumen of a vessel of a patient, the cylindrical element
acquires one or more fissures which extend from one open end of the
cylindrical element to the other open end of the cylindrical
element within a desired, predetermined period of time after the
stent is deployed in the patient. It has been determined that such
fragmentation within a predetermined period of time after
deployment allows for enlargement of the lumen of the vessel via
the process of arterial remodeling.
[0022] The regions of the stent with the different in vivo
lifetimes can be made in a variety of ways. Preferably, such stents
are made by producing regions having a first in vivo lifetime,
i.e., a shorter in vivo lifetime, in a polymeric layer having the
predetermined second, or longer, in vivo lifetime. The regions
having the first in vivo lifetime are produced by heating the
respective regions of the polymeric layer having a second in vivo
lifetime for a time and at a temperature sufficient to cause local
partial degradation of the polymeric chains. Such treatment, which
can be accomplished using a piloted hot needle, laser beam, or flow
of hot air, renders the polymer in the heated region more sensitive
to hydrolytic degradation. Alternatively, the regions having a
first in vivo lifetime may be produced in a polymeric layer having
a second in vivo lifetime by incorporating a sufficient number of
acidic ions into the respective regions of the polymeric layer.
Preferably, the acidic ions are provided by compounds that are not
soluble in blood.
[0023] Regions having a first in vivo lifetime can also be produced
in a polymeric film having a second in vivo lifetime by exposure of
the respective regions to beta radiation or gamma radiation for a
sufficient time to induce partial cleavage of the polymeric chains
within the respective regions. Provided the polymeric layer has a
thickness of less than 0.3 mm, regions having a first in vivo
lifetime can also be produced in a polymeric film having a second
in vivo lifetime by introducing areas of mechanical weakness into
the polymer. One method of introducing mechanical weakness is by
reducing the thickness of the polymer in the respective region or
forming holes therein. Regions having a first in vivo lifetime can
also be produced in a polymeric film having a second in vivo
lifetime by applying mechanical stress to the respective region.
However, this latter process is difficult to control and, thus, is
less preferred. Differing lifetimes can also be created by
providing one or more different coatings over different regions of
the biodegradable stent.
[0024] Another method for producing a polymeric layer in which one
region or a plurality of spaced apart regions have a first in vivo
lifetime and other regions have a second in vivo lifetime is to
incorporate strips or fibers of a faster degrading bioresorbable
polymer into a film made from a slower degrading polymer. For
example, a mesh or a parallel array of fibers or strips of PGA or
any other faster degrading bioresorbable polymer can be embedded
into the respective regions of a polymeric film of PLA that may be
designed to be slower degrading. Embedding can be achieved by
inserting the mesh or fibers between two melted sheets of the
slower degrading polymer. Provided the relative solubilities are
compatible, the fibers or mesh can be placed in an organic solution
of the slower degrading polymer and the desired polymeric film
formed by evaporation of the organic solvent. One example of a
method for embedding a mesh made from one polymer into a polymeric
layer made from a second polymer is described in U.S. Pat. No.
4,279,249 issued to Vert et al. on Jul. 21, 1981, which is
specifically incorporated herein by reference. A stent having the
desired shape and orientation of regions is then formed from the
polymeric layer by standard techniques such as stamping, employing
a laser beam, or any other technique used in the art to tool a
polymeric film.
[0025] The initial polymeric cylindrical device that is formed by
any of these processes can be configured to have the final
predetermined shape, length, wall thickness and diameter, all of
which are tailored to the application for which the stent is to be
utilized. For example, for cardiovascular applications the initial
polymeric device that is formed by these processes can have a final
predetermined length ranging from 0.5 cm to approximately 3 cm. For
certain applications, the initial polymeric cylindrical device can
have a final, predetermined diameter ranging from 0.50 mm to 8.0 mm
with a final, predetermined wall thickness ranging from 0.05 to 0.5
mm. Alternatively, the initial cylindrical device that is formed by
any of these processes can have a smaller diameter than the final
predetermined diameter.
[0026] In those instances where the initial polymeric cylindrical
device has a smaller diameter than the final predetermined
diameter, slits or openings are formed in the cylindrical device as
described above, and then the cylindrical device is deformed or
expanded to the final shape and diameter. This can be achieved by
inserting an expandable device such as a balloon into the
stent.
II. Educating and Crimping the Stent
[0027] While it is at the final predetermined shape, size, and
diameter, the cylindrical device is educated by heating the device
to a temperature above the Tg of the polymer from which the device
is formed. The device is heated for a time sufficient to erase any
former process-related memory and to impart a new memory of the
final predetermined shape and diameter to the polymeric cylindrical
device. It is believed that such conditions allow the polymer
chains to relax and reorganize themselves from an entanglement
typical of the former processing stages to an entanglement typical
of the high temperature at which the cylindrical device is
compatible with the final or deformed shape and size. When the
polymeric cylindrical device has an initial diameter that is less
than the final predetermined diameter, it is desired to heat to a
temperature well above the Tg of the polymer. This heating step
erases the anisotropic stresses promoted by the extrusion or
molding process and the former processing-related memory of the
polymer chains. Good results have been obtained by heating a
laser-precut polymeric cylindrical device formed from PLA75 and
deformed from a diameter of 1.0 mm to 4 mm at a temperature of
80.degree. C. for 30 minutes. Temperatures of from about 45.degree.
