U.S. patent application number 11/205254 was filed with the patent office on 2007-02-15 for fiber reinforced composite stents.
Invention is credited to Anthony J. Abbate, Joseph R. Callol, David C. Gale, Syed F. A. Hossainy, Bin Huang, Klaus Kleine, Timothy A. Limon, Srinivasan Sridharan.
Application Number | 20070038290 11/205254 |
Document ID | / |
Family ID | 37671250 |
Filed Date | 2007-02-15 |
United States Patent
Application |
20070038290 |
Kind Code |
A1 |
Huang; Bin ; et al. |
February 15, 2007 |
Fiber reinforced composite stents
Abstract
Polymeric composite stents reinforced with fibers for
implantation into a bodily lumen are disclosed.
Inventors: |
Huang; Bin; (Pleasanton,
CA) ; Gale; David C.; (San Jose, CA) ;
Sridharan; Srinivasan; (Morgan Hill, CA) ; Callol;
Joseph R.; (San Francisco, CA) ; Kleine; Klaus;
(Los Gatos, CA) ; Hossainy; Syed F. A.; (Fremont,
CA) ; Limon; Timothy A.; (Cupertino, CA) ;
Abbate; Anthony J.; (Santa Clara, CA) |
Correspondence
Address: |
SQUIRE, SANDERS & DEMPSEY LLP
1 MARITIME PLAZA
SUITE 300
SAN FRANCISCO
CA
94111
US
|
Family ID: |
37671250 |
Appl. No.: |
11/205254 |
Filed: |
August 15, 2005 |
Current U.S.
Class: |
623/1.16 ;
264/209.1; 623/1.22; 623/1.32; 623/1.34 |
Current CPC
Class: |
A61F 2002/91575
20130101; A61F 2/90 20130101; A61F 2230/0013 20130101; B29C 48/022
20190201; B29C 48/2886 20190201; D03D 3/02 20130101; A61F 2210/0076
20130101; B29C 48/10 20190201; B29C 45/0001 20130101; B29C 2043/028
20130101; B29C 48/0022 20190201; B29K 2101/00 20130101; B29C 43/02
20130101; B23K 26/402 20130101; B29C 59/021 20130101; B29C 45/0055
20130101; A61F 2210/0004 20130101; A61F 2220/005 20130101; A61F
2/04 20130101; B29C 59/02 20130101; B23K 26/50 20151001; B29K
2995/0056 20130101; A61F 2/91 20130101; A61L 31/129 20130101; B29C
48/18 20190201; B29L 2023/003 20130101; A61F 2/915 20130101; A61F
2002/91508 20130101; B23K 2103/50 20180801; B23K 2103/42 20180801;
B29C 53/566 20130101; B29C 45/0005 20130101; B29K 2995/006
20130101; A61F 2/06 20130101; B29K 2267/043 20130101; B29K 2105/12
20130101; A61F 2002/072 20130101; A61F 2250/0098 20130101; B29C
48/09 20190201; A61F 2210/0009 20130101; B29C 53/583 20130101; B29L
2031/7546 20130101 |
Class at
Publication: |
623/001.16 ;
623/001.22; 623/001.32; 623/001.34; 264/209.1 |
International
Class: |
A61F 2/88 20060101
A61F002/88; A61F 2/90 20060101 A61F002/90; B29C 47/20 20070101
B29C047/20 |
Claims
1. A method of making a stent comprising: forming a mixture
comprising a matrix polymer and a plurality of short fibers, the
fibers comprising a material having a melting temperature greater
than a melting temperature of the matrix polymer; disposing the
mixture in a tube or sheet forming apparatus to form a tube or a
sheet, wherein the apparatus is heated so that a temperature of the
mixture in the apparatus is greater than the melting temperature of
the matrix polymer and less than the melting temperature of the
material of the fibers such that at least a portion of the matrix
polymer is a polymer melt; and fabricating a stent from the tube or
sheet comprising the matrix polymer and the short fibers.
2. The method of claim 1, wherein the matrix polymer comprises a
biodegradable polymer.
3. The method of claim 1, wherein the material of the fibers
comprises a biodegradable polymer.
4. The method of claim 1, wherein the material of the fibers
comprises a biostable and/or erodible metal.
5. The method of claim 1, wherein the material of the fibers
comprises a radiopaque material.
6. The method of claim 1, wherein the mixture comprises an active
agent.
7. The method of claim 1, wherein at least some of the fibers
comprise an active agent.
8. The method of claim 1, wherein a length of at least a portion of
the short fibers is substantially smaller than a length of the
formed tube.
9. The method of claim 1, wherein the mixture is formed in a mixing
apparatus at a temperature greater than the melting temperature of
the matrix polymer and less than the melting temperature of the
material of the fibers.
10. The method of claim 1, wherein the tube comprises forming a
tube or a sheet by injection molding.
11. The method of claim 1, wherein forming the tube comprises
extruding a tube or a sheet from the mixture.
12. The method of claim 1, wherein the formed tube or sheet is
cooled to a temperature below the melting temperature of the matrix
polymer such that a majority of the matrix polymer in the formed
tube is amorphous, crystalline, or partially crystalline.
13. The method of claim 1, wherein fabricating a stent comprises
forming a pattern on the formed tube comprising a plurality of
interconnecting structural elements.
14. The method of claim 1, wherein fabricating a stent comprises
forming a tube from the sheet and forming a pattern on the formed
sheet comprising a plurality of interconnecting structural
elements.
15. The method of claim 1, further comprising radially deforming
the tube comprising the matrix polymer and the short fibers to
increase circumferential strength and rigidity of the tube and/or
to increase the circumferential alignment of at least some of the
fibers.
16. A stent made according to the method of claim 1.
17. A method of making a stent comprising: forming a tube
comprising at least one fiber layer and at least one polymer film
layer, fibers of at least one fiber layer comprising a material
having a melting temperature greater than a melting temperature of
at least one polymer film layer; heating the tube to a temperature
greater than the melting temperature of at least one polymer film
layer and less than the melting temperature of the material of the
fibers to melt at least a portion of the polymer of at least one
polymer film layer, at least a portion of at least one fiber layer
becoming embedded within at least a portion of the melted polymer
of at least one polymer film layer; cooling the heated tube; and
fabricating a stent from the cooled tube.
18. The method of claim 17, wherein at least one polymer film layer
comprises a biodegradable polymer.
19. The method of claim 17, wherein the material of the fibers of
at least one fiber layer comprises a biodegradable polymer.
20. The method of claim 17, wherein the material of the fibers of
at least one fiber layer comprises a biostable and/or erodible
metal.
21. The method of claim 17, wherein the material of the fibers of
at least one fiber layer comprises a radiopaque material.
22. The method of claim 17, wherein at least one fiber layer
alternates with at least one film layer.
23. The method of claim 17, wherein at least one fiber layer
comprises a woven structure.
24. The method of claim 17, wherein at least one of the polymer
film layers comprises an active agent.
25. The method of claim 17, wherein at least some of the fibers
comprise an active agent.
26. The method of claim 17, wherein forming the tube comprises
disposing at least one layer on a mandrel.
27. The method of claim 17, wherein the heated tube is cooled to a
temperature below the melting temperature of at least one polymer
film layer such that a majority of the polymer that was melted
becomes amorphous, crystalline, or partially crystalline.
28. The method of claim 17, wherein fabricating a stent comprises
forming a pattern on the tube comprising a plurality of
interconnecting structural elements.
29. The method of claim 17, further comprising radially deforming
the heated tube to increase circumferential strength and rigidity
of the tube.
30. The method of claim 17, wherein an orientation of fibers
relative to a cylindrical axis of the tube of at least one fiber
layer may be different from an orientation of fibers in another
fiber layer.
31. The method of claim 17, wherein an orientation of fibers
relative to a cylindrical axis of the tube in at least one fiber
layer is greater than 90.degree. and an orientation of fibers in
another fiber layer is less than 90.degree..
32. A stent made according to the method of claim 17.
33. A method of making a stent comprising: forming a layered sheet
comprising at least one fiber layer and at least one polymer film
layer, fibers of at least one fiber layer comprising a material
having a melting temperature greater than a melting temperature of
at least one polymer film layer; heating the layered sheet to a
temperature greater than the melting temperature of at least one
polymer film layer and less than the melting temperature of the
material of the fibers to melt at least a portion of the polymer of
at least one polymer film layer, at least a portion of the fibers
becoming embedded within at least a portion of the melted polymer
of at least one polymer film layer; cooling the heated layered
sheet; and fabricating a stent from the cooled sheet.
34. The method of claim 33, wherein at least one polymer film layer
comprises a biodegradable polymer.
35. The method of claim 33, wherein the material of the fibers of
at least one fiber layer comprises a biodegradable polymer.
36. The method of claim 33, wherein the material of the fibers of
at least one fiber layer comprises a biostable and/or erodible
metal.
37. The method of claim 33, wherein the material of the fibers of
at least one fiber layer comprises a radiopaque material.
38. The method of claim 33, wherein at least one fiber layer
alternates with at least one polymer film layer.
39. The method of claim 33, wherein fabricating the stent comprises
forming a tube from the layered sheet and forming a pattern on the
tube and/or the layered sheet comprising a plurality of
interconnecting structural elements.
40. The method of claim 33, wherein at least one of the polymer
film layers comprises an active agent.
41. The method of claim 33, wherein at least some of the fibers
comprise an active agent.
42. The method of claim 33, further comprising cooling the heated
layered sheet.
43. The method of claim 33, wherein the fibers are nanofibers.
44. The method of claim 33, wherein the heated sheet is cooled to a
temperature below the melting temperature of the polymer films such
that a majority of the polymer that was melted becomes amorphous,
crystalline and/or partially crystalline.
45. The method of claim 33, wherein fabricating a stent comprises
forming a pattern on the sheet comprising a plurality of
interconnecting structural elements and forming a tube from the
sheet.
46. The method of claim 33, wherein fabricating a stent comprises
forming a tube from the sheet and forming a pattern on the sheet
comprising a plurality of interconnecting structural elements.
47. A stent made according to the method of claim 33.
48. A method of making a stent comprising: forming a coating layer
comprising a coating polymer over a tube-shaped fiber layer
comprising a plurality of fibers, wherein the coating layer is
formed by applying a fluid comprising the coating polymer dissolved
in a solvent and by removing all or a majority of the solvent from
the applied fluid, the fibers comprising a material being insoluble
or having a relatively low solubility in the solvent, the material
comprising a melting temperature greater than a melting temperature
of the coating polymer; and fabricating a stent from the coated
fiber layer.
49. The method of claim 48, wherein the coating polymer comprises a
biodegradable polymer.
50. The method of claim 48, wherein the material of the fibers of
at least one fiber layer comprises a biodegradable polymer.
51. The method of claim 48, wherein the material of the fibers
comprises a biostable and/or erodible metal.
52. The method of claim 48, wherein the material of the fibers
comprises a radiopaque material.
53. The method of claim 48, wherein the coating layer comprises an
active agent.
54. The method of claim 48, wherein at least some of the fibers
comprise an active agent.
55. The method of claim 48, wherein the tube-shaped fibers layer
comprises a woven structure.
56. The method of claim 48, wherein applying the fluid comprises
spraying the fluid on and/or dipping the tube into the solvent.
57. The method of claim 48, further comprising heating the coating
layer to a temperature above the melting temperature of the coating
polymer and cooling the coating layer to a temperature below the
melting temperature of the coating polymer so that a majority of
the coating layer is amorphous, crystalline, and/or partially
crystalline.
58. The method of claim 48, further comprising radially deforming
the coated tube to increase circumferential strength and rigidity
of the tube.
59. The method of claim 48, wherein the fiber layer is disposed on
a mandrel.
60. The method of claim 48, wherein the fiber layer is disposed on
a mandrel over a coating layer comprising the coating polymer.
