U.S. patent application number 11/693684 was filed with the patent office on 2008-05-01 for temporal intraluminal stent, methods of making and using.
This patent application is currently assigned to Biosensors International Group. Invention is credited to Debashis Dutta, John E. Shulze, Shih-Horng Su.
Application Number | 20080103584 11/693684 |
Document ID | / |
Family ID | 39262617 |
Filed Date | 2008-05-01 |
United States Patent
Application |
20080103584 |
Kind Code |
A1 |
Su; Shih-Horng ; et
al. |
May 1, 2008 |
Temporal Intraluminal Stent, Methods of Making and Using
Abstract
A biodegradable polymer stent with radiopacity and a method of
making and using a stent with enhanced mechanical strength and/or
controlled degradation for use in a bodily lumen is described.
Inventors: |
Su; Shih-Horng; (Irvine,
CA) ; Shulze; John E.; (Rancho Santa Margarita,
CA) ; Dutta; Debashis; (Irvine, CA) |
Correspondence
Address: |
PERKINS COIE LLP
P.O. BOX 2168
MENLO PARK
CA
94026
US
|
Assignee: |
Biosensors International
Group
Hamilton
BM
|
Family ID: |
39262617 |
Appl. No.: |
11/693684 |
Filed: |
March 29, 2007 |
Related U.S. Patent Documents
|
|
|
|
|
|
Application
Number |
Filing Date |
Patent Number |
|
|
60862939 |
Oct 25, 2006 |
|
|
|
Current U.S.
Class: |
623/1.16 |
Current CPC
Class: |
A61F 2002/91591
20130101; A61F 2/86 20130101; A61F 2002/828 20130101; A61F
2002/91566 20130101; B29C 41/14 20130101; A61F 2210/0076 20130101;
B29C 41/34 20130101; A61F 2002/91525 20130101; A61F 2/88 20130101;
A61F 2002/91558 20130101; B29K 2067/046 20130101; A61F 2/91
20130101; A61F 2002/91541 20130101; A61F 2/915 20130101 |
Class at
Publication: |
623/1.16 |
International
Class: |
A61F 2/06 20060101
A61F002/06 |
Claims
1. A polymer stent comprising: a tubular structure formed of a
polymer and including an iodinated contrast agent, said structure
comprising one or more strength modules comprising one or more
radially expandable tubular elements, said strength modules being
interconnected by one or more axial linking elements for stent
flexibility, wherein said tubular structure is radially expandable
between at least an unexpanded diameter and an expanded diameter,
and wherein at least one of the strength modules has a locking
mechanism comprising a first and a second locking member, said
first locking member being fixedly attached at a valley of the
tubular element and said second locking member being fixedly
attached at the valley of the tubular element, wherein said first
and second locking members are located opposite each other on
radially expandable tubular elements such that said first and
second locking members are not interlocked with one another when
said tubular structure is in the unexpanded diameter and said first
and second locking members are interlocked with one another when
said tubular structure is in said expanded diameter, whereby the
tubular structure is locked at said expanded diameter.
2. The polymer stent of claim 1, wherein said first and second
locking members interlock by means of teeth or barbs.
3. The polymer stent of claim 1 wherein said tubular structure is
radially expandable between said unexpanded diameter and two or
more discrete expanded diameters, wherein said tubular structure is
lockable at any of said two or more expanded diameters.
4. The polymer stent of claim 1 wherein said tubular structure is
irreversibly expandable from said unexpanded diameter to said
expanded diameter.
5. The polymer stent of claim 1, wherein said stent is formed of a
polymer selected from the group consisting of biodegradable,
bioabsorbable, and bioerodible polymers.
6. The polymer stent of claim 1, further including at least one
pharmaceutical agent incorporated in the polymer and which is
released from the polymer.
7. The polymer stent of claim 1, wherein said iodinated contrast
agent is applied as a coating at least abluminally.
8. The polymer stent of claim 1 wherein said radially expandable
tubular elements comprises at least one strut forming a
substantially sinusoidal wave structure.
9. The polymer stent of claim 8, wherein said sinusoidal wave
structure has more than one peak per circumference.
10. The polymer stent of claim 1, further comprising: the strength
module having at least two circumferential restraint bands facing
opposite of a crown valley of the expandable tubular elements, said
expandable tubular elements having four or fewer crown peaks,
wherein the length of the circumferential restraint band defines a
size of the stent when deployed and a length of each
circumferential restraint band is less than a length of the
expandable tubular elements.
11. A method of making a polymer stent with enhanced mechanical
strength comprising the steps of: a.) dip-coating a mandrel with a
solution comprising one or more biocompatible polymers to form a
polymer tube at least one of the polymers including an iodinated
contrast agent; b.) spin-drying the polymer tube around its
longitudinal axis; c.) solvent-polishing and vacuum drying the
polymer tube; d.) repeating steps a-c until the polymer tube
reaches a desired thickness; e.) necking the polymer tube by
drawing the mandrel bearing the polymer tube through one or more
necking dies of decreasing diameter, wherein said necking is
carried out at a temperature above the glass transition temperature
of the polymer and below the melting temperature of the polymer;
f.) annealing the polymer tube with an inert gas; g.) removing the
polymer tube from the mandrel; and h.) creating a design in said
polymer tube.
12. The method of claim 11, wherein said design is created by laser
cutting said polymer tube.
13. The method of claim 11, wherein the solution comprising one or
more biocompatible polymers also comprises one or more active
pharmaceutical ingredients.
14. The method of claim 11, wherein the solution comprising one or
more biocompatible polymers is used in all repetitions of the
dip-coating step.
15. The method of claim 11, wherein said solution comprising one or
more biocompatible polymers comprises at least two solutions and
said repeating step comprises dip-coating the mandrel in a
different solution in each repetition.
16. A polymeric stent, comprising: a plurality of central lobes of
approximately the same size arranged in succession at spaced
intervals longitudinally defining a stent axis, there being a
leading end and a trailing end for each central lobe, the trailing
end of each central lobe, other than the last in the succession,
being connected to the leading end of the next successive central
lobe; a plurality of peripheral lobes adjoining each central lobe
regularly spaced circumferentially about each respective central
lobe, there being a leading peripheral lobe and a trailing
peripheral lobe for each central lobe, each leading peripheral lobe
adjoining the leading end of its corresponding central lobe, each
trailing peripheral lobe adjoining the trailing end of its
corresponding central lobe; and a plurality of longitudinal rods
attached to the central lobes at one or more points around the
periphery of the stent; wherein said at least a portion of at least
one of the central lobes, peripheral lobes, and longitudinal rods
are formed of a polymer including an iodinated contrast agent.
17. The polymer stent of claim 16, where in the central lobes and
peripheral lobes are formed of a continuous polymeric fiber.
18. The polymer stent of claim 18, wherein the polymeric fiber is
made by thermal extrusion.
19. The polymer stent of claim 17, wherein the polymeric fiber has
a single-fiber construction.
20. The polymer stent of claim 17, where the polymeric fiber has a
multiple-fiber ply construction.
21. The polymer stent of claim 17, wherein the polymeric material
of the fiber is selected from the group consisting of
bioresorbable, bioabsorbable, and bioerodible.
22. The polymer stent of claim 30, wherein the polymeric material
of the fiber comprises Poly-L-lactide (PLLA).
23. The polymer stent of claim 17, wherein the polymeric fiber cord
is impregnated with one or more active pharmaceutical ingredients
adapted to be released over time in a controlled manner into
adjacent tissue when the stent is implanted in a host patient.
24. The polymer stent of claim 16, wherein the stent has a furled
small-diameter state and an expanded large diameter state, and
wherein the peripheral lobes are only present in the furled
small-diameter state having been merged into the central lobes
during expansion to the large-diameter state.
25. The polymer stent of claim 24, wherein the peripheral lobes are
disposed exterior to the central lobes in the furled small diameter
state.
26. The polymer stent of claim 24, wherein the peripheral lobes are
confined within the central lobes in the furled small diameter
state.
27. The polymer stent of claim 17, wherein at least a portion of
the fiber includes an exterior coating that carries one or more
active pharmaceutical ingredients.
28. The polymer stent of claim 17, wherein the fiber generally
defines a helix in the expanded state.
29. A method of delivering a non-metallic stent to a vessel lumen
of a host, comprising; providing a continuous cord of non-metallic
material and including an iodinated contrast agent; winding the
cord to define an elongated stent having multiple successive coils,
each coil having a central lobe and a plurality of peripheral
lobes; attaching at least one longitudinal support element to the
stent at circumferentially spaced intervals, the longitudinal
support elements extending along the length of the stent; inserting
a balloon into the stent; positioning the stent at an implant site;
expanding the balloon to expand the stent while employing the
peripheral lobes to add circumferential length to the central
lobes, thus increasing the diameter of the stent; collapsing the
balloon; and withdrawing the balloon to leave the stent in place at
the implant site.