C. to about 120.degree. C. and times of 5 minutes or more should be
suitable for educating stents made from PLAx with 0<X<100,
PLAxGAy with 0<x<25 and 75<Y<100, or any
PLAxGAyGLz.
[0028] The polymeric cylindrical device is then crimped. "Crimping"
as used herein refers to a process that involves radial pressing on
a polymeric cylindrical device having slits, or openings in the
wall thereof to allow a decrease in the diameter of the device
without substantially affecting the thickness of the wall or struts
of the cylindrical device. Such process may also result in an
increase in length of the cylindrical device.
[0029] To crimp the educated cylindrical device, it is mounted onto
a device with a smaller diameter. The diameter of the educated
cylinder is reduced by heating the cylinder to a temperature below
the Tg of the polymer while evenly applying pressure on the
exterior surface of the wall of the cylindrical device.
[0030] In some embodiments, the polymeric stent is crimped onto an
inflatable device such as an inflatable balloon catheter. In this
instance, the stent assembly after crimping comprises an inflatable
balloon catheter and an expandable, educated, polymeric stent
snugly and stably disposed thereon. Slits or open spaces that allow
for a reduction in diameter of the cylindrical device without
substantially altering the wall thickness during crimping are
incorporated into the cylindrical device prior to the time the
cylindrical device is crimped on the inflatable balloon catheter.
The temperature at which the cylindrical device is heated during
crimping is high enough to allow reduction in diameter of the
cylindrical device but low enough to not erase the memory of the
final predetermined shape and diameter of the educated cylindrical
device. Ideally, the temperature is less than the glass transition
state of the polymer. More preferably, the temperature is at about
50.degree. C. Thus, the temperature at which the educated
cylindrical device is heated during crimping is less than the
temperature at which the cylindrical device is heated during
education of the cylindrical device. Further, the time it takes to
crimp the educated cylindrical device can vary, depending upon the
temperature, size and composition of the stent
[0031] In accordance with the present method, expansion of the
polymeric stent can be achieved by any means. In one embodiment, a
balloon is used merely as a carrier for the stent through the body.
In this preferred embodiment, the stent expansion occurs by the
positive recoil properties of the stent; thus, the expansion is
balloon inflation independent. In another preferred embodiment, a
balloon is inflates and/or heated to initiates the stent expansion.
It is contemplated that the positive recoil properties of the stent
would expand the stent to its final predetermined diameter. The
temperature used to initiate the stent expansion may be any
temperature at or below the Tg of the polymer. In a less preferred
embodiment, a balloon is inflated to expand the polymeric stent to
its final predetermined shape.
[0032] In another aspect, the method of the present invention
starts with a polymeric tube whose diameter initially is less than
the final predetermined diameter. Such tube is first heated to a
temperature close to or above the Tg of the polymer and expanded to
provide a cylindrical device whose diameter is equal to the final
desired diameter. Thereafter the cylindrical device is educated as
described above to provide an educated cylindrical device having a
memory of the final predetermined shape and diameter, and then
crimped on a balloon catheter as described above to provide an
assembly comprising the balloon catheter and an expandable,
educated, polymeric stent snugly and stably disposed thereon.
[0033] The present invention also provides an assembly comprising
an inflatable balloon catheter and a polymeric stent prepared in
accordance with the present method.
[0034] Advantageously, the stent of the present invention exhibits
little to no relaxation-related negative recoil when deployed in
the blood vessel of a subject. Advantageously, the assembly of the
present invention has a diameter that allows it to be easily
inserted into a blood vessel of the subject and advanced to a
target site. Advantageously, the stent of the present invention
exhibits expansion (positive recoil) and adaptation to the geometry
of the artery when the stent is not fully deployed up to its final
diameter during deployment. Positive recoil over several days will
create outward radial pressure for long periods of time. This
outward radial pressure aids in positive vascular remodeling by
assisting stabilizing the injured artery or vulnerable plaque,
assist in cellular progress to repair injury of original acute
expansion, assist in security of tissue flaps, and the like.
[0035] In addition, the stent of the present invention is stably
disposed on the balloon, meaning that a mechanical restraint is not
required to prevent the stent from rapidly expanding to its final
diameter during storage at room temperature. Thus, although not
required, the assembly of the present invention, optionally, also
comprises a retractable sheath covering the exterior surface of the
stent. Such sheath serves to prevent deformation of the stent and
preclude or slow expansion during storage.