61. A stent made according to the method of claim 48.
62. A method of making a stent comprising: disposing a plurality of
fibers within a mold for forming a structure; disposing a matrix
polymer that is partially or completely molten into the mold to at
least partially embed the fibers within the molten polymer, the
fiber comprising a material having a melting temperature greater
than a melting temperature of the matrix polymer, wherein a
temperature of the matrix polymer and the fibers in the mold is
less than the melting temperature of the fiber material; cooling
the molten polymer to form the structure; and fabricating a stent
from the cooled structure.
63. The method of claim 62, wherein the matrix polymer comprises a
biodegradable polymer.
64. The method of claim 62, wherein the material of the fibers
comprises a biodegradable polymer.
65. The method of claim 62, wherein the material of the fibers
comprises a biostable and/or erodible metal.
66. The method of claim 62, wherein the material of the fibers
comprises a radiopaque material.
67. The method of claim 62, wherein the material of the fibers
comprises a biodegradable polymer.
68. The method of claim 62, wherein the material of the fibers
comprises a biostable and/or erodible metal.
69. The method of claim 62, wherein the material of the fibers
comprises a radiopaque material.
70. The method of claim 62, further comprising disposing an active
agent into the mold.
71. The method of claim 62, wherein at least some of the fibers
comprise an active agent.
72. The method of claim 62, wherein disposing the fibers within a
mold comprises disposing the fibers around a mandrel disposed
within the mold.
73. The method of claim 62, wherein disposing the fibers within a
mold in a random or substantially random fashion.
74. The method of claim 62, wherein the structure comprises a tube
or a sheet.
75. The method of claim 62, wherein the structure comprises a
sheet, and wherein fabricating a stent comprises forming a tube
from the sheet and forming a pattern in the tube comprising a
plurality of interconnecting structural elements.
76. The method of claim 62, wherein the structure comprises a tube,
and wherein fabricating a stent comprises forming a pattern on the
tube comprising a plurality of interconnecting structural
elements.
77. The method of claim 62, wherein the structure is cooled to a
temperature below the melting temperature of the matrix polymer
such that a majority of the matrix polymer in the formed structure
is amorphous, crystalline, and/or partially crystalline.
78. The method of claim 62, wherein the structure comprises a tube,
and further comprising radially deforming the formed tube to
increase circumferential strength and rigidity of the tube.
79. A stent made according to the method of claim 68.
80. A method of making a stent comprising: disposing a plurality of
fibers in an extruder for forming a structure; conveying a matrix
polymer into the extruder, the fibers comprising a material having
a melting temperature greater than a melting temperature of the
matrix polymer; forming the structure with the extruder at a
temperature greater than the melting temperature of the matrix
polymer and less than the melting temperature of the material,
wherein at least some of the fibers becoming embedded within the
matrix polymer; and fabricating a stent from the cooled
structure.
81. The method of claim 80, wherein the matrix polymer comprises a
biodegradable polymer.
82. The method of claim 80, wherein the material of the fibers
comprises a biodegradable polymer.
83. The method of claim 80, wherein the material of the fibers
comprises a biostable and/or erodible metal.
84. The method of claim 80, wherein the material of the fibers
comprises a radiopaque material.
85. The method of claim 80, further comprising conveying an active
agent into the extruder.
86. The method of claim 80, wherein at least some of the fibers
comprise an active agent.
87. The method of claim 80, wherein disposing the plurality of
fibers in the extruder comprises disposing the fibers around a
mandrel.
88. The method of claim 80, wherein disposing the plurality of
fibers in the extruder comprises disposing the fibers in the
extruder in a random or substantially random fashion.
89. The method of claim 80, wherein the structure comprises a
sheet, and wherein fabricating a stent comprises forming a tube
from the sheet and forming a pattern in the tube comprising a
plurality of interconnecting structural elements.
90. The method of claim 80, wherein the structure comprises a tube,
and wherein fabricating a stent comprises forming a pattern in the
tube comprising a plurality of interconnecting structural
elements.
91. The method of claim 80, wherein the structure is cooled to a
temperature below the melting temperature of the matrix polymer
such that a majority of the matrix polymer in the formed structure
is amorphous, crystalline, or partially crystalline.
92. The method of claim 80, wherein the structure comprises a tube,
and further comprising radially deforming the formed tube to
increase circumferential strength and rigidity of the tube.
93. A stent made according to the method of claim 80.
94. A method of making a stent comprising: heating a fiber mesh
tube comprising two types of fibers, a first fiber comprising a
first polymer and the second fiber comprising a second polymer, the
first polymer having a softening temperature lower than a softening
temperature of the second polymer, wherein the tube is heated to a
temperature range between the softening temperature of the first
polymer and the softening temperature of the second polymer; and
applying pressure to the tube so as to flatten at least some of the
fibers of the tube to reduce a radial profile of the tube.
95. The method of claim 94, wherein the first polymer comprises a
biostable and/or biodegradable polymer.
96. The method of claim 94, wherein the second polymer comprises a
biostable and/or biodegradable polymer.
97. The method of claim 94, wherein the tube is disposed over a
mandrel during heating.
98. The method of claim 94, wherein pressure is applied with a
heated crimper.
99. The method of claim 94, wherein the diameter of the tube is
fixed during heating.
100. The method of claim 94, wherein the tube is heated and
pressure applied at or near a fabricated diameter of the tube.
101. The method of claim 94, wherein the temperature of the tube is
maintained in the temperature range for a selected period of time
to allow heat setting of the tube.
102. The method of claim 94, further comprising radially expanding
the tube prior to, during, or subsequent to heating and/or applying
pressure to flatten at least some of the fibers.
103. The method of claim 94, wherein the first polymer has a lower
melting temperature than the second polymer.
104. The method of claim 94, wherein the temperature range is below
the melting temperature of the first polymer and the second
polymer.
105. The method of claim 94, wherein the temperature range is below
the glass transition temperature the second polymer.
106. The method of claim 94, wherein the applied pressure reduces a
radial profile of at least some of the net points of the
fibers.
107. A stent made according to the method of claim 94.
108. A method of making a stent comprising: heating a fiber mesh
tube, at least some of the fibers of the tube comprising a first
polymer and a second polymer, the first polymer having a softening
temperature lower than a softening temperature of the second
polymer, wherein the tube is heated to a temperature range between
the softening temperature of the first polymer and the softening
temperature of the second polymer; and applying pressure to the
tube so as to flatten at least some of the fibers of the tube to
reduce a radial profile of the tube.
109. The method of claim 108, wherein at least some of the fibers
of the tube comprise an inner core and an outer covering, the inner
core comprising the first polymer and the outer covering comprising
the second polymer.
110. The method of claim 108, wherein at least some of the fibers
of the tube comprise a mixture of the first polymer and the second
polymer.
111. The method of claim 108, wherein the first polymer comprises a
biostable and/or biodegradable polymer.
112. The method of claim 108, wherein the second polymer comprises
a biostable and/or biodegradable polymer.
113. The method of claim 108, wherein the tube is disposed over a
mandrel during heating.
114. The method of claim 108, wherein pressure is applied with a
heated crimper.
115. The method of claim 108, wherein the diameter of the tube is
fixed during heating.
116. The method of claim 108, wherein the tube is heated and
pressure applied at or near a fabricated diameter of the tube.
117. The method of claim 108, wherein the temperature of the tube
is maintained in the temperature range for a selected period of
time to allow heat setting of the tube.
118. The method of claim 108, further comprising radially expanding
the tube prior to, during, or subsequent to heating and/or applying
pressure to flatten at least some of the fibers.
119. The method of claim 108, further comprising crimping the tube
prior to, during, or subsequent to heating the tube and/or applying
pressure to flatten at least some of the fibers.
120. The method of claim 108, wherein the first polymer has a lower
melting temperature than the second polymer.
121. The method of claim 108, wherein the temperature range is
below the melting temperature of the first polymer and the second
polymer.
122. The method of claim 108, wherein the temperature range is
below the glass transition temperature of the second polymer.
123. The method of claim 108, wherein the applied pressure reduces
a radial profile of at least some of the net points of the
fibers.
124. A stent made according to the method of claim 108.
125. A method of making a stent comprising: coupling a metallic
film to at least a portion of a surface of a polymeric tube; and
fabricating a stent from the tube with the metallic film so that
the metallic film is over at least a portion of a surface of the
stent.
126. The method of claim 125, wherein the metallic film comprises a
biostable and/or erodible metal.
127. The method of claim 125, further comprising forming a coating
above at least a portion of the metallic film above a portion of
the surface of the stent.
128. The method of claim 125, wherein the metallic film is coupled
to at least a portion of a surface of a polymeric tube with a
biocompatible adhesive.
129. The method of claim 125, wherein the polymeric tube comprises
a biostable and/or biodegradable polymer.
130. The method of claim 125, wherein the metallic film comprises a
band circumferentially aligned around a surface of the tube.
131. The method of claim 125, wherein the metallic film comprises a
longitudinal strip longitudinally aligned along the surface of the
tube.
132. The method of claim 125, wherein fabricating the stent
comprises forming a pattern in the tube having the metallic film,
the pattern comprising a plurality of interconnecting structural
elements.
133. A stent made according to the method of claim 125.
134. A method of making a stent comprising: forming a tube having a
metallic film in between two radial polymeric layers; and
fabricating a stent from the tube.
135. The method of claim 134, wherein the tube is formed by
extruding an outer polymeric tubular layer over an inner polymeric
tubular layer, the inner layer comprising a metallic film disposed
above a surface of the inner layer.
136. The method of claim 134, wherein the polymeric layers comprise
a biostable and/or biodegradable polymer.
137. The method of claim 134, wherein the metallic film comprises a
biostable and/or erodible metal.
138. The method of claim 134, wherein the metallic film comprises a
band circumferentially aligned around a circumference of the tube
in between the layers.
139. The method of claim 134, wherein the metallic film comprises a
longitudinal strip longitudinally aligned along the tube in between
the polymer layers.
140. The method of claim 134, wherein fabricating the stent
comprises forming a pattern in the tube, the pattern comprising a
plurality of interconnecting structural elements.
141. A stent made according to the method of claim 134.
142. A method of making a stent comprising: elongating a polymeric
tube so that a diameter of the stent decreases; positioning a
metallic band around a circumference of the elongated tube; heating
the elongated polymeric tube with the metallic band positioned
around the tube; allowing the heated tube to radially expand so as
to couple the metallic band to the tube; and fabricating a stent
from the expanded tube.
143. The method of claim 142, wherein the metallic band comprises a
biostable and/or erodible metal.
144. The method of claim 142, wherein the heated tube radially
expands to at least a diameter of the metallic band.
145. The method of claim 142, wherein the polymeric tube comprises
a biostable and/or biodegradable polymer.
146. The method of claim 142, wherein fabricating the stent
comprises forming a pattern in the expanded tube, the pattern
comprising a plurality of interconnecting structural elements.
147. A stent made according to the method of claim 142.
148. A radially expandable stent comprising a plurality of
interconnecting structural elements, the structural elements
comprising fibers at least partially embedded in a matrix polymer,
the fibers including a material having a melting temperature
greater than a melting temperature of the matrix polymer, the
fibers configured to provide mechanical reinforcement to the stent
due to a higher strength and modulus along an axis of the fibers
than the matrix polymer.
149. A radially expandable stent comprising a plurality of
interconnecting structural elements, the structural elements
comprising at least one radial fiber layer and at least one radial
polymer film layer, the fibers including material with a melting
temperature greater than a melting temperature than at least one
polymer film layer, at least one fiber layer being at least
partially embedded within at least one polymer film layer, the
fibers configured to provide mechanical reinforcement to the stent
due to a higher strength and modulus along an axis of the fibers
than the polymer film layer.
150. A radially expandable stent comprising a plurality of
structural elements, the structural elements comprising at least
two radial fiber layers and at least one radial polymer film layer,
at least a portion of at least one fiber layer being embedded
within at least a portion of at least one polymer film layer,
wherein an orientation of fibers relative to a cylindrical axis of
the stent of at least one fiber layer is different from an
orientation of fibers in another fiber layer.