30. The method of claim 29, further comprising delivering at least
one therapeutic agent to host tissue adjacent to the stent from
agents impregnated in the stent in a time controlled manner.
Description
CROSS-REFERENCE TO RELATED APPLICATIONS
[0001] This application claims the benefit of U.S. Provisional
Application No. 60/862,939 filed Oct. 25, 2006, which is
incorporated herein by reference.
TECHNICAL FIELD
[0002] The present application relates to a biodegradable polymer
stent with radiopacity and a method of making and using the
stent.
BACKGROUND
[0003] A stent is an endoprosthetic implant, generally tubular in
shape, typically having an open or latticed tubular construction
which is expandable to be inserted into an anatomical lumen to
provide mechanical support to the lumen and to maintain or to
re-establish a flow channel within said lumen. Stents are known for
use in blood vessels such as the aorta, carotid artery, or coronary
artery or arteries, to treat arterial blockage or aneurysm, for
example. In additions, stents are known for use in maintaining
patency of body lumens or channels besides blood vessels; these
include bile duct stents, urethral stents, and the like. As an
example, an endovascular stent may be inserted into a blood vessel
during angioplasty, and is designed to prevent early collapse of a
vessel that has been weakened or damaged by angioplasty. Insertion
of endovascular stents has been shown to reduce negative remodeling
of the vessel while healing of the damaged vessel wall proceeds
over a period of months.
[0004] During the healing process, inflammation caused by
angioplasty and stent implant injury often causes smooth muscle
cell proliferation and regrowth inside the stent, thus partially
closing the flow channel, and thereby reducing or eliminating the
beneficial effect of the angioplasty/stenting procedure. This
process is called restenosis. Blood clots may also form inside of
the newly implanted stent due to the thrombotic nature of the stent
surfaces, even when biocompatible materials are used to form the
stent. While large blood clots may not form during the angioplasty
procedure itself or immediately post-procedure due to the current
practice of injecting powerful anti-platelet drugs into the blood
circulation, some thrombosis is always present, at least on a
microscopic level on stent surfaces, and it is thought to play a
significant role in the early stages of restenosis by establishing
a biocompatible matrix on the surfaces of the stent whereupon
smooth muscle cells may subsequently migrate in and
proliferate.
[0005] Stents can be of a permanent or temporary nature. Temporary
stents which are made from biodegradable material may be
advantageous, particularly in cases of recurrent vessel narrowing
in which it is desirable to insert a subsequent stent at or near
the site of initial stent placement, or where a stent is needed
only temporarily to counteract post-surgical swelling that may
cause obstruction of a bodily lumen, such as obstruction of the
urethra after prostate surgery.
[0006] Bioabsorbable/Biodegradable/Bioerodible stents are typically
made of synthetic polymers that are biocompatible and are broken
down by biological means. Biodegradable stents are also known
wherein the outer surfaces or even the entire bulk of polymer
material is porous. For example, PCT Publication No. WO 99/07308,
which is commonly owned with the present application, discloses
such stents, and is expressly incorporated by reference herein.
[0007] Stents are also known which contain APIs (active
pharmaceutical ingredients), which are generally intended to reduce
or eliminate thrombosis or restenosis. Such APIs are often
dispersed or dissolved in either a durable or biodegradable polymer
matrix, which is applied as a coating over at least a portion of
the filament surface. After implantation, the API diffuses out of
the polymer matrix and preferably into the surrounding tissue.
[0008] A variety of agents specifically claimed to inhibit smooth
muscle-cell proliferation, and thus inhibit restenosis, have been
proposed for release from endovascular stents. Rapamycin
(sirolimus), an immunosuppressant reported to suppress both smooth
muscle cell and endothelial cell growth, has been shown to have
effectiveness against restenosis, when delivered from a polymer
coating on a stent (see, for example, U.S. Pat. Nos. 5,288,711 and
6,153,252). Also, PCT Publication No. WO 97/35575 and WO
2003/090684 describe the macrocyclic triene immunosuppressive
compound everolimus and related compounds, which have been proposed
for treating restenosis. U.S. Pat. No. 6,159,488 describes the use
of a quinazolinone derivative; U.S. Pat. No. 6,171,609 and U.S.
Pat. No. 5,716,981 describe the use of paclitaxel (taxol). U.S.
Pat. No. 5,288,711 describes the use of both heparin and rapamycin.
Tranilast, a membrane-stabilizing agent thought to have
anti-inflammatory properties is disclosed in U.S. Pat. No.
5,733,327. As described in U.S. Pat. No. 6,231,600, a mixture of
polymer and therapeutic substance can be coated onto the surface of
a stent, which is then coated with a second layer of polymer. The
first layer may contain polymer mixed with a therapeutic substance
and the second layer may contain polymer mixed with heparin. In
U.S. Pat. No. 6,939,376, Shulze et al. disclose a stent for
inhibiting restenosis, which is comprised of a stent body and a
biodegradable drug-release coating which contains poly(D,L-lactide)
polymer and an immunosuppressive compound which is eluted with time
at the vascular site of injury. U.S. Pat. No. 6,808,536 discloses
local delivery of rapamycin or its analogs from an intravascular
stent, either directly from tiny micropores or channels in the
stent body or mixed or bound to a polymer coating applied on stent,
grooves or channels which are smaller in dimension than the stent
struts. Also, U.S. Pat. No. 6,904,658 contains reference to the use
of a porous plated layer to contain and elute therapeutic drug.
[0009] It is difficult to visualize non-metal, polymer based stents
because they are radiolucent. Since optimal stent placement
requires real time visualization to allow the cardiologist to track
the stent in vivo there is a need to increase the radiopacity of
non-metallic polymer based stents. Iodinated contrast media is a
common type of intravenous radiographic dye containing iodine that
enhances the visibility of vascular structures during radiographic
procedures.
[0010] Present stents vary widely by geometry. Polymer tubular
stent blanks are generally injection molded or extruded, and then
die-cut, machined, or laser-cut into the desired geometry or
openwork. Alternatively, rolling one or more sheets of metals or
polymer may form tubular metal or polymer stent blanks. Stents may
also be composed of extruded polymer filaments that are woven into
a braid-like structure (see U.S. Pat. No. 6,368,346). To achieve
the reticular or openwork nature of the stent body, stents
generally comprise radially expandable tubular elements or "bands"
which often have a zigzag or sinusoidal structure and which are
interconnected by linking elements or "linkers" that typically run
in a generally longitudinal direction.
[0011] Steinke (U.S. Pat. No. 6,623,521) discloses a locking stent,
which may be degradable. The stent is formed from a flat sheet, or
sheets, of metal or plastic and bears sliding and locking radial
elements or struts. The radial elements may bear a ratcheting
mechanism that permits one-way sliding of the radial elements.
[0012] U.S. Pat. No. 6,022,371 (Killion) discloses a continuous
circumference tubular stent with a unitarily formed locking arm
that selectively locks the stent at a desired diameter.
[0013] U.S. Pat. No. 6,540,777 (Stenzel) discloses a stent
comprising a plurality of interconnected cells, at least one of
which is a lockable cell with a first and second locking member
which may lock with one another. Also disclosed is a stent
comprising a plurality of interconnected bands with a pincer
locking member extending toward an adjacent band having a tongue
locking member.
[0014] U.S. Pat. No. 6,156,062 (McGuinness) discloses a stent
comprising a strip of material with a groove along one edge and a
tongue along the other edge, to maintain a helical configuration.
No locking mechanism is disclosed.
[0015] Application Serial No. U.S. 2004/0249442A1 (Fleming)
discloses a stent comprising a lattice having a closed and an open
configuration. The lattice is composed of hoops or struts that
interlock with one another while moving from a closed to an open
configuration, and the hoops interlock with one another by means of
teeth on the struts.
[0016] U.S. Pat. No. 6,368,346 (Jadhav) discloses biocompatible and
biodegradable stents made of blended polymers.
[0017] U.S. Pat. No. 5,441,515 (Khosravi) discloses a ratcheting
stent comprising a cylindrical sheet having overlapping edges that
interlock. The stent may be biodegradable and may be
drug-releasing.
[0018] U.S. Pat. No. 5,817,328 and U.S. Pat. No. 6,419,945 disclose
buffered resorbable internal fixation polymer devices for bone
repair.