III. Procedures for Determining Times and Temperatures for
Educating and Crimping the Stent of the Present Invention
[0036] FIG. 1 shows one embodiment of the invention. FIG. 1 is a
method for determining the temperatures and times suitable for
educating the cylindrical device. This method results in a stent
that is resistant to negative recoil, and in fact has positive
recoil. Once the stent is created (step 101), the stent is crimped
onto a balloon catheter according to one of the above methods (step
102). The temperature of at which the stent is heated during
crimping is high enough to allow reduction in diameter of the stent
but low enough to not erase the memory of the final predetermined
shape and diameter of the educated stent. Ideally, the temperature
is less than the glass transition state of the polymer. More
preferably, the temperature is at about 50.degree. C. Thus, the
temperature at which the educated stent is heated during crimping
is less than the temperature at which the stent is heated during
education of the stent. Further, the time it takes to crimp the
educated stent can vary, depending upon the temperature, size and
composition of the stent. The method of FIG. 1 may be used to
assist in determining the ideal temperature to use to educate the
stent.
[0037] The balloon is then partially or fully inflated to initiate
stent expansion (step 103). The balloon is removed (step 104) and
the stent is stored at a temperature appropriate for storage (step
105). In one preferred embodiment, the storage temperature is
37.degree. C. While in storage, the stent may increase in diameter
because of the positive recoil properties of the stent. After a
predetermined amount of time, the stent will be examined to find if
it exhibits negative recoil (step 106). In one preferred
embodiment, the amount of time is 4 to 6 weeks. In another
preferred embodiment, the amount of time is the estimated for an
artery wall to recover from PTC angioplasty. If the stent exhibits
little to no negative recoil when stored under these conditions,
the times and temperatures employed for educating the stent are
suitable (step 108). In those cases where the polymeric stent
exhibits a small amount of recoil, the cylindrical device can be
educated at a diameter slightly larger than the final predetermined
diameter to compensate for the small amount of negative recoil
(step 109). Temperatures and times suitable for educating the stent
to a reduced diameter can be assessed by allowing the stent-mounted
balloon catheter of the present assembly to stay at room
temperature or at the storage temperature. If the crimped stent
stays collapsed at the small diameter corresponding to the deflated
balloon under these conditions, the times and temperatures employed
during suitable are suitable (step 107).
[0038] Optimization of the imparted stent mechanical properties
such as positive recoil can be achieved by storing the finished
product at a room temperature below 20.degree. C. Preferably, the
finish product is refrigerated at about 6.degree. to 8.degree.
C.
IV. Deployment of the Stent
[0039] The polymer-based stent is first preheating for a period of
3 to 6 min at around 37.degree. C. The preheating of the stent can
occur by any means, including heating in saline, the blood stream,
or hot air. After the preheating period, the polymer-based stent
assembly of the present invention is introduced into a duct, tube,
or vessel, e.g. a blood vessel of a mammalian subject, preferably
in conjunction with a guiding catheter, and advanced to a target
site, e.g. the site of stenotic lesion. After it is located at the
target site the balloon is rapidly inflated to initiate expansion
of the stent. Alternatively, the stent may be placed on a
deployment device that is capable of localized heating of the stent
when the stent is correctly positioned. During this process the
diameter of the stent increases, but the thickness of the walls of
the stent remain substantially the same.
[0040] It is further contemplated that fracturing of the plaque and
deployment of the stent may be done concurrently. If a balloon is
used in such cases, the balloon is inflated to a pressure of about
8 to 12 atmospheres to crack the plaque and expand the stent.
Alternatively, the vessel may be pre-dilated using angioplasty
without the stent. Thereafter, the stent is introduced into the
desired site on a separate catheter, preferably an expanding
balloon catheter.
[0041] In addition to coronary arteries, the present stent may be
used in other arteries such as for example, femeroiliac arteries,
the carotid artery, vertebro-basilar arteries, as well as in the
interior of other hollow passageways such as for example veins,
ureters, urethrae, bronchi, biliary and pancreatic duct systems,
the gut, eye ducts, and spermatic and fallopian tubes. Indeed, it
is further contemplated that certain aspects of the present
invention include devices that are used as substitutes for veins,
arteries, and ductal or tubal structures in the body.
[0042] It is to be recognized that aspects of the present invention
are applicable to other medical devices. For example, the disclosed
formulations can be employed to create a passive marker on an
interventional or surgical device, such as a biopsy needle or other
hand-held devices. In addition, entire medical devices or portions
thereof can embody the imagable material of the present
invention.
[0043] While only the presently preferred embodiments have been
described in detail, as will be apparent to those skilled in the
art, alternatives, additions, modifications and improvements may be
made to the device and method disclosed herein without departing
from the scope of the invention. Accordingly, it is not intended
that the invention be limited, except as by the appended
claims.
* * * * *