151. The stent of claim 150, wherein the orientation of fibers
relative to the cylindrical axis of the stent in at least one fiber
layer is greater than 90.degree. and an orientation of fibers in
another fiber layer is less than 90.degree..
152. A radially expandable stent woven from at lease two types of
fibers, a first fiber comprising a first polymer and the second
fiber comprising a second polymer, the first polymer having a
softening temperature lower than a softening temperature of the
second polymer, wherein at least some of the fibers have a
flattened radial profile that reduces the radial profile of the
tube.
153. A radially expandable stent woven from fibers comprising a
first polymer and a second polymer, the first polymer having a
softening temperature lower than a softening temperature of the
second polymer, wherein at least some of the fibers have a
flattened radial profile that reduces the radial profile of the
tube.
154. A radially expandable stent comprising metallic film coupled
to a plurality of portions of a surface of the stent, wherein the
metallic film is sufficiently radiopaque to allow the stent to be
visualized during use.
155. The device of claim 153, wherein the metallic film comprises a
biostable and/or erodible metal.
156. The device of claim 153, further comprising a coating above at
least some of the plurality of portions of the surface having the
metallic film.
157. The device of claim 153, wherein the polymeric tube comprises
a biostable and/or biodegradable polymer.
158. The device of claim 153, wherein the plurality of portions are
circumferentially aligned.
159. The device of claim 153, wherein the plurality of portions are
longitudinally aligned.
160. A radially expandable stent comprising a plurality of
interconnecting structural elements, the structural elements having
two radial polymeric layers with metallic film embedded in a
plurality of locations in between the layers, and wherein the
metallic film is sufficiently radiopaque to allow the stent to be
visualized during use.
161. The device of claim 160, wherein the polymer layers comprise a
biostable and/or biodegradable polymer.
162. The device of claim 160, wherein the metallic film comprises a
biostable and/or erodible metal.
163. The device of claim 160, wherein the plurality of portions are
circumferentially aligned.
164. The device of claim 160, wherein the plurality of portions are
longitudinally aligned.
Description
BACKGROUND OF THE INVENTION
[0001] 1. Field of the Invention
[0002] This invention relates to radially expandable implantable
medical devices such as stents for implantation into a bodily
lumen. In particular, the invention relates composite stents
reinforced with fibers.
[0003] 2. Description of the State of the Art
[0004] This invention relates to radially expandable
endoprostheses, which are adapted to be implanted in a bodily
lumen. An "endoprosthesis" corresponds to an artificial device that
is placed inside the body. A "lumen" refers to a cavity of a
tubular organ such as a blood vessel.
[0005] A stent is an example of such an endoprosthesis. Stents are
generally cylindrically shaped devices, which function to hold open
and sometimes expand a segment of a blood vessel or other
anatomical lumen such as urinary tracts and bile ducts. Stents are
often used in the treatment of atherosclerotic stenosis in blood
vessels. "Stenosis" refers to a narrowing or constriction of the
diameter of a bodily passage or orifice. In such treatments, stents
reinforce body vessels and prevent restenosis following angioplasty
in the vascular system. "Restenosis" refers to the reoccurrence of
stenosis in a blood vessel or heart valve after it has been treated
(as by balloon angioplasty, stenting, or valvuloplasty) with
apparent success.
[0006] The treatment of a diseased site or lesion with a stent
involves both delivery and deployment of the stent. "Delivery"
refers to introducing and transporting the stent through a bodily
lumen to a region, such as a lesion, in a vessel that requires
treatment. "Deployment" corresponds to the expanding of the stent
within the lumen at the treatment region. Delivery and deployment
of a stent are accomplished by positioning the stent about one end
of a catheter, inserting the end of the catheter through the skin
into a bodily lumen, advancing the catheter in the bodily lumen to
a desired treatment location, expanding the stent at the treatment
location, and removing the catheter from the lumen.
[0007] In the case of a balloon expandable stent, the stent is
mounted about a balloon disposed on the catheter. Mounting the
stent typically involves compressing or crimping the stent onto the
balloon. The stent is then expanded by inflating the balloon. The
balloon may then be deflated and the catheter withdrawn. In the
case of a self-expanding stent, the stent may be secured to the
catheter via a retractable sheath or a sock. When the stent is in a
desired bodily location, the sheath may be withdrawn which allows
the stent to self-expand.
[0008] The stent must be able to satisfy a number of mechanical
requirements. First, the stent must be capable of withstanding the
structural loads, namely radial compressive forces, imposed on the
stent as it supports the walls of a vessel. Therefore, a stent must
possess adequate radial strength. Radial strength, which is the
ability of a stent to resist radial compressive forces, is due to
strength and rigidity around a circumferential direction of the
stent. Radial strength and rigidity, therefore, may also be
described as, hoop or circumferential strength and rigidity.
[0009] Once expanded, the stent must adequately maintain its size
and shape throughout its service life despite the various forces
that may come to bear on it, including the cyclic loading induced
by the beating heart. For example, a radially directed force may
tend to cause a stent to recoil inward. Generally, it is desirable
to minimize recoil.
[0010] In addition, the stent must possess sufficient flexibility
to allow for crimping, expansion, and cyclic loading. Longitudinal
flexibility is important to allow the stent to be maneuvered
through a tortuous vascular path and to enable it to conform to a
deployment site that may not be linear or may be subject to
flexure. Finally, the stent must be biocompatible so as not to
trigger any adverse vascular responses.
[0011] The structure of a stent is typically composed of
scaffolding that includes a pattern or network of interconnecting
structural elements often referred to in the art as struts or bar
arms. The scaffolding can be formed from wires, tubes, or sheets of
material rolled into a cylindrical shape. Conventional methods of
constructing a stent from a polymer material involve extrusion,
blow molding, or injection molding a polymer tube based on a single
polymer or polymer blend and then laser cutting a pattern into the
tube. The scaffolding is designed so that the stent can be radially
compressed (to allow crimping) and radially expanded (to allow
deployment). A conventional stent is allowed to expand and contract
through movement of individual structural elements of a pattern
with respect to each other.
[0012] Additionally, a medicated stent may be fabricated by coating
the surface of either a metallic or polymeric scaffolding with a
polymeric carrier that includes an active or bioactive agent or
drug. Polymeric scaffolding may also serve as a carrier of an
active agent or drug.
[0013] Furthermore, it may be desirable for a stent to be
biodegradable. In many treatment applications, the presence of a
stent in a body may be necessary for a limited period of time until
its intended function of, for example, maintaining vascular patency
and/or drug delivery is accomplished. Therefore, stents fabricated
from biodegradable, bioabsorbable, and/or bioerodable materials
such as bioabsorbable polymers should be configured to completely
erode only after the clinical need for them has ended.
[0014] In general, there are several important aspects in the
mechanical behavior of polymers that affect stent design. Polymers
tend to have lower strength than metals on a per unit mass basis.
Therefore, polymeric stents typically have less circumferential
strength and radial rigidity than metallic stents of the same or
similar dimensions. Inadequate radial strength potentially
contributes to a relatively high incidence of recoil of polymeric
stents after implantation into vessels.
[0015] Another potential problem with polymeric stents is that
their struts or bar arms can crack during crimping and expansion,
especially for brittle polymers. The localized portions of the
stent pattern subjected to substantial deformation tend to be the
most vulnerable to failure. Furthermore, in order to have adequate
mechanical strength, polymeric stents may require significantly
thicker struts than a metallic stent, which results in an
undesirably larger profile.
[0016] Additionally, another factor to consider in stent design is
radiopacity. In addition to meeting the mechanical requirements
described above, it is desirable for a stent to be radiopaque, or
fluoroscopically visible under x-rays. "Radiopaque" refers to the
ability of a substance to absorb x-rays. Accurate stent placement
is facilitated by real time visualization of the delivery of a
stent. A cardiologist or interventional radiologist can track the
delivery catheter through the patient's vasculature and precisely
place the stent at the site of a lesion. This is typically
accomplished by fluoroscopy or similar x-ray visualization
procedures. For a stent to be fluoroscopically visible it must be
more absorptive of x-rays than the surrounding tissue. Radiopaque
materials in a stent may allow for its direct visualization.
[0017] A significant shortcoming of polymers as compared to metals
(and polymers generally composed of carbon, hydrogen, oxygen, and
nitrogen) is that they are radiolucent with no radiopacity.
Polymers tend to have x-ray absorption similar to body tissue.
[0018] Additionally, there are manufacturing difficulties in
placing small markers on stents as well as challenges in keeping
very small markers attached to the stent. If the maximum
permissible size of the marker is too small to be visible on a
fluoroscope, multiple markers may be necessary. This makes
manufacturing even more challenging.
[0019] Therefore, it would be desirable to have methods of making
biodegradable polymeric stents that are both strong and
flexible.
SUMMARY OF THE INVENTION
[0020] Certain embodiments of the present invention are directed to
a method of making a stent that may include forming a mixture
having a matrix polymer and a plurality of short fibers such that
fibers include a material having a melting temperature greater than
a melting temperature of the matrix polymer. The method may further
include disposing the mixture in a tube or sheet forming apparatus
to form a tube or a sheet such that the apparatus is heated so that
a temperature of the mixture in the apparatus is greater than the
melting temperature of the matrix polymer and less than the melting
temperature of the material of the fibers. At least a portion of
the matrix polymer may be a polymer melt. A stent may be fabricated
from the tube or sheet including the matrix polymer and the short
fibers.
[0021] Further embodiments of the present invention are directed to
a method of making a stent that may include forming a tube having
at least one fiber layer and at least one polymer film layer such
that fibers of at least one fiber layer include a material having a
melting temperature greater than a melting temperature of at least
one polymer film layer. The method may further include heating the
tube to a temperature greater than the melting temperature of at
least one polymer film layer and less than the melting temperature
of the material of the fibers to melt at least a portion of the
polymer of at least one polymer film layer. At least a portion of
at least one fiber layer may become embedded within at least a
portion of the melted polymer of at least one polymer film layer.
The heated tube may then be cooled and a stent fabricated from the
cooled tube.
[0022] Additional embodiments of the present invention are directed
to a method of making a stent that may include forming a layered
sheet having at least one fiber layer and at least one polymer film
layer such that fibers of at least one fiber layer include a
material having a melting temperature greater than a melting
temperature of at least one polymer film layer. The method may
further include heating the layered sheet to a temperature greater
than the melting temperature of at least one polymer film layer and
less than the melting temperature of the material of the fibers to
melt at least a portion of the polymer of at least one polymer film
layer. At least a portion of the fibers may become embedded within
at least a portion of the melted polymer of at least one polymer
film layer. The heated layered sheet may then be cooled and a stent
fabricated from the cooled sheet.
[0023] Additional embodiments of the present invention are directed
to a method of making a stent that may include forming a coating
layer comprising a coating polymer over a tube-shaped fiber layer
having a plurality of fibers. The coating layer may be formed by
applying a fluid including the coating polymer dissolved in a
solvent and by removing all or a majority of the solvent from the
applied fluid. The fibers may include a material that is insoluble
or having a relatively low solubility in the solvent. The material
may have a melting temperature greater than a melting temperature
of the coating polymer. The method may further include fabricating
a stent from the coated fiber layer.
[0024] Other embodiments of the present invention are directed to a
method of making a stent that may include disposing a plurality of
fibers within a mold for forming a structure. The method may
further include disposing a matrix polymer that is partially or
completely molten into the mold to at least partially embed the
fibers within the molten polymer. The fiber may include a material
having a melting temperature greater than a melting temperature of
the matrix polymer. A temperature of the matrix polymer and the
fibers in the mold may be less than the melting temperature of the
material. The method may further include cooling the molten polymer
to form the structure and fabricating a stent from the cooled
structure.
[0025] Further embodiments of the present invention are directed to
a method of making a stent that may include disposing a plurality
of fibers in an extruder for forming a structure. The method may
further include conveying a matrix polymer into the extruder. The
fibers may include a material having a melting temperature greater
than a melting temperature of the matrix polymer. The structure may
be formed with the extruder at a temperature greater than the
melting temperature of the matrix polymer and less than the melting
temperature of the material in such a way that at least some of the
fibers become embedded within the matrix polymer. A stent may then
be fabricated from the cooled structure.