[0019] U.S. Pat. No. 6,932,930 discloses method to make synthetic
polymer strong for stent application.
[0020] Unlike traditional metal stents, biodegradable stents are
capable of bulk loading of multiple APIs and are temporary
implants. Biodegradable stents, however, have typically suffered
from insufficient mechanical strength and/or undesirable
physical/mechanical elastic polymer recoil. In addition, the
degradation time of such stents has been uncontrolled, being
dependent mainly on the molecular weight of the polymer resin used.
The present stents and methods provide means to adjust the polymer
degradation rate and/or to enhance the mechanical strength of the
polymer tube or fiber used for biodegradable stent fabrication.
BRIEF DESCRIPTION OF DRAWINGS
[0021] FIG. 1 is a perspective illustration of the
three-dimensional structure of an expanded fiber stent;
[0022] FIG. 2 is a perspective illustration of a side view of the
fiber stent of FIG. 1;
[0023] FIG. 3 is a cross-sectional side illustration of the fiber
stent of FIG. 1;
[0024] FIG. 4 is an enlarged cross-sectional illustration of the
fiber stent of FIG. 1;
[0025] FIG. 5 is a plan illustration of a process for manufacturing
a tube stent;
[0026] FIG. 6 is a perspective illustration of the
three-dimensional structure of an expanded tube stent with a
circumferential restraint mechanism facing opposite of the crown
valleys;
[0027] FIG. 7 is a side illustration of the tube stent of FIG.
6;
[0028] FIG. 8 is a plan illustration showing an enlargement of the
tube stent of FIG. 6;
[0029] FIG. 9 is a scanned image of a tube stent that is crimped
onto a balloon catheter;
[0030] FIG. 10 is a scanned image of a tube stent in an expanded
state;
[0031] FIG. 11 is a plan illustration of a stent having an axial
locking design;
[0032] FIG. 12 is a table graph comparing the compression extension
as measured by the average load at compression extension (cm (N))
of pure PLLA (.box-solid.) and phosphate salt containing PLLA
(.quadrature.) tubes at 0 and 5 months of pre-immersion in
water.
[0033] FIG. 13 is a radiograph image of a PLLA stent (#1), a stent
formed of iohexol in PLLA at 26 wt. % (#2), a stent formed of
iohexol in PLLA at 50 wt. % (#3), and a PLLA stent coated with
iohexol (#4) and a guiding catheter.
[0034] FIG. 14 is a radiograph image of a PLLA stent (#1 and #4), a
stent coated with iohexol and BA9-PLLA after 2 minute water
immersion (#2) and after 30 seconds of water immersion (#3) and a
guiding catheter.
SUMMARY
[0035] The present stent preferably has at least one of the
following features: (1) it has an all-polymer construction with
similar mechanical function to conventional metallic stents; (2) it
is expandable with an expansion ratio that can be customized to
meet various needs; (3) it can be deployed at body temperature with
low inflation pressure (3 atm); (4) it is a temporary, fully
biodegradable implant; (5) it has a regulated degradation rate due
to biocompatible buffers that accelerates hydrolysis; (6) it may be
a local drug or gene delivery device; (7) it may be a local
radiation therapy device; and (8) it can include fibers with
various functions (mechanical support, acute drug burst, long-term
drug release, etc.), enabling a variety of treatment options
including multiple functions with a single stent and using a single
stent-implant procedure.
[0036] An alternative embodiment preferably has at least one of the
following features:
[0037] (1) it has an all-polymer construction with similar
mechanical function to conventional metallic stents; (2) it is
expandable with an expansion ratio that can be customized to meet
various needs; (3) it can be deployed at body temperature with low
inflation pressure (3 atm); (4) it is a temporary, fully
biodegradable implant; (5) it has a regulated degradation rate due
to biocompatible buffers that accelerates hydrolysis; (6) it may be
a local drug or gene delivery device; (7) it may be a local
radiation therapy device; (8) it has enhanced mechanical strength
by forming due to dip-coating and necking process; (9) it is formed
at a low temperature, below the melting temperature of the polymer
and right above the glass transition temperature of the polymer;
and (10) it can include fibers with various functions (mechanical
support, acute drug burst, long-term drug release, etc.), enabling
a variety of treatment options including multiple functions with a
single stent and using a single stent-implant procedure.
[0038] An alternative embodiment preferably has at least one of the
following features: (1) it has an all-polymer construction with
similar mechanical function to conventional metallic stents; (2) it
is expandable with an expansion ratio that can be customized to
meet various needs; (3) it can be deployed at body temperature with
low inflation pressure (3 atm); (4) it is a temporary, fully
biodegradable implant; (5) it has a regulated degradation rate due
to biocompatible buffers that accelerates hydrolysis; (6) it may be
a local drug or gene delivery device; (7) it may be a local
radiation therapy device; (8) it can include a temporary iodinated
contrast agent for increased visibility; and (9) it can include
fibers with various functions (mechanical support, acute drug
burst, long-term drug release, etc.), enabling a variety of
treatment options including multiple functions with a single stent
and using a single stent-implant procedure.
[0039] In one aspect, the stent is a temporary implant. The
temporary implant permits the stress against the vessel wall to be
decreased, where subsequent intervention is not necessary
especially for young people and vulnerable patients, such as
diabetics.
[0040] In another aspect, the stent is capable of delivering
therapeutic agents incorporated in the stent body and/or coated on
the polymer surface.
[0041] A further aspect of the present stent is that it is possible
to vary applications and control degradation of the stent by
selection of the polymer composition, the polymer molecular weight,
fiber cord diameter and processing conditions, thus controlling the
degradation rate, drug release rate and period of mechanical
support.
[0042] An additional aspect of the present stent is that it has
improved radiopacity that allows the stent to be visibly tracked
during interventional procedures.
[0043] These and other aspects and advantages of the present
invention will become apparent to those of ordinary skill in the
art from the following detailed description of the preferred
embodiment when considered in conjunction with the accompanying
drawings in which like numerals in the several views refer to
corresponding parts.
DETAILED DESCRIPTION OF THE INVENTION
I. Definitions
[0044] The following terms have the definitions given herein,
unless indicated otherwise "Inhibiting restenosis" means reducing
the extent of restenosis observed following a vascular
"overstretch" injury, as measured by a reduction in average
percentage of vascular stenosis at a selected time following stent
placement, e.g., 1-6 months.
[0045] "Radiopaque" refers to a material that prevents the passage
of electromagnetic radiation making the material fluoroscopically
visible under x-rays.
II. Stents
[0046] A. Materials
[0047] The present stents are formed of one or more polymer(s) or
co-polymers. In an embodiment, and as described further below, the
stent body is formed of a plurality of linked tubular members by
filaments. The stent body may be formed of biocompatible polymers
which may be biodegradable including, but not limited to;
bioresorbable, bioabsorbable, or bioerodible. A variety of natural,
synthetic, and biosynthetic polymers are biodegradable. Generally,
polymer backbones that contain chemical linkages such as anhydride,
ester, or amide bonds, among others are biodegradable
(www.sigmaaldrich.com). The mechanism for degradation is generally
by hydrolysis or enzymatic cleavage of these bonds that results in
division of the polymer backbone. Bioerosion of polymers generally
works by conversion of the polymer that is at least partly
insoluble water into one that is at least partly water-soluble.
When the polymer is admixed with a therapeutic agent, as the
polymer surrounding the drug is eroded, the drug is released.
[0048] In some embodiments the stent is formed of a biodegradable
polymer. The rate of biodegradation may be controllable by a number
of factors including, without limitation, the degree of
crystallinity, the material molecular weight, and the use of
biocompatible buffers which mitigate pH change at degradation
site(s) and accelerate hydrolysis via dissolving pathways. In some
embodiments the polymer stent releases one or more therapeutic
agents, which may be released in a desired order and at a desired
rate.
[0049] The use of biodegradable materials allows the stent to be
decomposed and resorbed in tissues and/or be absorbed by the cells.
Such materials include, but are not limited to, polymers of the
linear aliphatic polyester and glycolide families, as discussed
below. Other materials contemplated for the stent embodiments of
the present invention include biocompatible polymers, such as of
the type from the polyethylene, polyester and polypropylene
families and plastics such as a polymer from the linear aliphatic
polyester family. Exemplary polymers include, but are not limited
to, poly(lactic acid), poly(glycolic acid) or polycaprolactone, and
their associated copolymers, degradable polymers such as
polyorthoester, polyanhydride, polydioxanone and
polyhydroxybutyrate or combinations thereof.