[0026] Additional embodiments of the present invention are directed
to a method of making a stent that may include heating a fiber mesh
tube including two types of fibers. A first fiber may include a
first polymer and the second fiber may include a second polymer.
The first polymer may have a softening temperature lower than a
softening temperature of the second polymer. The tube may be heated
to a temperature range between the softening temperature of the
first polymer and the softening temperature of the second polymer.
The method may further include applying pressure to the tube so as
to flatten at least some of the fibers of the tube to reduce a
radial profile of the tube.
[0027] Some further embodiments of the present invention are
directed to a method of making a stent that may include heating a
fiber mesh tube. At least some of the fibers of the tube may
include a first polymer and a second polymer. The first polymer may
have a softening temperature lower than a softening temperature of
the second polymer. The tube may be heated to a temperature range
between the softening temperature of the first polymer and the
softening temperature of the second polymer. The method may further
include applying pressure to the tube so as to flatten at least
some of the fibers of the tube to reduce a radial profile of the
tube.
[0028] Some further embodiments of the present invention are
directed to a method of making a stent that may include coupling a
metallic film to at least a portion of a surface of a polymeric
tube. The method may further include fabricating a stent from the
tube with the metallic film so that the metallic film is over at
least a portion of a surface of the stent.
[0029] Other embodiments of the present invention are directed to a
method of making a stent that may include forming a tube having a
metallic film in between two radial polymeric layers. The method
may further include fabricating a stent from the tube.
[0030] Certain other embodiments of the present invention are
directed to a method of making a stent that may include elongating
a polymeric tube so that a diameter of the stent decreases. The
method may further include positioning a metallic band around a
circumference of the elongated tube. The elongated polymeric tube
with the metallic band positioned around the tube may then be
heated. The method may further include allowing the heated tube to
radially expand so as to couple the metallic band to the tube. A
stent may be fabricated from the expanded tube.
[0031] Additional embodiments of the present invention are directed
to a radially expandable stent including a plurality of
interconnecting structural elements including fibers at least
partially embedded in a matrix polymer. The fibers may include a
material having a melting temperature greater than a melting
temperature of the matrix polymer. The fibers may be configured to
provide mechanical reinforcement to the stent due to a higher
strength and modulus along an axis of the fibers than the matrix
polymer.
[0032] Other embodiments of the present invention are directed to a
radially expandable stent including a plurality of interconnecting
structural elements including at least one radial fiber layer and
at least one radial polymer film layer. The fibers may include
material with a melting temperature greater than a melting
temperature than at least one polymer film layer. At least one
fiber layer may be at least partially embedded within at least one
polymer film layer. The fibers may be configured to provide
mechanical reinforcement to the stent due to a higher strength and
modulus along an axis of the fibers than the polymer film
layer.
[0033] Additional embodiments of the present invention are directed
to a radially expandable stent including a plurality of structural
elements including at least two radial fiber layers and at least
one radial polymer film layer. At least a portion of at least one
fiber layer may be embedded within at least a portion of at least
one polymer film layer. An orientation of fibers relative to a
cylindrical axis of the stent of at least one fiber layer may be
different from an orientation of fibers in another fiber layer.
[0034] Certain embodiments of the present invention are directed to
a radially expandable stent woven from at lease two types of
fibers. A first fiber may include a first polymer and the second
fiber may include a second polymer. The first polymer may have a
softening temperature lower than a softening temperature of the
second polymer. At least some of the fibers may have a flattened
radial profile that reduces the radial profile of the tube.
[0035] Other embodiments of the present invention are directed to a
radially expandable stent woven from fibers comprising a first
polymer and a second polymer. The first polymer may have a
softening temperature lower than a softening temperature of the
second polymer such that at least some of the fibers have a
flattened radial profile that reduces the radial profile of the
tube.
[0036] Additional embodiments of the present invention are directed
to a radially expandable stent including a metallic film coupled to
a plurality of portions of a surface of the stent such that the
metallic film is sufficiently radiopaque to allow the stent to be
visualized during use
[0037] Further embodiments of the present invention are directed to
a radially expandable stent including a plurality of
interconnecting structural elements such that the structural
elements may have two radial polymeric layers with metallic film
embedded in a plurality of locations in between the layers. The
metallic film may be sufficiently radiopaque to allow the stent to
be visualized during use.
BRIEF DESCRIPTION OF THE DRAWINGS
[0038] FIG. 1 depicts a three-dimensional view of a stent.
[0039] FIG. 2 depicts a schematic plot of the rate of
crystallization of a polymer as a function of temperature.
[0040] FIG. 3 depicts a schematic representation of a mixture of a
continuous polymer phase and a discrete fiber phase.
[0041] FIG. 4 depicts a fiber-reinforced tube with short
fibers.
[0042] FIG. 5 depicts an embodiment of a method of fabricating a
fiber reinforced tube.
[0043] FIG. 6A depicts a two-dimensional radial cut-off view of a
tube formed with a fiber layer and two polymer layers.
[0044] FIG. 6B depicts an expanded view of the layers from FIG.
6A.
[0045] FIG. 7 depicts a tube of helically wound fiber mesh.
[0046] FIG. 8 depicts a two-dimensional view of layers of a tube
formed with fiber layers and polymer layers.
[0047] FIG. 9 depicts a fiber mesh tube disposed on a mandrel.
[0048] FIG. 10 depicts alignment of struts or structural elements
with fibers.
[0049] FIG. 11 depicts a radial cross-section of a composite
fiber.
[0050] FIG. 12 depicts a radial cross-section of a system for
heating and flattening fibers of a fiber stent.
[0051] FIG. 13 depicts an expanded view of fibers in the system of
FIG. 12 prior to flattening the fibers.
[0052] FIG. 14 depicts an expanded view of fibers in the system of
FIG. 12 showing flattening of the fibers.
[0053] FIG. 15 depicts a polymeric tube with a circumferentially
aligned metallic band coupled or adhered to the surface of the
tube.
[0054] FIG. 16 depicts a polymeric tube with a longitudinally
aligned strip of metallic film coupled or adhered to the surface of
the tube.
[0055] FIG. 17 depicts a stent with a circumferentially aligned
metallic film on its surface.
[0056] FIG. 18 depicts a stent with a longitudinally aligned
metallic film on its surface.
[0057] FIG. 19 depicts a cross-sectional view of a sidewall of a
portion of a structural element of a stent with a coating above a
metallic film on a polymeric substrate.
[0058] FIG. 20 depicts a cross-sectional view of a sidewall of a
portion of a structural element of a stent with a metallic film
embedded between an abluminal layer and a luminal layer.
DETAILED DESCRIPTION OF THE INVENTION
[0059] Various embodiments of the present invention relate to
composite polymeric biodegradable implantable medical devices and
methods of making such devices. In general, a composite implantable
medical device is a device which is made up of two or more
macroscopically distinct materials that have different properties.
The composite device as a whole may have desirable properties of
two or more of the distinct materials. Therefore, desirable
mechanical and/or degradation properties may be obtained through
the use of a polymer composite structure.
[0060] For the purposes of the present invention, the following
terms and definitions apply:
[0061] The "glass transition temperature," T.sub.g, is the
temperature at which the amorphous domains of a polymer change from
a brittle vitreous state to a solid deformable or ductile state at
atmospheric pressure. In other words, the T.sub.g corresponds to
the temperature where the onset of segmental motion in the chains
of the polymer occurs. When an amorphous or semicrystalline polymer
is exposed to an increasing temperature, the coefficient of
expansion and the heat capacity of the polymer both increase as the
temperature is raised, indicating increased molecular motion. As
the temperature is raised the actual molecular volume in the sample
remains constant, and so a higher coefficient of expansion points
to an increase in free volume associated with the system and
therefore increased freedom for the molecules to move. The
increasing heat capacity corresponds to an increase in heat
dissipation through movement. T.sub.g of a given polymer can be
dependent on the heating rate and can be influenced by the thermal
history of the polymer. Furthermore, the chemical structure of the
polymer heavily influences the glass transition by affecting
mobility.
[0062] "Stress" refers to force per unit area, as in the force
acting through a small area within a plane. Stress can be divided
into components, normal and parallel to the plane, called normal
stress and shear stress, respectively. Tensile stress, for example,
is a normal component of stress applied that leads to expansion
(increase in length). In addition, compressive stress is a normal
component of stress applied to materials resulting in their
compaction (decrease in length). Stress may result in deformation
of a material, which refers to change in length. "Expansion" or
"compression" may be defined as the increase or decrease in length
of a sample of material when the sample is subjected to stress.
[0063] "Strain" refers to the amount of expansion or compression
that occurs in a material at a given stress or load. Strain may be
expressed as a fraction or percentage of the original length, i.e.,
the change in length divided by the original length. Strain,
therefore, is positive for expansion and negative for
compression.
[0064] "Solvent" is defined as a substance capable of dissolving or
dispersing one or more other substances or capable of at least
partially dissolving or dispersing the substance(s) to form a
uniformly dispersed mixture at the molecular- or ionic-size level.
The solvent should be capable of dissolving at least 0.1 mg of the
polymer in 1 ml of the solvent, and more narrowly 0.5 mg in 1 ml at
ambient temperature and ambient pressure. The "strength" of a
solvent refers to the degree to which a solvent may dissolve a
polymer. The stronger a solvent is, the more polymer the solvent
can dissolve.
[0065] Furthermore, a property of a material that quantifies a
degree of strain with applied stress is the modulus. "Modulus" may
be defined as the ratio of a component of stress or force per unit
area applied to a material divided by the strain along an axis of
applied force that results from the applied force. For example, a
material has both a tensile and a compressive modulus. A material
with a relatively high modulus tends to be stiff or rigid.
Conversely, a material with a relatively low modulus tends to be
flexible. The modulus of a material depends on the molecular
composition and structure, temperature of the material, amount of
deformation, and the strain rate or rate of deformation. For
example, below its T.sub.g, a polymer tends to be brittle with a
high modulus. As the temperature of a polymer is increased from
below to above its T.sub.g, its modulus decreases.
[0066] "Above" a surface or layer is defined as higher than or over
a surface or layer measured along an axis normal to a surface or
layer, but not necessarily in contact with the surface or
layer.
[0067] "Vicat Softening Temperature" (VST) is a measure of the
temperature at which a polymer starts to soften at specified test
conditions according to ISO 306. It is determined with a standard
indenter (a flat-ended needle of 1 mm 2 circular cross section)
penetrating into the surface of a test specimen under a predefined
load. The temperature at 1 mm penetration is quoted as the VST in
Co. VST gives an indication of a material's ability to withstand
limited short-term contact with a heated object.
[0068] The term "elastic deformation" refers to deformation of an
object in which the applied stress is small enough so that the
object moves towards its original dimensions or essentially its
original dimensions once the stress is released. However, an
elastically deformed polymer material may be prevented from
returning to an undeformed state if the material is below the
T.sub.g of the polymer. Below T.sub.g, energy barriers may inhibit
or prevent molecular movement that allows deformation or bulk
relaxation.
[0069] "Elastic limit" refers to the maximum stress that a material
will withstand without permanent deformation. The "yield point" is
the stress at the elastic limit and the "ultimate strain" is the
strain at the elastic limit. The term "plastic deformation" refers
to permanent deformation that occurs in a material under stress
after elastic limits have been exceeded.
[0070] The term "implantable medical device" is intended to
include, but not limited to, self-expandable stents,
balloon-expandable stents, stent-grafts, and grafts. In general, an
implantable medical device, such as a stent, can have virtually any
structural pattern that is compatible with a bodily lumen in which
it is implanted. The embodiments of the invention described herein
are generally applicable to implantable medical devices.
[0071] Typically, a stent is composed of a pattern or network of
circumferential rings and longitudinally extending interconnecting
structural elements of struts or bar arms. In general, the struts
are arranged in patterns, which are designed to contact the lumen
walls of a vessel and to maintain vascular patency. A myriad of
strut patterns are known in the art for achieving particular design
goals.
[0072] FIG. 1 depicts a three-dimensional view of a stent 10 which
shows struts 15. The embodiments disclosed herein are not limited
to stents or to the stent pattern illustrated in FIG. 1. The
embodiments are easily applicable to other patterns and other
devices. The variations in the structure of patterns are virtually
unlimited.
[0073] A stent such as stent 10 may be fabricated from a tube by
forming a pattern with a technique such as laser cutting.
Representative examples of lasers that may be used include an
excimer, carbon dioxide, and YAG. In other embodiments, chemical
etching may be used to form a pattern on the elongated tube.
[0074] In some embodiments, the diameter of the polymer tube prior
to fabrication of an implantable medical device may be between
about 0.2 mm and about 5.0 mm, or more narrowly between about 1 mm
and about 3 mm. Unless otherwise specified, the "diameter" of the
tube refers to the outside diameter of the tube.
[0075] Various embodiments of fabricating composite polymeric
implantable devices are disclosed herein. The continuous and
discrete phases may include polymeric or metallic materials or a
combination of polymeric and metallic materials.
[0076] In general, polymers can be biostable, bioabsorbable,
biodegradable, or bioerodable. Biostable refers to polymers that
are not biodegradable. The terms biodegradable, bioabsorbable, and
bioerodable, as well as degraded, eroded, and absorbed, are used
interchangeably and refer to polymers that are capable of being
completely eroded or absorbed when exposed to bodily fluids such as
blood and can be gradually resorbed, absorbed, and/or eliminated by
the body. In addition, a medicated stent may be fabricated by
coating the surface of the stent with an active agent or drug, or a
polymeric carrier including an active agent or drug.
[0077] A stent made from a biodegradable polymer is intended to
remain in the body for a duration of time until its intended
function of, for example, maintaining vascular patency and/or drug
delivery is accomplished. After the process of degradation,
erosion, absorption, and/or resorption has been completed, no
portion of the biodegradable stent, or a biodegradable portion of
the stent will remain. In some embodiments, very negligible traces
or residue may be left behind. The duration can be from about a
month to a few years, but is typically in the range of six to
eighteen months.
[0078] Biodegradation of polymers generally refers to changes in
physical and chemical properties that occur in a polymer upon
exposure to bodily fluids as in a vascular environment. The changes
in properties may include a decrease in molecular weight,
deterioration of mechanical properties, and decrease in mass due to
erosion or absorption. Mechanical properties may correspond to
strength and modulus of the polymer. Deterioration of the
mechanical properties of the polymer decreases the ability of a
stent, for example, to provide mechanical support in a vessel. The
decrease in molecular weight may be caused by, for example,
hydrolysis and/or metabolic processes. Hydrolysis is a chemical
process in which a molecule is cleaved into two parts by the
addition of a molecule of water.
[0079] Consequently, the degree of bulk degradation of a polymer is
strongly dependent on the diffusivity, and hence the diffusion rate
of water in the polymer. Several characteristics or parameters of
the degradation process are important in designing biodegradable
devices. These include an average erosion rate of a device, the
erosion profile, the half-life of the degrading polymer, and
mechanical stability of a device during the degradation process.
The "average erosion rate" may be an average erosion rate over any
selected time interval: Average erosion
rate=(m.sub.2-m.sub.1)/(t.sub.2-t.sub.1) where "m" refers to mass
of the device, "t" refers to a time during erosion, and m.sub.1 and
m.sub.2 are the masses of the device at t.sub.1 and t.sub.2 during
erosion. For instance, the selected time interval may be between
the onset of degradation and another selected time. Other selected
times, for example, may be the time for about 25%, 50%, 75%, or
100% (complete erosion) of the device to erode. Complete erosion
may correspond approximately to the time required for treatment by
the device.
[0080] The "half-life" of a degrading polymer refers to the length
of time for the molecular weight of the polymer to fall to one half
of its original value. See e.g., J. C. Middleton and A. J. Tipton,
Biomaterials, Vol. 21 (23) (2000) pp. 2335-2346.
[0081] In addition, metals may be considered to be biostable or
bioerodible. Some metals are considered bioerodible since they tend
to erode or corrode relatively rapidly when exposed to bodily
fluids. Biostable metals refer to metals that are not bioerodible.
Biostable metals have negligible erosion or corrosion rates when
exposed to bodily fluids.
[0082] In general, metal erosion or corrosion involves a chemical
reaction between a metal surface and its environment. Erosion or
corrosion in a wet environment, such as a vascular environment,
results in removal of metal atoms from the metal surface. The metal
atoms at the surface lose electrons and become actively charged
ions that leave the metal to form salts in solution.
[0083] Representative examples of bioerodible metals that may be
used to fabricate an implantable medical device may include, but
are not limited to, magnesium, zinc, and iron. In one embodiment, a
bioerodible metallic stent may be completely eroded when exposed to
bodily fluids, such as blood, between about a week and about three
months, or more narrowly, between about one month and about two
months.
[0084] As indicated above, implantable medical devices, such as a
stent, should be capable of exhibiting relatively high strength and
rigidity, as well as flexibility since devices have varied
mechanical requirements during use arising from stress imposed on
the device, both before and during treatment. "Use" includes
manufacturing, assembling (e.g., crimping a stent on balloon),
delivery of a stent through a bodily lumen to a treatment site, and
deployment of a stent at a treatment site. For example, a stent
requires radial or hoop strength and rigidity to resist radial
compressive forces.
[0085] The stress imposed on a stent, for example, during use
subjects individual structural elements to stress. During
deployment, the scaffolding and/or coating of a stent can be
exposed to stress caused by the radial expansion of the stent body.
In addition, the scaffolding and/or coating may be exposed to
stress when it is mounted on a catheter from crimping or
compression of the stent. After deployment, radial compressive
forces subject scaffolding and/or coating to stress. These stresses
can cause the scaffolding to fracture. Failure of the mechanical
integrity of the stent while the stent is localized in a patient
can lead to serious risks for a patient. For example, there is a
risk of embolization caused by a piece of the polymeric scaffolding
and/or coating breaking off from the stent.
[0086] Conventional methods of constructing a stent from a polymer
material involve forming a polymer tube based on a single polymer
or polymer blend and then laser cutting a pattern into the tube.
Alternatively, a polymer tube may be formed from sheets or films
that are rolled and bonded. Polymer tubes and sheets may be formed
by various methods, including, but not limited to extrusion,
injection molding, or blow molding.
[0087] In extrusion, a polymer melt of a single polymer or polymer
blend is conveyed through an extruder which is then formed into a
tube. Extrusion tends to impart large forces on the molecules in
the axial direction of the tube due to shear forces on the polymer
melt. The shear forces arise from forcing the polymer melt through
a die and pulling and forming the polymer melt into the small
dimensions of a tube. As a result, polymer tubes formed by
conventional extrusion methods tend to possess a significant degree
of axial polymer chain alignment. However, such conventionally
extruded tubes tend to possess no or substantially no polymer chain
alignment in the circumferential direction.
[0088] Due to stresses imposed on an implantable medical device
during use, it is important for the mechanical stability of a
device to have an adequate magnitude of strength both in axial and
circumferential directions. The direction of stress in structural
members can be in various directions between axial and
circumferential. Therefore, an adequate balance of axial and
circumferential strength is also important for mechanical
stability. The relative amount of axial and circumferential
orientation may depend on a number of factors such as the stent
pattern.
[0089] A radially expandable device, such as a stent, without an
adequate magnitude and balance of strength in axial and
circumferential directions may tend to be more prone to mechanical
instability. For example, a stent made from a tube with an adequate
magnitude and balance of strength in the radial and axial
directions may be less susceptible to cracking during crimping and
deployment. Therefore, it may be desirable to fabricate an
implantable medical device with desired strength and balance in the
axial and circumferential directions.
[0090] Desired strength in both directions may be achieved in a
number of ways. Some embodiments may include fabricating a device
from a tube of a composite structure with fibers mixed with a
continuous phase. In other embodiments, polymer chain alignment may
be induced along the circumferential direction to increase
strength.
[0091] In certain embodiments, a polymer for a fiber may be
selected that can form crystalline regions with a high modulus and
the polymer for a continuous phase may include a relatively
flexible polymer. Representative examples of polymers that may be
used for fiber reinforcement include, but are not limited to,
poly(L-lactide) and polyglycolide. Representative polymers that may
be used for a continuous phase may include, but are not limited to,
poly(DL-lactide) and poly(.epsilon.-caprolactone).
[0092] An implantable medical device, such as a stent, with an
adequate magnitude and balance of both circumferential strength and
modulus may be less susceptible to cracking during the crimping
process. In addition, increased circumferential strength and
modulus may allow a decrease in strut width, or generally, a
decrease in form factor of a stent. Implantable medical devices
fabricated from tubes with adequate strength in both the axial and
circumferential directions may possess mechanical properties
similar to or better than metal stents with an acceptable wall
thickness and strut width.
[0093] It is well know by those skilled in the art that molecular
orientation or alignment of polymer chains in a polymer is a
particularly important phenomenon that strongly influences bulk
polymer properties. For example, strength, modulus, yield stress
behavior, and elongation to break are a few of the important
properties that may be influenced by orientation of polymer chains
in a polymer. Orientation refers to the degree of alignment of
polymer chains along a longitudinal or covalent axis of the polymer
chains. The degree of molecular orientation in a polymeric material
may be induced by applying stress along a preferred direction.
[0094] Polymers in the solid state may have amorphous regions and
crystalline regions. Crystalline regions include highly oriented
polymer chains in an ordered structure. An oriented crystalline
structure tends to have high strength and high modulus (low
elongation with applied stress) along an axis of alignment of
polymer chains. On the other hand, amorphous polymer regions
include relatively disordered polymer chains that may or may not be
oriented in a particular direction. However, a high degree of
molecular orientation may be induced even in an amorphous region.
An oriented amorphous region also tends to have high strength and
high modulus along an axis of alignment of polymer chains.
Additionally, for some polymers under some conditions, induced
orientation in an amorphous polymer may be accompanied by
crystallization of the amorphous polymer.
[0095] A polymer may be completely amorphous, partially
crystalline, or almost completely crystalline. A partially
crystalline polymer includes crystalline regions separated by
amorphous regions. The polymer chains of the crystalline regions
are not all necessarily oriented in the same direction. However, a
high degree of orientation of crystallites may be induced by
applying stress to a semi-crystalline polymer. The stress may also
induce orientation of polymer chains in amorphous regions of a
polymer.
[0096] Polymer tubes fabricated in a conventional manner using
extrusion, blow molding, or injection molding based on a single
polymer or polymer blend tend to have inadequate strength and
rigidity in the circumferential direction. This is due to low
polymer chain alignment in the circumferential direction.
[0097] Various embodiments of the present invention include
implantable medical devices, such as stents, and methods of
fabricating such devices from a composite including a continuous
phase and a discrete phase. The continuous phase may include a
polymeric matrix and the discrete phase may include fibers mixed,
dispersed, and/or embedded in the matrix. Additionally, either or
both the continuous phase or the discrete phase may include an
active agent.
[0098] In some embodiments, the discrete phase may include
radiopaque materials. The radiopaque materials may include, for
example, metals; alloys; or mixtures of polymers and metal or
alloys. In one embodiment, the discrete phase may include metallic
fibers, wires, bands, or strips. The metals may include erodible
metals, biostable metals, or mixtures of biostable and bioerodible
metals. Representative metals that may be used in the discrete
phase may include, but are not limited to, magnesium, zinc, iron,
platinum, and gold.
[0099] A "fiber" may be defined as a unit of matter having a length
substantially longer than its width or diameter. As used herein, a
fiber can include, but is not limited to, a filament, a strip, or a
wire.
[0100] In some embodiments, a polymeric fiber may be formed using
any of a number of methods known in the art including, but not
limited to, melt spinning, wet spinning, dry spinning, gel
spinning, electrospinning, or an atomizing process. Fibers may be
fabricated with relatively high polymer chain orientation along the
fiber axis, and thus relatively high strength and stiffness.