[0050] Representative bio-compatible absorbable polymers include
poly(lactic acid)(PLA), poly(L-lactic acid), poly(D,L-lactic acid),
polyglycolic acid (PGA), poly(D-lactic-co-glycolic acid),
poly(L-lactic-co-glycolic acid), poly(D,L-lactic-co-glycolic acid),
poly(.epsilon.-caprolactone), poly(valerolactone),
poly(hydroxybutyrate), polydioxanone, poly(hydroxyl butyrate),
poly(hydrovalerate), etc., including copolymers such as polyglactin
(a co-polymer of lactic acid and glycolic acid (PGA-PLA)),
polyglyconate (a co-polymer of trimethylene cargonate and
glycolide), a co-polymer of polyglycolic acid and
.epsilon.-caprolactone, a co-polymer of poly(lactic acid) and
.epsilon.-caprolactone, poly(lactic acid)-poly(ethylene glycol)
block co-polymer, and poly(ethylene
oxide)-poly(butyleneteraphthalate), poly(lactic
acid-co-trimethylene carbonate), poly(.epsilon.-caprolactone
copolymer), poly(L-lactic acid copolymers), etc. It will be
appreciated that biodegradable stents may be made from single
polymers or co-polymers (for example, a co-polymer of L-lactide and
.epsilon.-caprolactone as described in U.S. Pat. No. 5,670,161 or a
terpolymer of L-lactide, glycolide and .epsilon.-caprolactone as
described in U.S. Pat. No. 5,085,629. Biodegradable stents may also
be formed of blended homopolymers such as those described in U.S.
Pat. No. 6,368,346, including blends having similar compositions to
the above copolymers. Homopolymers, blended polymers and
co-polymers may have different characteristics including varying
susceptibility to hydrolytic decomposition and thus may be
preferred under circumstances in which faster or slower absorption
is desired.
[0051] Another distinct advantage of polymer stents is that they
are more compatible to MRI imaging since the polymer is not a
ferromagnetic material. This property makes polymers less likely to
cause signal loss during the imaging process and maintain the
vessel lumen visibility. Further, subsequent analysis may be
performed non-invasively.
[0052] The stents may further include a non-toxic radiopaque
marker, such as, for example, barium sulfate or bismuth trioxide,
into the polymer prior to stent formation, as disclosed in U.S.
Pat. No. 6,368,356 to increase the radiopacity of the stents. In a
preferred embodiment, the stent is coated with one or more
radiopaque layers of non-ionic, water-soluble, iodinated contrast
medium having a molecular weight of approximately 1 milligram
(+/-20%) and a thickness of about 0.5 to 5 microns. In an
embodiment, the radiopaque coating is thin and temporary as by
bioabsorption. Preferably, the iodinated contrast is water-soluble
for faster absorption by body tissue. Typically, the contrast media
also has a low osmolality to reduce tonicity, chemical toxicity,
hypersensitivity, and other potentially adverse side effects. It
will be appreciated that a combination of contrast agents may be
used in the same layer or in separate layers. It will further be
appreciated that one or more contrast agents may be included in the
stent polymer and one or more different or same contrast agents may
be coated on the stent. Alternatively, the contrast media may be
hydrophobic. Hydrophobic contrast media may be utilized in
applications that require slower rates of degradation or excretion.
Examples are disclosed in U.S. Pat. No. 7,008,614 to Kitaguchi.
[0053] Previous stents have included heavy metal coatings to confer
radiopacity. These coatings, however, do not disappear once the
biodegradable stents are absorbed by the tissues. The present
radiopaque coatings create a biodegradable stent with temporary
radiopacity without introducing permanent and harmful materials
into the human body. In an embodiment, after about 2-3 minutes,
there is little or no radiopaque material left in the tissues.
[0054] Contrast agents that may be used to confer biodegradable
stents with radiopacity include, but are not limited to, iopamidol,
iohexol, iopromide, and iodixanol. The molecular structure of these
agents provides both comfort to the patient and needed visibility
for the cardiologists. In general, the chemical structures consist
of a hydrophobic region masked with hydrophilic regions that
increase solubility and decreases binding with blood or other
vascular constituents. A preferred contrast agent is iohexol.
[0055] The selected radiopaque compounds may be coated directly on
a polymer stent, included within a polymer coating, impregnated
within the stent structure, sandwiched between the stent and a
further coating, or any combination of these techniques.
Preferably, the selected radiopaque compounds are incorporated
within a biodegradable polymer(s).
[0056] Where the contrast agent is coated on the stent, the agent
may be formulated as a solution in a solvent such as a
methanol-based solvent as described in Example 3. The solution is
then applied to a stent with any appropriate method such as
spraying. The solvent is evaporated to leave a thin covering of the
radiopaque material over at least a portion of the stent. In a
preferred embodiment, the abluminal stent surface is completely
covered. The resulting stent is radiopaque under typical
visualization techniques. As seen in FIG. 13 (#4), a stent coated
with iohexol was visible with imaging.
[0057] The stent embodiment also optionally includes a process for
combining the radiopaque compound(s) with the biodegradable
polymer. The layer of contrast media can be applied by spray
coating, dip-coating, co-extrusion, compress molding,
electroplating, painting, plasma vapor deposition, sputtering,
evaporation, ion implantation, or use of a fluidized bed. In a
preferred embodiment the polymer backbone of the stent is
impregnated with the contrast agent, and the drug is subsequently
applied on top using one of the aforementioned processes. As
described in Example 5, the contrast agent is suspended in a
solution with the polymer to a desired final weight amount. As seen
FIG. 13 (#2 and #3), the radiopacity of the impregnated stents
appears to increase with an increase in the weight percentage of
contrast agent. In an alternate embodiment, the temporary iodinated
contrast agent is sandwiched between the drug on the stent backbone
and a thin polymer (PLLA or other biodegradable polymer) on the
abluminal surface. In an exemplary method as described in Example
4, the polymer stent is first coated with the contrast agent. The
stent is subsequently coated with the therapeutic agent or the
therapeutic agent in a polymer coating.
[0058] In an exemplary embodiment, iohexol, a type of iodinated
contrast media, is used as the radiopaque material. As seen in
Example 5, iohexol was incorporated in poly(L-lactic acid) (PLLA)
polymers and coated on a stent. These stents remained radiopaque
after exposure to water at thirty seconds and two minutes.
[0059] B. Active Pharmaceutical Ingredients
[0060] The stent of the present invention may also be used to
deliver one or more APIs. These agents may be released from the
stent in desired sequence and with controllable timing in order to
have a desired effect on host cell responses. Various types of APIs
may be mixed with the polymer solution at desired weight
percentages from 0.1 wt % to 55 wt %. The present stents are
manufactured without the extreme heating needed for polymer
extrusion methods, thus decreasing the likelihood of heat
inactivation of any temperature-sensitive therapeutic agents. In an
embodiment, the biodegradable polymer is admixed with one or more
of a variety of APIs or therapeutic agents. In other embodiment,
the API's may be deposited on the fiber surface or into the lumen
of the hollow fibers alone or in combination with APIs admixed in
the stent polymer. It will be appreciated that by adding a
fluoroscopic impermeable agent at the time of spinning the fibers
or assembling the stents, the status of the introduced luminal
stent may be observed with conventional fluoroscopic equipment.
Non-limiting examples of therapeutic agents useful with the present
stent include anti-restenosis drugs, anti-proliferative drugs,
immunosuppressive compounds, anti-thrombogenic drugs,
anti-fibrotic/fibrinolytic compounds, and cytotoxic compounds. In
preferred embodiments the agents are anti-restenosis,
anti-proliferative drugs such as rapamycin (sirolimus), everolimus,
paclitaxel, zotarolimus, Biolimus A9.RTM., pimecrolimus and
tacrolimus, anti-thrombogenetic/anti-coagulate drugs such as
heparin/enoxaparin/low-molecular-weight heparin,
hirudin/bivalirudin/lepirudin/recombinant hirudin, aprotinin,
clopidogrel, prasugrel, argatroban, anti-fibrotic or fibrinolytic
drugs such as tranilast, colchicine, streptokinase, two-chain
urokinase-type (tcu- plasminogen activator, urokinase), tissue-type
plasminogen activator PA (t-PA), and single-chain urokinase-type PA
(scu-plasminogen activator). If the polymer is biodegradable, in
addition to release of the drug through the process of diffusion,
the API may also be released as the polymer degrades or resolves,
making the agent more readily available to the surrounding tissue
environment. When biodegradable polymers are used as drug delivery
coatings, porosity is variously disclosed to aid tissue ingrowth,
make the erosion of the polymer more predictable, and/or to
regulate or enhance the rate of drug release, as, for example,
disclosed in U.S. Pat. Nos. 6,099,562, 5,873,904, 5,342,348,
5,707,385, 5,824,048, 5,527,337, 5,306,286, and 6,013,853.