[0101] "Spinning" of polymeric fibers generally involves the
extrusion or forcing of a thick, viscous fluid, which is either a
polymer melt or solution, through the tiny holes of a device called
a spinneret to form continuous filaments of semi-solid polymer. The
spinneret has a multiplicity of holds through which polymer melt or
solution pass through. In their initial state, the fiber-forming
polymers are solids and therefore must be first converted into a
fluid state for extrusion. This is usually achieved by melting, if
the polymers are thermoplastic (i.e., they soften and melt when
heated), or by dissolving them in a suitable solvent if they are
non-thermoplastic. If they cannot be dissolved or melted directly,
they must be chemically treated to form soluble or thermoplastic
derivatives.
[0102] In melt spinning, the fiber-forming polymer is melted for
extrusion through the spinneret and then solidified by cooling. Wet
spinning involves forming a fiber from a polymer dissolved in a
solvent. The polymer solution is pumped through a spinneret that is
submerged in a chemical bath. The dissolved polymer is immiscible
in the chemical bath. As the filaments emerge from the spinneret,
the polymer precipitates from solution and solidifies.
[0103] Dry spinning also involves forming fibers from a polymer
solution. The polymer solution is pumped through the spinneret.
However, instead of precipitating the polymer by dilution or
chemical reaction, solidification is achieved by evaporating the
solvent in a stream of air or inert gas.
[0104] Gel Spinning is a type of wet spinning, but is a special
process used to obtain high strength or other special fiber
properties. In this process, ultra-high molecular weight polymer is
dissolved in a solvent at very low concentration. The concentration
is much lower than that typically used in wet spinning and dry
spinning processes. The polymers or fibers precipitate from
solution and solidify in a chemical bath or in a chilled water
bath. The fiber is then drawn to orient the polymer molecules. The
draw-down ratio is also typically much higher than for wet spinning
and dry spinning processes.
[0105] The draw-down ratio is defined as the ratio of the length of
a drawn fiber to the original length of the fiber. The draw-down
ratio for gel spinning can be up to 40:1, while the drawn-down
ratio for wet or melt spinning can be about 3-15:1.
[0106] The draw-down ratio is defined as the ratio of the length of
a drawn fiber to the length of an as-spun fiber. The as-spun fiber
refers to a solidified fiber formed from solution or melt. The
draw-down ratio for gel spinning can be up to 40:1, while the
drawn-down ratio for wet or melt spinning can be about 3-15:1.
[0107] In a dry-jet-wet spinning method, the polymer is not in a
true liquid state during extrusion. The polymer chains are bound
together at various points in liquid crystal form. The chains are
not completely separated, as they would be in a true solution. This
produces strong inter-chain forces in the resulting filaments that
can significantly increase the tensile strength of the fibers. In
addition, the liquid crystals are aligned along the fiber axis by
the shear forces during extrusion. The filaments emerge with a
relatively high degree of orientation relative to each other,
further enhancing strength. The filaments first pass through air
and then are cooled further in a liquid bath. The draw-down ratio
in dry-jet-wet spinning is typically less than 1.03:1.
[0108] Electrospinning and atomizing processes may be used to
produce nanofibers. A "nanofiber" refers to a fiber with a
dimension in the range of about 1 nm to about 10,000 nm.
Electrospinning makes use of electrostatic and mechanical force to
spin fibers from the tip of a fine orifice or spinneret. In
electrospinning, a polymer is dissolved in a solvent or a polymer
melt and is placed in a spinneret (e.g., a glass pipet) sealed at
one end. The spinneret is maintained at positive or negative charge
by a power supply, for example. When the electrostatic repelling
force overcomes the surface tension force of the polymer solution
or melt, the liquid spills out of the spinneret and forms an
extremely fine continuous filament.
[0109] The strength and modulus of the spun fibers may be increased
by drawing. Drawing involves applying tension along the fiber axis.
Fibers may be drawn while extruded fibers are solidifying and/or
after they have hardened. Drawing tends to pull the molecular
chains together and orient them along the fiber axis, creating a
considerably stronger and rigid fiber along the fiber axis.
[0110] As indicated above, favorable mechanical and degradation
properties of a stent may be obtained by fabricating the stent as a
composite. Individual characteristics of the stent (i.e., rigidity,
strength, longitudinal flexibility, degradation rate) may be
provided by one or more of the macroscopically distinct materials
that make up the composite. Thus, one benefit of a composite
structure is that individual characteristics of a stent may be
tuned independently or more independently than a stent fabricated
from a single polymer or blend.
[0111] To provide strength and rigidity to a stent, the fiber may
be fabricated to be relatively strong and stiff with a high modulus
along the fiber axis. The continuous phase may be configured to
have different properties than the discrete fiber phase. For
instance, the continuous phase may be configured to have a lower
modulus, and thus greater flexibility than the discrete phase.
Therefore, the continuous phase may be configured to provide the
required flexibility for the stent. As indicated above, polymers
below their T.sub.g tend to be relatively brittle or inelastic and
are more flexible and more easily deformed than above their
T.sub.g. Therefore, a flexible continuous phase may be obtained by
using polymers with a T.sub.g above a body temperature.
[0112] Additionally, the degradation behavior of the fiber and
continuous phases may be configured to have various combinations.
In one embodiment, degradation rates of the fiber and continuous
phase may be approximately the same. Alternatively, the degradation
rates of the fiber and the continuous phases may be different. The
degradation rate of the fiber may be faster or slower than the
continuous phase. As discussed herein, the degradation rate of
phases may be controlled the choice of polymer, the molecular
weight, and the crystallinity of the polymers of the phases.
[0113] Moreover, there are numerous ways that the properties of
polymers in the stent may be controlled or modified. These include
a suitable choice of polymers or chemical component groups in
polymers for the discrete and continuous phases since different
polymers have different mechanical properties and degradation
rates. In addition, as described below, various properties depend
on the molecular weight of polymers. In addition, certain
properties of polymers are also related to the degree of
crystallinity in a polymer. Thus, these properties of the discrete
and continuous phases may be modified independently by choice of
polymers and modifying the molecular weight and crystallinity of
the polymers in these phases.
[0114] Mechanical properties such as strength and modulus, the
degradation behavior of polymer, and melting temperatures depend
upon the molecular weight. In general, the higher the molecular
weight, the stronger and stiffer (higher modulus) a polymer is.
Therefore, the strength and modulus of a fiber may be further
enhanced by fabricating a fiber with a higher molecular weight.
[0115] Additionally, the degradation rate of a polymer decreases as
the molecular weight increases. Also, the melting temperature
increases with molecular weight. In some embodiments, the same type
of polymer with different molecular weights may be used for both
the fiber and continuous phase in a composite for a device. Due to
the different molecular weight of the continuous and discrete
phases, the mechanical properties, degradation behavior, and
melting temperature of the phases may be different.
[0116] As indicated above, the degree of crystallinity in a polymer
is related to the mechanical properties such as the strength and
modulus of a material. The higher the degree of crystallinity, the
stronger and stiffer a polymer is along the direction of molecular
orientation of crystalline structures in the polymer.
[0117] In addition, the degree of crystallinity is also related to
the diffusion rate of fluids, and hence, the erosion rate of a
biodegradable polymer. In general, the diffusion rate of a fluid
through a polymer decreases as the degree of crystallinity
increases. Therefore, it is expected that the diffusion rate of
water and bodily fluids is lower in crystalline and
semi-crystalline polymers than in amorphous polymers. Thus, the
erosion rate of a biodegradable polymeric region may be controlled
by modifying the degree of crystallinity in a continuous polymeric
phase of a composite, for example.
[0118] In one embodiment, the crystallinity of a polymer may be
modified by heating the polymer. Heating a polymer can alter the
degree of crystallinity and/or size of crystalline regions in a
polymer material. The degree of crystallinity may be altered by
heating the polymer within a particular temperature range. Heating
a polymer material to a temperature below the glass transition
temperature, T.sub.g, of the polymer does not significantly alter
the molecular structure, and hence, the mechanical properties of
the material. Below T.sub.g, energy barriers to segmental motion of
the chains of a polymer inhibit or prevent alteration of molecular
structure of a polymeric material.
[0119] In general, crystallization may occur in a polymeric
material that is heated to a temperature between T.sub.g and the
melting temperature, T.sub.m, of the polymer. As a result, heating
a polymer to a temperature between the T.sub.g and the T.sub.m of
the polymer increases the modulus of the polymer.
[0120] FIG. 2 depicts a schematic plot of the rate of
crystallization of a polymer as a function of temperature.
(Rodriguez, F., Principles of Polymer Systems, 2 ed., McGraw Hill
(1982)) FIG. 2 shows that the rate of polymer crystallization
increases as the temperature is increased from below the T.sub.g of
the polymer or is decreased from above the T.sub.m of the polymer.
The rate of crystallization reaches a maximum 16 somewhere between
the T.sub.g and the T.sub.m. FIG. 2 shows that effectively no
crystallization occurs below the T.sub.g or above the T.sub.m.
[0121] In addition, as indicated above, an amorphous polymer may be
formed by heating a polymer material. Above the T.sub.m, a
polymeric material becomes a disordered melt and cannot crystallize
and any crystallinity present is destroyed. Quenching a polymer
melt from above the T.sub.m of the polymer to a temperature below
the T.sub.g of the polymer may result in the formation of a solid
amorphous polymer. The resulting amorphous polymer material may
have a lower modulus and be a more flexible or a less stiff
material than before heating.
[0122] In certain embodiments, a method of fabricating a
fiber-reinforced stent may include forming a mixture including a
matrix polymer and a plurality of short or staple fibers. The
fibers may include a material having a melting temperature greater
than a melting temperature of the matrix polymer.
[0123] In one embodiment, the matrix polymer may be a biostable or
biodegradable polymer or a combination thereof. The material of the
fibers may also be a biostable or biodegradable polymer or a
combination thereof. In some embodiments, the material of the
fibers may be a biostable and/or erodible metal. In an embodiment,
the fibers may be combination of a polymeric and a metallic
material. For example, the fibers may be a mixture of polymeric and
metallic particles.
[0124] As described above, the mixture formed may be a composite
material. FIG. 3 depicts a schematic representation of the mixture.
The matrix polymer may correspond to a continuous phase 20 and
short fibers 25 may correspond to a discrete phase.
[0125] In one embodiment, the short fibers may be composed of the
same or similar polymeric material as the continuous polymeric
phase. Alternatively, the short fibers may be a mixture of fibers
with different properties. For example, the short fibers may be a
mixture of fibers having different degradation rates and/or
mechanical properties.
[0126] In one embodiment, the mixture may be formed by mixing the
matrix polymer and the fibers in a mixing apparatus at a
temperature that is greater than the melting temperature of the
matrix polymer and less than the melting temperature of the fiber
material. Therefore, a polymeric melt continuous phase containing
the matrix polymer may be mixed with a discrete fiber phase which
is below the fiber material melting temperature.
[0127] FIG. 4 depicts a fiber-reinforced tube with short fibers.
Tube 30 includes a plurality of short fibers 35 embedded in a
continuous polymer phase 38. As shown in FIG. 4, fibers 35 are
oriented in arbitrary directions with respect to the axis of the
tube. The fibers provide mechanical reinforcement axially,
circumferentially, and orientations between the two. Thus, fibers
enhance the mechanical stability of the tube and a stent formed
from the tube.
[0128] In addition, circumferential strength can be further
enhanced through radial expansion of the tube. Radial expansion
enhances the circumferential strength of the tube due to induced
polymer chain alignment of the continuous phase and induced
circumferential alignment of the short fibers.
[0129] Embodiments of the method may further include disposing the
mixture in a tube or sheet forming apparatus to form a tube or a
sheet. The apparatus may be heated so that a temperature of the
mixture in the apparatus is greater than the melting temperature of
the matrix polymer and less than the melting temperature of the
fiber material. In some embodiments, at least a portion of the
matrix polymer may be a polymer melt. In addition, the mixture may
then be cooled below the melting temperature of the matrix
polymer.
[0130] As indicated above, a polymer melt may be cooled in such a
way so as to control the degree of crystallinity of the formed
tube. Thus, in some embodiments, the formed tube or sheet may be
cooled to a temperature below the melting temperature of the matrix
polymer such that a majority of the matrix polymer in the formed
tube is either amorphous or crystalline.
[0131] In one embodiment, the forming apparatus may be an injection
molding apparatus.
[0132] The mixture may be injected at a temperature above the
melting temperature of the matrix polymer and less than the melting
temperature of the fiber such that at least a portion of the matrix
polymer is a polymer melt. The mold may be heated by a heating
device or in a chamber so that the temperature of the mixture in
the mold is above the melting temperature of the matrix polymer and
less than the melting temperature of the fibers.
[0133] Alternatively, in another embodiment, the mixture may be
placed into the mold at a temperature below the melting temperature
of the matrix polymer. The heated mold may then melt the matrix
polymer by heating the mixture to a temperature above the melting
temperature of the matrix polymer and less the melting temperature
of the fibers.
[0134] In another embodiment, the forming apparatus may be an
extruder. The mixture may be conveyed into the extruder at a
temperature below the melting temperature of the matrix polymer.
Alternatively, the mixture may be conveyed into the extruder at a
temperature above the melting temperature of the matrix polymer and
less than the melting temperature of the fiber such that at least a
portion of the matrix polymer is a polymer melt. The mixture may be
heated in the extruder so that its temperature is above the melting
temperature of the matrix polymer and less than the melting
temperature of the fiber.
[0135] Additionally, the method may further include fabricating a
stent from the tube or sheet. As indicated above, a stent may be
fabricated from a tube by forming a pattern on the tube including a
plurality of interconnecting structural elements. Also, a sheet may
be rolled into a tube and a pattern may be formed onto the
tube.
[0136] FIG. 5 depicts a schematic representation of an embodiment
of a method of fabricating a fiber reinforced tube, as described
above. A matrix polymer 40 is conveyed into an extruder 45 as
either a polymer melt or a solid. Extruder 45 melts and mixes
polymer 40 to form a relatively low viscosity fluid 50. Fluid 50 is
fed from extruder 45 into a mixing apparatus 55. Fibers 60 are also
fed into mixing apparatus 55. Mixing apparatus 55 mixes fluid 50
with fibers 60 to produce mixture 65. Mixture 65 is conveyed
through a die 70 into a forming apparatus 75 to form a tube or a
sheet. Forming apparatus 75 may be, for example, an injection
molding apparatus or an extruder.
[0137] In some embodiments, the short fibers may be made by forming
fibers as described above, and cutting them into short lengths. In
one embodiment, a length of at least a portion of the short fibers
is substantially smaller than a diameter of the formed tube. For
example, in some embodiments, the short fibers may be less than
0.05 mm long. In other embodiments, the short fibers may be between
0.05 and 8 mm long or more narrowly between 0.1 and 0.4 mm long or
0.3 and 0.4 mm long.
[0138] In other embodiments, a method of fabricating a
fiber-reinforced stent may include forming a tube including at
least one fiber layer and at least one polymer film layer. In one
embodiment, at least one fiber layer alternates with at least one
film layer. In an embodiment, fibers in a fiber layer may include
at least one material having a melting temperature greater than
melting temperatures at least one polymer film layer.
Alternatively, the method may include forming a layered sheet
including at least one fiber layer and at least one polymer film
layer.
[0139] In one embodiment, at least one polymer film layer may
include a biodegradable polymer. The material of the fibers of at
least one fiber layer may include a biostable or biodegradable
polymer or a combination thereof. In an embodiment, the material of
the fibers of at least one fiber layer may include a biostable
and/or erodible metal. In an embodiment, the fibers of at least one
fiber layer may be a combination of a polymeric and a metallic
material. For example, the fibers may be a mixture of polymeric and
metallic particles.
[0140] In one embodiment, the fibers of the fiber layers may be
composed of the same or similar polymeric material as the polymer
film layers. Alternatively, the fiber layers may be a mixture of
fibers with different properties. Some embodiments may include at
least one fiber layer with different properties than another fiber
layer. For example, different fiber layers may have different
degradation rates and/or mechanical properties.
[0141] In other embodiments, the polymer film layers may have the
same or similar properties. Alternatively, at least one polymer
film layer may have different properties than another polymer film
layer. For example, different polymer film layers may have
different degradation rates and/or mechanical properties.
[0142] In one embodiment, the fiber layer may be a woven structure.
A woven structure may refer to any structure produced from between
one and several hundred or more fibers that are woven, braided,
knitted, helically wound, and/or intertwined in any manner, at
angles between 0.degree. and 180.degree. degrees with the
cylindrical axis of the tube, depending upon the overall geometry
and dimensions desired.
[0143] FIG. 6A depicts a two-dimensional radial cut-off view of a
tube formed with a fiber layer 80 between two polymer film layers
82 and 84. FIG. 6B depicts an expanded view of the layers. Fiber
layer 80 is at least partially embedded in polymer from polymer
film layers 82 and 84 due to melting of the polymer film
layers.
[0144] FIG. 7 depicts a tube 90 of helically wound fiber mesh
including two sets of helically wound fibers 92 and 94. Tube 90 has
a cylindrical axis 96. Coordinate system 98 shows the relative
orientation with respect to axis 96. Fibers 92 have a relative
orientation greater than 90.degree. and fibers 92 have a relative
orientation less than 90.degree..
[0145] In some embodiments, an orientation of fibers in one fiber
layer may be different from an orientation of fibers in another
fiber layer. This may further enhance the mechanical stability of
stent. One embodiment may include one fiber layer with a set of
fibers with an orientation greater than 90.degree. and another
fiber layer with a set of fibers with an orientation less than
90.degree..
[0146] For example, FIG. 8 depicts a two-dimensional view of layers
of a tube formed with fiber layers 100 and 104 and polymer film
layers 108, 112, and 116. Fiber layer 100 may include fibers with
an orientation greater than 90.degree., such as fibers 92 in FIG.
7. Fiber layer 104 may include fibers with an orientation less than
90.degree., such as fibers 94 in FIG. 7.
[0147] In one embodiment, the tube may be formed by disposing the
layers over a mandrel. For example, FIG. 9 depicts a helically
wound fiber mesh 120 disposed on a mandrel 124. A polymer film
layer may be disposed over mandrel 124 before disposing fiber mesh
120 over mandrel 124. A polymer film layer may be disposed over
mandrel 124 followed by another fiber layer, another polymer film
layer, and so on.
[0148] Additionally, the method may further include heating the
tube or sheet to a temperature greater than the melting
temperatures of at least one polymer film layer and less than the
melting temperature of a material of the fibers. Heating the tube
or sheet may melt at least a portion of the polymer of the polymer
film layers.
[0149] In one embodiment, at least a portion of the fiber layers
may become embedded within at least a portion of the melted polymer
of a polymer film layer. In some embodiments, the heated tube may
be cooled and a stent may then be fabricated from the cooled tube.
As described above, the heated tube may be cooled in such a way to
control the degree crystallinity of the cooled polymer film layers
of the formed tube.
[0150] As indicated above, nanofibers may be used in fabricating
the stent. Nanofibers are particularly desirable when fabricating a
layered structure since a larger number of layers may be formed. In
general, the more the number of layers, the stronger the composite
structure. The number of layers may be limited if fibers larger
than nanofibers are used since the structure may become thicker
than desirable.
[0151] In other embodiments, a method of fabricating a
fiber-reinforced stent may include forming a coating layer
including a coating polymer over a tube-shaped fiber layer having a
plurality of fibers. Alternatively, the fibers may be formed into a
sheet. The plurality of fibers shaped into a tube may be a woven
structure, as described above.
[0152] In one embodiment, the coating polymer may include a
biostable and/or biodegradable polymer or a combination thereof. A
material of the fibers of at least one fiber layer may include a
biostable and/or biodegradable polymer or a combination thereof. In
an embodiment, the material of the fibers may include a biostable
and/or erodible metal. In another embodiment, the fibers may be a
combination of a polymeric and a metallic material. For example,
the fibers may be a mixture of polymer and metallic particles.
[0153] In one embodiment, the coating layer may be formed by
applying a fluid including the coating polymer dissolved in a
solvent. The material of the fibers may be insoluble or have a
relatively low solubility in the solvent. In some embodiments, the
coating may include an active agent. The fluid may include an
active agent dissolved or dispersed in the fluid. In an embodiment,
the material of the fiber may have a melting temperature greater
than a melting temperature of the coating polymer. Additionally,
all or a majority of the solvent may be removed from the applied
fluid.
[0154] Furthermore, the fluid may be applied on the tube in a
variety of ways known in the art. For example, the fluid may be
sprayed on the tube or the tube may be dipped in the fluid. In one
embodiment, the fiber layer may be disposed on a mandrel and then
dipped in and/or sprayed with the fluid.
[0155] In one embodiment, the fiber layer may be disposed on a
mandrel over a polymer layer including the coating polymer or
another type of polymer previously formed on the mandrel. The
previously formed polymer layer on the mandrel may be formed by
dipping and/or spraying, as described above.
[0156] In some embodiments, after forming the coating, the tube or
sheet may be heated to a temperature above the melting temperature
of the coating polymer and below the melting temperature of the
material of the fiber. The tube or sheet may then be cooled to a
temperature below the melting temperature of the coating polymer
such that a majority of the coating polymer in the formed tube or
sheet is amorphous, crystalline, or partially crystalline.
[0157] In some embodiments, a method of fabricating a
fiber-reinforced stent may include
[0158] disposing a plurality of fibers within a mold for forming a
structure. The structure may be, for example, a tube or a sheet.
The fibers may be disposed within the mold in a number of ways. One
embodiment may include disposing short fibers, as described above,
in a random or substantially random fashion within the mold. In
another embodiment, long fibers may be wound around a mandrel,
disposed in the mold, in a helical or other fashion. In one
embodiment, a woven structure, as described above, may be disposed
in the mold.
[0159] Additionally, the method may further include disposing a
matrix polymer that is partially or completely molten into the mold
to at least partially embed the fibers within the molten polymer.
In one embodiment, the fibers may include a material having a
melting temperature greater than a melting temperature of the
matrix polymer. The temperature of the molten polymer and the
fibers within the mold may be less than a melting temperature of
the material of the fiber.
[0160] In one embodiment, the matrix polymer may be a biostable or
biodegradable polymer or a combination thereof. The material of the
fibers may also be a biostable or biodegradable polymer or a
combination thereof. In some embodiments, the material of the
fibers may be a biostable and/or erodible metal. In an embodiment,
the fibers may be combination of a polymeric and a metallic
material. For example, the fibers may be a mixture of polymer and
metallic particles.
[0161] The molten polymer may then be cooled to form the structure
and a stent may be fabricated from the cooled structure. As
indicated above, the polymer melt may be cooled in such a way to
control the degree crystallinity of the matrix polymer.
[0162] A stent may be formed from a tube by forming a pattern in
the tube including a plurality of interconnecting structural
elements. As indicated above, a stent may be fabricated from a
sheet by forming a tube from the sheet and forming a pattern in the
tube including a plurality of interconnecting structural
elements.
[0163] In one embodiment, a medicated stent may be fabricated by
disposing an active agent into the mold. The active agent may be
mixed or dispersed within the molten matrix polymer. Alternatively,
the active agent may be mixed or dispersed with the fiber. In
another embodiment, a coating including an active agent may be
applied to the stent.
[0164] In further embodiments, a method of fabricating a
fiber-reinforced stent may include
[0165] disposing a plurality of fibers in an extruder for forming a
structure. The structure may be, for example, a tube or a sheet. As
described above, the fibers may be disposed within the extruder in
a number of ways. One embodiment may include disposing short
fibers, as described above, in a random or substantially random
fashion within the extruder. In another embodiment, long fibers may
be wound around a mandrel, disposed in the mold, in a helical or
other fashion. In one embodiment, a woven structure, as described
above, may be disposed in the extruder.