[0061] C. Mechanical strength
[0062] The present stent has not only structural benefits to the
stent, and thus to the patient, but also allows the stent to be
manufactured with less polymer material, which has advantages of
cost as well as of decreasing the exposure of the patient to
foreign material. The enhanced mechanical strength is provided via
highly oriented polymer molecules.
[0063] As described further below and depicted in FIG. 5, the tube
stents may be formed by a process for making a polymer stent with
enhanced mechanical strength. The process includes the spin drying
240 and necking 260 steps that orient the polymer chains of the
tube in the radial and axial directions, respectively. Furthermore,
the steps of the process involve do not require extreme heating
that is necessary for heat extrusion techniques typically used to
form stents. The moderate manufacturing temperature conveys
advantages especially where a temperature sensitive API is to be
delivered via biodegradable tube stent 200. For example, heat
extrusion typically used to manufacture polymer stents must be
performed at temperatures above the melting temperature (T.sub.m)
of the polymer, which in the case of the polymer poly-L-lactic acid
is approximately 173.degree. C. In contrast, the necking 260
process of the present invention is carried out at a temperature
between the glass transition temperature (T.sub.g) and the T.sub.m,
or approximately 55.degree. to 60.degree. C. for poly-L-lactic
acid, and the manufacturing steps which precede the necking 260
step are carried out at room temperature.
[0064] It will be appreciated that the fibers of the fiber stent
may also be formed by this process. In another embodiment, the
mechanical strength of the fiber stent may be enhanced by the
configuration of the fibers, described further below, and/or by the
manufacturing process.
[0065] D. Buffers
[0066] The stent of the present invention may further be
manufactured in such a way that the stent resorption rate can be
controlled. In one embodiment, this is accomplished by addition of
buffer salts to alter the stent resorption rate. In this
embodiment, one or more buffers, including but not limited to, a
phosphate buffer salt, a citrate buffer salt, or NaCl buffer may be
loaded in the polymer solution alone or in conjunction with one or
more APIs in order to adjust the degradation of the polymer, and
thus, the stent. Without being limited as to theory, it is thought
that the buffer that is incorporated into the polymer quickly
diffuses out of the stent once the stent contacts fluid, thus
creating microscopic holes or channels. Water molecules can then
permeate the stent through those holes or channels. For example,
PLLA polymer decomposition is hydrolysis-driven and subject to the
influence of water content. Resorption of the polymer occurs when
the long molecular chain is broken down into many single molecules
forming lactic acid and then nearby cells uptake the lactic acid.
Thus, controlling the amount of buffer powders loaded into the
polymer solution, the buffer salt diffusion rate and the stent
resorption rate are controllable.
[0067] The concentration of buffer salt is generally from about
0.01% to 15% wt % of the stent; preferably 0.01% to 10% wt %; more
preferably 2% to 8% wt %. As described in Example 1, a polymer
stent was formed of PLLA with 6% by weight phosphate salt.
Referring to FIG. 12, the PLLA tube stent with phosphate salt
buffer at 6% wt % resulted in a substantially faster degradation
measured at five months as compared to a PLLA tube stent without
buffer.
[0068] E. Stent Geometry
[0069] 1. Fiber Stent
[0070] FIG. 1 shows a fiber stent 100 constructed in accordance
with the invention made from one or more polymer fibers.
Preferably, the fiber stent is a luminal stent consisting of a
tubular member produced by "knitting" a biodegradable polymer yarn,
fiber, or cord. Such a fiber stent is preferably deployed from a
luminal stent deployment device comprising the luminal stent which
is fitted over a balloon forming portion in the vicinity of a
distal end of a delivery catheter.
[0071] In a preferred embodiment, the fibers are "biodegradable
fibers" that can be decomposed and resorbed by tissues. The fibers
can be solid, hollow, or a combination of solid and hollow.
Preferably, the fibers are decomposed within about 1 to 60 months
after insertion into the body, more preferably in about 3 to 15
months, even more preferably in about 6 to 12 months. The
biodegradable fibers may be formed of biodegradable materials as
described above. In a preferred embodiment, the polymer fibers are
formed of PLLA. In a non-limiting embodiment, these fibers are
generally a filament thread of about 5 to about 1,000 .mu.m in
diameter. Preferably, the filament thread is sized such that a
stent composed of these fibers is firm enough to easily maintain a
cylindrical form. Monofilament threads are particularly preferred
for use herein. The mean molecular weight of preferred
biodegradable polymers is between about 10,000 to about 800,000 DA.
It will be appreciated that selection of the biodegradable polymer
may depend on the total radial strength necessary to support
various sized vessel lumen. The polymer fiber may be formed by any
suitable means. In one embodiment, the polymer fiber is formed by
thermal extrusion as known in the art. In another embodiment, the
fiber is formed by the method as described above and illustrated in
FIG. 5.
[0072] It will be appreciated that at least some of the stent's
polymer fibers may be admixed with one or more APIs as described
above. It will further be appreciated that at least some of the
polymer fibers may be coated with one or more APIs.
[0073] In some embodiments, the polymer fibers used for the stent
fabrication are loaded with a non-steroid type anti-inflammation
agent, such as turmeric alone or in combination with further
API(s). The turmeric-loaded fibers significantly reduce
inflammation at the stent implant site by reducing the adhesion of
inflammatory cells. The impregnated or coated APIs can be prepared
in doses that are controllably delivered over a predetermined time
period.
[0074] Furthermore, by taking advantage of the fact that the fiber
stent produced from biodegradable polymer fibers fully degrades
after a predetermined time from the site into which it has been
introduced, carcinostatics or anti-thrombotic agents may be mixed
into or attached to the fibers for concentrated administration of
these agents to the site of lesion.
[0075] The fiber stent of the present invention provides adequate
mechanical support for the vessel lumen following the
interventional procedure. Further, the fiber stent, by being
absorbed over controllable periods of time, avoids chronic
mechanical disturbance of the vessel wall. The residual stress
against the vessel wall is eliminated while the stent is hydrolyzed
and the fibers are endothelialized. The fiber stent APIs are
released in a controlled fashion during hydrolysis and effective
concentrations at target lesions are maintained.
[0076] The fiber stent may be introduced into and placed at the
site of angioplasty by a catheter fitted with a balloon and
deployed by dilating the balloon or any other method as known in
the art. The fiber stent may retain its shape for several weeks to
several months, usually about 2 months to about 2 years, after
placement and hydrolyze in several months, usually about 6-12
months. It will be appreciated that the stent may hydrolyze over a
longer period of time such as 2 years.
[0077] The methods of using the fiber stent are intended to provide
structural support and, optionally, local drug administration to
the interior of a body lumen. In one embodiment, the methods are
designed to minimize the risk and/or extent of restenosis in a
patient who has received localized vascular injury, or who is at
risk of vascular occlusion. Typically the vascular injury is
produced during an angiographic procedure to open a partially
occluded vessel, such as a coronary or peripheral vascular artery.
In the angiographic procedure, a balloon catheter is placed at the
occlusion site, and a distal-end balloon is inflated and deflated
one or more times to force the occluded vessel open. This vessel
expansion, particularly involving surface trauma at the vessel wall
where plaque may be dislodged, often produces enough localized
injury that the vessel responds over time by inflammation, smooth
muscle cell proliferation leading to positive remodeling, and
reocclusion. Not surprisingly, the occurrence or severity of this
process, known as restenosis, is often related to the extent of
vessel stretching and injury produced by the angiographic
procedure. Particularly where overstretching is 35% or more,
restenosis occurs with high frequency and often with substantial
severity, i.e., vascular occlusion.
[0078] The fiber stent is typically placed in its contracted state
typically at the distal end of a catheter, either within the
catheter lumen, or in a contracted state on a distal end balloon.
The distal catheter end is then guided to the injury site, or the
site of potential occlusion, and released from the catheter, e.g.,
by using a trip wire to release the stent into the site if the
stent is self-expanding, or by expanding the stent on a balloon by
balloon inflation, until the stent contacts the vessel walls, in
effect, implanting the stent into the tissue wall at the site.
[0079] FIGS. 1-4 show an exemplary fiber stent. Referring now to
FIG. 1, the fiber stent 100 is comprised of coiled fiber material.
The fiber material is a polymer fiber or ply of multiple polymer
fibers as described above. Preferably, the polymer comprises PLLA.
The use of PLLA to construct the fiber stent is advantageous
because it is biodegradable. The degradation mechanism of the fiber
stent is generally via hydrolysis at the ester bonds. Degradation
may occur over a period of about three months to three years,
depending on several factors, in particular, the molecular weight
of the polymer and the type of buffer employed. PLLA is also
advantageous because it may be impregnated and/or coated with drugs
or other therapeutic agents for local treatment of tissue at the
stent implant site. It will be appreciated that other biodegradable
polymers will have the advantages as described for PLLA.
[0080] The fiber material is coiled to form at least one large
central lobe 160 that is further comprised of a plurality of
peripheral lobes 180 within the large central lobe 160 and
connecting segments 130 disposed between the peripheral lobes 180.
In a preferred embodiment, the plurality of peripheral lobes
comprises at least three peripheral lobes per central lobe. In an
alternative configuration, the peripheral lobes 180 may be disposed
on the abluminal side of the large central lobe 160 (not shown).
The arbitrary bands 120 define the putative starting and ending
point of each of the peripheral lobes 180. The large central lobes
160 form the super structure of the fiber stent 100. At least three
longitudinal rods 170 are attached on the abluminal surface of the
large central lobes 160, preferably using a viscous PLLA-chloroform
solution. The longitudinal rods 170 may be composed of the same
material as the large central lobes 160 and peripheral lobes 180.
In the embodiment shown in FIG. 1, the stent comprises nine central
lobes 160 formed of three peripheral lobes 180 linked by connecting
bands 130. The peripheral lobes are disposed on the luminal side of
the central lobe. The stent further comprises at least one
longitudinal or reinforcing rod 170 disposed on the abluminal
surface of the fiber stent. Preferably two or more longitudinal
rods are disposed on the abluminal surface of the fiber stent. More
preferably, three or more longitudinal rods are disposed on the
abluminal surface of the fiber stent. The longitudinal may be
attached to one or more of the central lobes at multiple points
along the stent. As seen in the figure, the central lobes are
approximately the same size and are arranged in succession at
spaced intervals to define the stent longitudinal axis. Each
central lobe has a leading end 162 and a trailing end (not shown).
Except for the first and last central lobes, the trailing end of
each central lobe is connected to the leading end of the next
successive central lobe. It will be appreciated that the stent may
be formed of a continuous fiber where the trailing end of each
central lobe leads continuously into the leading end of the next
successive central lobe. As further seen in the figure, the
peripheral lobes may be regularly or substantially regularly spaced
about the circumference of the central lobe. In another embodiment,
the peripheral lobes may be irregularly spaced about the
circumference of the central lobe (not shown). The plurality of
peripheral lobes further includes a leading peripheral lobe
following the central lobe leading end and a trailing peripheral
lobe prior to the central lobe trailing end. One or more additional
peripheral lobes may further be positioned between the leading and
trailing peripheral lobes. Preferably, the leading peripheral lobe
adjoins the leading end of the central lobe and the trailing
peripheral lobe adjoins the trailing end of the central lobe.
[0081] FIGS. 2-4 illustrate an alternative embodiment poly-layered
fiber stent 110 wherein, the large central lobes 161, peripheral
lobes 181, and the longitudinal rods 171 may comprise a
multiple-fiber ply material. For example, the large central lobes
161 and peripheral lobes 181 may be formed from a double-fiber or
higher ply material, and each of the at least three longitudinal
rods 171 may be formed from a triple-fiber ply material for added
rigidity. Further, the poly-layered fiber stent 110 may have a
hollow lumen 150 that is capable of storing APIs for release after
implantation.
[0082] Also, by way of example, the length of a preferred
embodiment is 18 mm and the initial diameter is 1.9 mm. In this
embodiment, the final diameter, after balloon expansion, is
preferably about 3.25 mm. Preferably, the stent length is about 8
mm to about 30 mm. In some embodiments, the stent length is up to
about 60 mm. Typical coronary artery diameters are about 2 mm to
about 4 mm and the expanded diameter of coronary stents is
generally suitably dimensioned. It will be appreciated that other
body lumens can have diameters up to about 1 cm and stents for
these lumens have a suitable expanded diameter. The length of the
fiber stent 110 can be increased by increasing the number of large
central lobes 161 and peripheral lobes 181. The peripheral lobe 181
and large central lobe 181 diameters may be adjusted to set the
final diameter of the stent. For example, coronary stents
commercially available have a final expanded diameter range of 2-5
mm. It will be appreciated that the stent may be appropriately
sized for the lumen and/or application.
[0083] In practicing the present invention, the stent is placed in
its contracted state where the central lobes of the stent are in a
furled, small diameter state. The stent is typically at the distal
end of a catheter, either within the catheter lumen, or in a
contracted state on a distal end balloon. The distal catheter end
is then guided to the injury site, or the site of potential
occlusion, and released from the catheter, e.g., by using a trip
wire to release the stent into the site, if the stent is
self-expanding, or by expanding the stent on a balloon by balloon
inflation, until the stent contacts the vessel walls, in effect,
implanting the stent into the tissue wall at the site.
[0084] Once deployed at the site, the stent begins to release
active compound into the cells lining the vascular site, to inhibit
cellular proliferation.
[0085] 2. Tube Stent
[0086] In another embodiment, the stent is a biodegradable polymer
tube stent. Typically, the stent is formed in a cylindrical sheet
according to the process as illustrated in FIG. 5. Designs may be
laser cut 280 from the tubes using excimer laser technology with a
wavelength of less than about 310 nm.
[0087] The biodegradable polymer tube is built layer by layer on a
mandrel 210. Typically, the mandrel is a Teflon.RTM. mandrel or
Teflon.RTM.-coated metal mandrel. It will be appreciated that other
mandrels or structures that support the stent form can be used. A
biodegradable polymer solution 220 is made by dissolving the
biodegradable polymer resins in a suitable solvent, such as (but
not limited to) chloroform or dioxane. The solution viscosity is
generally from about 1 to about 2000 centipoise; preferably from
about 10 to about 500 centipoise.
[0088] Buffers such as phosphate buffer or citrate buffer salts may
be loaded in the polymer solution alone or in conjunction with one
or more APIs. Other biocompatible buffered salts are readily known
to those in the medical art, including, but not limited to Ringers
solution and lactose.
[0089] The tube stent is formed by dip coating 230 the mandrel into
the biodegradable polymer solution one or more times until the
polymer coating is a desired thickness. The dip 230 coating speed
generally ranges from about 1 millimeter/minute to about 10
meter/minute. A typical speed range is from 1 meter/minute to 5
meter/minute. The coated mandrel is then dried 240. Preferably, the
coated mandrel is spin dried 240 around the longitudinal axis of
the mandrel in a laminar flow hood, leaving a thin polymer layer
upon evaporation of solvent. The speed of spin drying 240 can be
from about 1 to about 100,000 rpm with a typical range being from
about 100 to about 4000 rpm. The spin drying 240 step enhances the
radial strength of the polymer tubing via circumferential
orientation. The orientation of the spin in the drying step 240 may
be repeated in one direction or the orientation may alternate, for
instance from the clockwise direction to the counter-clockwise
direction. This polymer layer is then solvent polished 250 and
dried 240 one or more times, leaving behind a layer of thin and
smooth polymer tubing (not shown). Preferably, solvent polishing
250 uses pure solvent, which can be the same or different from the
solvent used for preparing the polymer solution 220. A typical
solvent is chloroform. The whole cycle, including dip 230 coating,
spin drying 240 and solvent polishing 250, is preferably performed
at or about room temperature (from 10.degree. C. to 30.degree. C.).
However, it will be appreciated that a temperature range of about
-20.degree. C. to 80.degree. C. is possible. The cycle of dipping,
drying, and polishing is preferably repeated in order to increase
the thickness of the tubing until a desired thickness, 0.0875
millimeter to 1.25 millimeter, for example, is reached. In one
embodiment, the cycle is repeated until at least about 46 layers of
polymer are laid down to form a complete tube. As described above,
the various cycles need not be conducted with equivalent polymer
solutions 220. The polymer, buffer and/or the API may be varied in
nature or concentration from layer to layer, as illustrated above
with sequential dip 230 coatings.
[0090] The polymer tubing may be mechanically strengthened even
further by necking 260 and annealing 270 processes. In the necking
260 treatment, the outer diameter thickness of the tubing is
reduced as the tubing is drawn through necking dies (not shown),
while the inner diameter remains constant. This necking 260 process
enhances the axial strength of the tubing by aligning the polymeric
molecules along the longitudinal axis. The necking 260 process
takes place at a temperature above T.sub.g, the polymer's glass
transition temperature, and below T.sub.m, the polymer's melting
temperature. For example, the necking temperature for poly L-lactic
acid is approximately 55.degree. C. to 60.degree. C. The necked
tube is then annealed 270 by blowing air onto the surface of the
necked tube once it comes out of the necking die. The necking 260
and annealing 270 may be repeated until the desired tube outer
diameter is achieved.
[0091] It will be appreciated that the area drawn down ratio in the
necking 260 step affects the strengthening effect. The area drawn
down ratio can be from about 1.01 to about 20.00, preferable 3.5 to
6.0. The area drawn down ratio is calculated as follows:
[0092] X is the diameter of the polymer-coated mandrel before
necking;
[0093] Y is the diameter of the mandrel;
[0094] Z is the diameter of the polymer-coated mandrel after
necking.
[0095] The area drawn down ratio
(R)=X.sup.2-Y.sup.2/Z.sup.2-Y.sup.2
[0096] Following necking 260, the tube is typically annealed 270
with pure inert gases. Suitable inert gases include, but are not
limited to, nitrogen, argon, neon, helium, or other noble gas.
Annealing 270 also increases the mechanical strength, by increasing
the crystallinity of the polymer and also regulates access of water
for hydrolysis.
[0097] Tube stents formed by this method had significantly enhanced
mechanical strength as compared to convention thermal extruded tube
stents. The maximum load at break for thermal extruded
poly(L-lactic acid) (PLLA) tube with a wall thickness 0.007" was
46.33.+-.1.66 NT. For tubes of same specifications that were made
by the present method had maximum load of 153.13.+-.1.66 NT.
[0098] In general, the strength of tube stents decreases as the
hydrolysis initiates. For example, PLLA becomes weaker as the
hydrolysis rate increases. In the compression extension comparison
test as described in Example 1, PLLA containing phosphate salt and
pure PLLA specimens were first immersed in water for 5 months and
tested for the radial strength. FIG. 12 shows the degraded
phosphate salt containing PLLA tube was 49.8% weaker than degraded
pure PLLA tube. FIG. 12 indicates that phosphate salt containing
PLLA tubes are subject to faster hydrolysis that led to the
accelerated loss of radial strength.
[0099] In an alternative embodiment, the method of making the
polymer tube stent with enhanced mechanical strength includes a
repeated dip-coating process 230, necking 260 the polymer tube at a
temperature above the glass transition temperature of the polymer
and below the melting temperature of the polymer, annealing 270 the
polymer tube and excimer laser cutting a stent of desired design
from the polymer tube, and finally cut 280 to form a
circumferential restraint when expanded.
[0100] As seen in FIG. 8, the polymer tube stent 300 of the instant
invention comprises a sinusoidal strut band 320 and a
circumferential restraint band 330. Adjacent bands are connected by
either a fixed axial link 340 or flexible axial link 350. As is
shown in FIG. 9, two adjacent radially expandable circumferential
restraint bands 330 are linked together at their respective valleys
by flexible axial link 350. The sinusoidal strut band 320 is linked
to the circumferential restraint band 330 at their respective
crowns via the fixed axial link 340. The sinusoidal strut band
comprises a substantially sinusoidal wave structure with at least
one peak and valley around the circumference of the stent.
Preferably, the sinusoidal strut band includes more than one peak
and valley around the circumference of the stent. During expansion
of the polymer tube stent 300 from the unexpanded to the expanded
state, the radially expandable circumferential restraint band 330
straightens out to compensate for the increase in radial diameter.
As is shown in FIG. 10, the radially expandable circumferential
restraint 330 is straightened out to form a complete hoop that will
contact the vessel lumen. When the radially expandable
circumferential restraint band 330 is straightened out, it locks
the polymer tube stent 300 in the expanded state. The polymer tube
stent 300 of the present invention may have as few as four peaks
per circumference, thus increasing the radius of the bend of each
band 330, 340 or as many as twelve peaks or more to accommodate
larger vessel lumens. The number of peaks on bands 330, 340 may be
adjusted to reduce mechanical stress and strain levels in the
bands, particularly during deployment.
[0101] In another embodiment, the stent comprises one or more
strength modules comprising one or more radially expandable tubular
elements. Preferably the radially expandable tubular elements
comprise a substantially sinusoidal wave structure of at least one
crown peak and crown valley. In one embodiment, the expandable
tubular elements have four or fewer crown peaks. The strength
modules are interconnected by one or more axial linking elements
that add flexibility to the strength module(s). The strength module
further has at least two circumferential restraint bands facing
opposite of a crown valley of the expandable tubular elements. In
an embodiment, the length of the circumferential restraint band
restraint band defines the size of the stent when the stent is
expanded. In this embodiment, the length of each circumferential
restraint band is less than a length of the expandable tubular
element.
[0102] An alternative embodiment, which comprises a biolock polymer
stent 400 made of elastic polymer is shown in FIG. 11. The polymer
stent 400 is made up of a tubular structure that is made up of one
or more radially expandable bands 405, 406 interconnected by long
fixed links 450, short fixed links 451, and spring connectors 452
so that the polymer stent 400 is radially expandable between an
unexpanded diameter and at least one expanded diameter, with a
locking mechanism made up of a first locking member such as an
axial peg 410 and a second locking member such as an axial receiver
420. The axial peg 410 protrudes from the valley 430 of band 406,
opposite of axial receiver 420 that is attached to the valley 440
of band 405. As the stent is expanded from its unexpanded state to
its expanded state, the axial peg 410 makes contact and engages
axial receiver 420, locking the stent in its expanded diameter. The
locking mechanism is not engaged when the polymer stent 400 is in
an unexpanded diameter and is engaged when the tubular structure is
in an expanded diameter. The spring connector 452 is disposed
between pairs of axial peg 410 and axial receiver 420 to increase
radial flexibility.
[0103] In some embodiments the polymer stent 400 is lockable at
multiple expanded diameters and in some embodiments the locking is
irreversible. In one embodiment, the axial peg includes teeth or
barbs that are dimensioned to fit within the axial receiver. The
axial receiver may further be shaped with one or more positions to
house the barbs of the axial peg and hold the axial peg in
position. In this manner, the stent can be expanded and locked at
one or more positions. It will be appreciated that the stent may
originally be locked into a first position and then further be
expanded to a second, or any number of, position(s). Thus, the
expansion of the stent can be increased and locked into position by
the locking of the axial peg and axial receiver. The polymer stent
400 has increased radial strength due to augmented force sharing.
The radial force is shared at the interface of the axial peg 410
and the axial receiver 420 in the locking mechanism, rather than on
the circumferential strength of the stent struts alone.
[0104] The tube stent may also include a mechanism for controlling
the resorption rate of the polymer tube and consequently, of the
stent. Buffer powders may be incorporated into the polymer solution
which is then used to form the tube stent. These buffers quickly
diffuse out of the tube stent, once it contacts fluid, thus
creating microscopic holes. Water molecules can then permeate the
tube stent through those holes. PLLA polymer decomposition is
hydrolysis-driven and subject to the influence of water content.
Resorption of the polymer occurs when the long molecular chain is
broken down into many single molecules forming lactic acid and then
nearby cells uptake the lactic acid. Thus, controlling the amount
of buffer powders loaded into the polymer solution, the buffer salt
diffusion rate and the tube stent resorption rate are
controllable.
[0105] The tube stent embodiment also optionally includes a process
for loading the polymer tubes with multiple APIs or a single API in
different concentrations in layered fashion at the time of tube
synthesis. Further, it may be desirable to provide different APIs
within the different layers of the stent. For example, one may
provide an immunosuppressant or anti-restenosis agent in the outer
or lateral layer or layers of the stent, an anti-inflammatory agent
in the middle layer or layers of the stent, and an
anti-thrombogenic agent in the medial (inner) layer or layers of
the stent. This is accomplished by changing the drug content in the
polymer as various layers of the polymer tube (which will become
the stent) are built up through dip coating. The timing and
duration of the release kinetics can be tuned by adjusting the
sequence of API, buffer, and polymer. Furthermore, as mentioned
above, the manufacturing process for the polymer tube stent is
carried out at moderate temperatures, which allows use of a much
wider range of APIs than is possible with thermally extruded
polymers.
[0106] Although the invention has been described with respect to
particular embodiments and applications, it will be appreciated
that various changes and modifications may be made without
departing from the invention. The following examples illustrate
various aspects of making and using the stent invention herein.
They are not intended to limit the scope of the invention.
III. EXAMPLES
[0107] Materials and Methods
[0108] Iohexol was purchased from Amersham (product
#0407-1414-80).
[0109] Methanol was purchased from EMD (Product #MX0488).
[0110] Phosphatidylcholine was purchased from Sigma-Aldrich (PN
P3556, 20 mg)
[0111] Radiograph images were taken with an OEC Model 9600 ESP
C-ARM 60 Hz with a magnification setting of MAG2.
Example 1
Preparation of Biodegradable Polymer Tubes
[0112] A biodegradable polymer tube was built layer by layer on a
mandrel by dipping the mandrel into a biodegradable polymer
solution of 12% wt % PLLA in CCl.sub.3H (chloroform). The mandrel
was dipped 46 times in the PLLA solution with a rate of dipping of
.about.0.1 meter per second.
[0113] The coated mandrel was then spin dried around the
longitudinal axis in a laminar flow hood, leaving a thin polymer
layer upon evaporation of solvent. The spin in the drying step was
repeated. The resulting a polymer tube had a thickness of
.about.0.2 mm and was 6% by weight phosphate salt buffer. This
polymer layer was then solvent polished with chloroform or the pure
solvent in which the polymer is dissolved and dried,
[0114] leaving behind a layer of thin and smooth polymer
tubing.
The outer diameter thickness of the tube stent was reduced by
drawing the tube stent through necking dies, while keeping the
inner diameter of the stent constant.
[0115] Following necking, the tubing was annealed with pure inert
nitrogen.
[0116] The average load at compression was measured before and
after 5 months of immersion in water as measured by a radial force
test performed on an Instron (Norwood, Mass.) force
delivery/measuring system. The average load at compression for
stents with and without buffer was tested with the results shown in
FIG. 12.
Example 2
Design and Fabrication of Stent
[0117] The stent pattern was designed using CAD software. Flat
layout designs and uncut biodegradable polymer tubing were sent to
a laser working studio for laser cutting. Several laser cutting
facilities are commercially available such as, Resonetics (Nashua,
N.H.) and Spectralytics (Dassel, Minn.). Stent designs were cut
from biodegradable polymer tube stents with an excimer laser with a
wavelength of less than 310 nm.
[0118] Those skilled in the art will appreciate that the inventive
stents, in the disclosed embodiments or variations thereof, provide
mechanical and therapeutic advantages over conventional stents. In
addition, advantageous treatments will suggest themselves to the
skilled practitioner considering the foregoing description of the
inventions. By virtue of the biodegradable polymeric nature of the
inventive stents, the same vessel site can be retreated at a later
time if needed, including staging procedures during growth of the
patient. Similarly, successive treatments of a tissue that is
changing size can be facilitated with the disclosed stents. It
should also be noted that the inventive stents can be implanted at
a site of healthy tissue for diagnostic purposes or therapeutic
treatment of adjacent tissue.
Example 3
Radiopaque Stent with Iodinated Contrast Agent
[0119] PLLA polymer stents 0.8-1.2 cm long with PLLA fiber diameter
of 0.01905 cm and fiber length of 15-22 cm were used. Iohexol was
dissolved in methanol to a concentration of 350 mg/mL. The pure
iohexol solution was then sprayed onto the top layer of the PLLA
stents to a coating thickness of about 0.01''. The measured dose on
all stent samples was 1000 .mu.g/stent. Once the methanol
evaporated, iohexol covered the abluminal stent surface completely.
The radiopacity of the coated stent was observed under the c-arm
after exposure to water for 30 seconds with the results shown in
FIG. 13 (#4). The radiopacity of a control stent formed of pure
PLLA was also tested (#1).
Example 4
[0120] Stent with BA9-PLLA Coating Solution on Top of Iodinated
Contrast Coating to Create Radiopacity
[0121] PLLA polymer stents 0.8-1.2 cm long with PLLA fiber diameter
of 0.01905 cm and fiber length of 15-22 cm were coated with
iohexol. The stent characteristics are shown in Table 1.
TABLE-US-00001 TABLE 1 Stent Characteristics Stent length (cm) 0.8
1.2 Total fiber length (cm) 15 22 Quantity 1 3
[0122] Iohexol was first dissolved in methanol to a concentration
of 350 mg/mL. The pure iohexol solution was then spray-coated onto
the PLLA stents. The measured dose on all stent samples was 1000
.mu.g/cm of stent. The BA9-PLLA coating solution was sprayed on top
of the iohexol coating to completely cover the abluminal surface.
The coated stents were then immersed in water for 30 seconds or two
minutes before observation.
TABLE-US-00002 TABLE 2 Coated stent Stent 1 2 3 4 Fiber length (cm)
16 16 22 23 Bare stent weight (mg) 2.457 2.287 2.504 2.338 2.455
2.29 2.506 2.339 2.456 2.289 2.506 2.338 Average 2.456 2.288667
2.505333 2.338333 Std. Deviation 0.001 0.001528 0.00115 0.000577
Final iohexol coating 3.396 3.486 3.545 3.601 weight (mg) 3.399
3.49 3.542 3.603 3.399 3.49 3.542 3.603 Average 0.942 1.2 1.037667
1.264 Std. Deviation 0.001732 0.002309 0.001732 0.001155 Estimated
iohexol 0.277222 0.343971 0.292878 0.350884 weight % Final BA9
coating 3.818 4.011 3.96 4.152 weight (mg) 3.817 4.013 3.96 4.15
3.817 4.013 3.96 4.15 Average 0.419333 0.523667 0.417 0.548333 Std.
Deviation 0.000577 0.001155 0 0.001155 Estimated BA9 0.209667
0.261833 0.2085 0.274167 weight (mg)
[0123] The radiopacity of the coated stents when exposed to water
for 30 seconds, two minutes and two control stent formed of pure
PLLA were tested and are shown in FIG. 14.
Example 5
Impregnating Stent with Iodinated Contrast Agent to create
Radiopacity
[0124] The PLLA backbone of the stent was impregnated with a
contrast agent. Iohoxel fine powder was suspended in
PLLA-chloroform solution with a final weight of 26 or 50 weight
percent of iohexol. In The radiopacity of the stents was observed
under the c-arm after exposure to water for 30 seconds with the
results shown in FIG. 13 (#2 and #3). The radiopacity of a control
stent formed of pure PLLA was also tested (#1). As seen in the
figure, radiopacity increased with an increase in iohexol weight
percentage.
[0125] Although preferred embodiments have been described and
illustrated, it should be understood that various changes,
substitutions and alterations can be made therein without departing
from the spirit and scope of the invention as defined by the
appended claims.
Example 6
Hydrophobic Iohexol Coating
[0126] A. Preparation of Phosphatidylcholine-iohexol liposome
[0127] Phosphatidylcholine (PC, available from Sigma-Aldrich,
product number P3556, 20 mg) is dissolved in 10 mL chloroform in a
50-ml round-bottom flask with a long extension neck, and the
chloroform is then removed under reduced pressure by a rotary
evaporator. The system is then purged with nitrogen and PC is
re-dissolved in the chloroform to form the solvent phase.
[0128] The aqueous phase (50 mg Iohexol in 1 mL distilled water) is
then added, the system is kept continuously under nitrogen, and the
resulting two-phase system is sonicated briefly (2-5 min) in a
bath-type sonicator (Bransonic Ultrasonic Cleaner, 1510R-MTH) until
the mixture becomes either a clear one-phase dispersion or a
homogeneous opalescent dispersion that does not separate for at
least 30 min after sonication. The sonication temperature is
20-25.degree. C. The mixture is then placed on the rotary
evaporator and chloroform is removed under reduced pressure (water
aspirator) at 20-25.degree. C., rotating at approximately 200
rpm.
[0129] During evaporation of chloroform, the system generally
froths. As the majority of the solvent is removed, the material
first forms a viscous gel and subsequently (within 5-10 min) it
becomes an aqueous suspension. At this point excess water can be
added (but this is not necessary) and the suspension evaporated for
an additional 15 min at 20.degree. C. to remove traces of solvent.
The preparation is then centrifuged to remove nonencapsulated
iohexol and residual chloroform. Finally, the PC-iohexol liposome
remains at 450.degree. C. for at least 30 min to completely remove
water. It is estimated 1.7-2.5 mg iohexol per mg PC.
[0130] B. Spray coating of PC-iohexol onto biodegradable stents
[0131] PC-iohexol liposome (10 mg) is suspended in (3 ml) ethylene
acetate and sonicated for 30 minutes. The solution is then
spray-coated onto stents. The spray coating process continues until
the net coating weight reaching 1.5 mg per stent. Then stents are
vacuum dried for 48 hours to remove ethylene acetate.
[0132] Reference: Reverse phase evaporation method. Henze et al,
Radio-opaque liposomes for the improved visualization of focal
liver disease by computerized tomography. Comput Med Imaging Graph.
1989 November-December;13(6):455-62.
* * * * *