[0166] Additionally, the method may further include conveying a
matrix polymer into the extruder. In one embodiment, the matrix
polymer may have a melting temperature less than a melting
temperature of a material of the fiber. In addition, the structure
may then be formed with the extruder at a temperature greater than
the melting temperature of the matrix polymer and less than the
melting temperature of the material of the fiber. In an embodiment,
at least some of the fibers may become embedded within matrix
polymer.
[0167] In one embodiment, the matrix polymer may be a biostable or
biodegradable polymer or a combination thereof. The material of the
fibers may also be a biostable or biodegradable polymer or a
combination thereof. In some embodiments, the material of the
fibers may be a biostable and/or erodible metal. In an embodiment,
the fibers may be combination of a polymeric and a metallic
material. For example, the fibers may be a mixture of polymer and
metallic particles.
[0168] The molten polymer may then be cooled to form the structure
and a stent may be fabricated from the cooled structure. A stent
may be fabricated from a tube or a stent as described above. As
indicated above, the molten polymer melt may be cooled in such a
way to control the degree crystallinity of the matrix polymer.
[0169] In some embodiments, a medicated stent may be fabricated by
conveying an active agent into the extruder. In other embodiments,
at least some of the fibers may include an active agent.
[0170] As indicated above, a pattern including a plurality of
interconnecting structural elements may be formed by laser cutting
the pattern. In one embodiment, the pattern can be cut such that at
least a portion of the structural elements may be aligned or
substantially aligned with an orientation of at least some of the
fibers. For example, FIG. 10 illustrates a pattern of struts or
structural elements 130 that are aligned with the orientation of
fiber segments 132, 134, 136, and 138.
[0171] In further embodiments, the circumferential strength and
rigidity of a fiber reinforced stent may be enhanced by radial
expansion of a tube. Polymer chain alignment in the continuous
polymer phase may be induced along the circumferential direction to
increase strength. In addition, radial expansion may also induce
alignment of fibers in the tube, further enhancing circumferential
strength.
[0172] The degree of polymer chain alignment induced with applied
stress may depend upon the temperature of the polymer. For example,
below the T.sub.g of a polymer, polymer segments may not have
sufficient energy to move past one another. In general, polymer
chain alignment may not be induced without sufficient segmental
mobility. Above T.sub.g polymer chain alignment may be readily
induced with applied stress since rotation of polymer chains, and
hence segmental mobility, is possible. Between T.sub.g and the
melting temperature of the polymer, T.sub.m, rotational barriers
exist. However, the barriers are not great enough to substantially
prevent segmental mobility. As the temperature of a polymer is
increased above T.sub.g, the energy barriers to rotation decrease
and segmental mobility of polymer chains tends to increase. Thus,
as the temperature increases, polymer chain alignment is more
easily induced with applied stress.
[0173] Rearrangement of polymer chains may take place when a
polymer is stressed in an elastic region and in a plastic region of
the polymer material. A polymer stressed beyond its elastic limit
to a plastic region generally retains its stressed configuration
and corresponding induced polymer chain alignment when stress is
removed. The polymer chains may become oriented in the direction of
the applied stress. The stressed polymer material may have a higher
tensile strength and modulus in the direction of the applied
stress.
[0174] Additionally, heating a polymer may facilitate deformation
of a polymer under stress, and hence, modification of the
mechanical properties of the polymer. A polymer deformed
elastically with stress facilitated with heating may retain induced
polymer chain alignment by cooling the polymer before relaxing to
or towards an unstrained state.
[0175] In some embodiments, a polymer tube may be deformed at a
temperature below the T.sub.g of the polymer of the continuous
phase. Alternatively, it may be desirable to deform the tube in a
temperature range greater than or equal to the T.sub.g of the
continuous phase polymer and less than or equal to the T.sub.m of
the polymer. As indicated above, a polymeric material deforms much
more readily due to segmental motion of polymer chains above
T.sub.g. Deformation induces polymer chain alignment that may occur
due to the segmental motion of the polymer chains. Therefore,
heating the polymer tube prior to or contemporaneously with
deformation may facilitate deformation particularly for polymers
with a T.sub.g below an ambient temperature. Heating the tube
contemporaneously with the deformation may be desirable since the
deformation may occur at a constant or nearly constant temperature.
Therefore, the induced polymer alignment and material properties
may be constant or nearly constant.
[0176] In some embodiments, a fiber-reinforced polymer tube may be
deformed radially by increasing a pressure in a polymer tube, for
example, by conveying a fluid into the tube. Tension and/or torque
may also be applied to the tube. The tube may be positioned in an
annular member or mold. The mold may act to control the degree of
radial deformation of the tube by limiting the deformation of the
outside diameter or surface of the tube to the inside diameter of
the mold. The inside diameter of the mold may correspond to a
diameter less than or equal to a desired diameter of the polymer
tube.
[0177] The polymer tube may also be heated prior to, during, and
subsequent to the deformation. In general, it is desirable for the
temperature during deformation to be greater than or equal to a
glass transition temperature of the polymer and less than or equal
to a melting temperature of the polymer. The polymer tube may be
heated by the fluid and/or the mold.
[0178] Certain embodiments may include first sealing, blocking, or
closing a polymer tube at a distal end. The end may be open in
subsequent manufacturing steps. A fluid, (conventionally an inert
gas such as air, nitrogen, oxygen, argon, etc.) may then be
conveyed into a proximal end of the polymer tube to increase the
pressure in the tube. The pressure of the fluid in the tube may act
to deform the tube.
[0179] The increased pressure may deform the tube radially and/or
axially. The fluid temperature and pressure may be used to control
the degree of radial deformation by limiting deformation of the
inside diameter of the tube as an alternative to or in combination
with using the mold. In addition, it may be desirable to increase
the pressure to less than about an ultimate stress of the
continuous phase polymer to inhibit or prevent damage to the tube.
The continuous phase polymer may be deformed plastically or
elastically. As indicated above, a polymer elongated beyond its
yield point tends to retain its expanded configuration, and hence,
tends to retain the induced molecular orientation.
[0180] Additionally, the pressure inside the tube and the
temperature of the tube may be maintained above ambient levels for
a period of time to allow the polymer tube to be heat set. In one
embodiment, the temperature of the deformed tube may be maintained
at greater than or equal to the T.sub.g of the continuous phase
polymer and less than or equal to the T.sub.m of the continuous
polymer for a selected period to time. The selected period of time
may be between about one minute and about two hours, or more
narrowly, between about two minutes and about ten minutes. "Heat
setting" refers to allowing polymer chains to equilibrate to
different configurations in response to an elevated temperature. In
this case, the polymer chains are allowed to adopt highly oriented
structure at an elevated temperature. Polymer chain alignment is a
time and temperature dependent process, therefore, a period of time
may be necessary to allow polymer chains to realign at a given
temperature that are stable in a deformed state of a polymeric
material. Heat setting may also be facilitated by tension.
[0181] Further embodiments of the present invention relate to
stents composed primarily or completely of polymeric fibers coiled
or braided into a mesh tube or stent structure. A braided stent can
provide sufficient radial strength, however, the radial profile of
such a mesh, fiber structure can be higher than desirable. In
particular, the "net points," which refer to the points of overlap
of fibers, tend to increase the radial profile of a stent. Net
points 142 are illustrated in FIG. 9. Embodiments of methods
described herein allow fabrication of a fiber mesh stent with a
sufficiently small profile and sufficiently high radial
strength.
[0182] Certain embodiments of a method of fabricating a stent may
include making a tube or stent structure from at least two types of
fibers. In one embodiment, a first fiber may include a first
polymer and a second fiber may include a second polymer.
[0183] Other embodiments of a method of fabricating a stent may
include making a tube or stent structure from composite fibers
including the first polymer and the second polymer. In some
embodiments, the fibers of the tube may include an inner core
including the first polymer and an outer layer or covering
including the second polymer. In other embodiments, the outer
covering may be the first polymer and the inner core may be the
second polymer. In alternative embodiments, the fibers of the tube
may be a mixture or blend of the first polymer and the second
polymer.
[0184] In certain embodiments, the first polymer may have a
softening temperature (T.sub.s) lower than a softening temperature
of the second polymer. Also, the first polymer may have a T.sub.m
and a T.sub.g less than the T.sub.m and T.sub.g of the second
polymer.
[0185] For example, in a stent including two different types of
fibers, a first fiber may be made from poly (l-lactic acid) and a
second fiber may be made from 10:90 poly(l-lactide-co-glycolide)
(10% lactide, 90% glycolide). The poly (L-lactic acid) has a
melting temperature of 175.degree. C. and
poly(l-lactide-co-glycolide) has a melting temperature of
200.degree. C. Additionally, an exemplary composite fiber may be
made from 50% poly (l-lactic acid) and 10:90
poly(l-lactide-co-glycolide).
[0186] Various embodiments of fabricating a stent having two types
of fibers and/or fibers composed of a mixture of the first polymer
and the second polymer may include heating the tube to a
temperature range between a softening temperature of the first
polymer and the melting temperature of the first polymer. In one
embodiment, pressure is applied to the tube so as to flatten at
least some of the fibers of the tube to reduce the radial profile
of the tube. A heated fiber made of a first polymer may tend to
flatten as pressure is applied, reducing the profile of the tube.
Likewise, a heated fiber including the first polymer may also tend
to flatten. In particular, the applied pressure may reduce the
radial profile of the tube of at least some of the net points of
the fibers.
[0187] In some embodiments, the second polymer may be adapted to
provide high strength to the stent structure during heating,
flattening of the fibers, and during use of the stent. In an
embodiment, the temperature range may be below the T.sub.g of the
second polymer. In this case, the second polymer may remain
relatively rigid during heating a flattening of the first fiber.
The temperature range may also be below the T.sub.s of the second
polymer. Alternatively, the temperature range may be above the
T.sub.s of the second polymer.
[0188] In some embodiments, a cross-section of a fiber may be
circular. Alternatively, a cross-section of a fiber may have an
oblong shape, for example, an oval or elliptical shape. Fibers with
an oblong-shaped cross-section may allow greater surface coverage
of a vessel, in addition to providing a smaller radial profile to
the stent.
[0189] In one embodiment, the tube may be disposed over a mandrel
during heating and flattening of the fibers. Pressure may be
applied for flattening the fibers by a pressure tube disposed over
the tube. In an alternate embodiment, the fibers may be heated and
flattened in a heated crimper. Heat may be applied to the tube by a
heated mandrel. The tube may also be heated by blowing a heated
fluid onto the tube such as an inert gas, e.g., argon, air, oxygen,
nitrogen, etc. Additionally, the tube may be allowed to heat set on
the mandrel.
[0190] In an embodiment, the tube may be heated and the pressure
applied at or near a fabricated diameter of the tube. The tube may
be allowed to heat set by maintaining the tube in the temperature
range for a selected period of time.
[0191] In some embodiments, the method may include radial expansion
the tube prior to, during, or subsequent to heating and/or applying
pressure to flatten at least some of the fibers. As described
above, radial expansion may induce molecular orientation in the
fiber polymer that tends to increase the tensile strength of the
fiber. In other embodiments, the tube may be crimped prior to,
during, or subsequent to heating the tube and/or applying pressure
to flatten at least some of the fibers.
[0192] FIG. 11 depicts a radial cross-section of composite fiber
140. Fiber 140 has an inner core 142 of a first polymer and an
outer layer 144 of a second polymer. FIG. 12 depicts a radial
cross-section of a system 150 that can be used for heating and
flattening fibers of a fiber stent. Overlapping fibers 152 are
between a sliding wedge-type crimper 154 and a stationary mandrel
156. Wedge-type crimper 154 is heated and can apply pressure to the
fibers of the stent. A region 160 including fibers 152, crimper
154, and mandrel 156 in FIG. 12 is shown in an expanded view in
FIGS. 13 and 14. FIG. 13 illustrates fibers 152 prior to applying
pressure to the fiber by inward movement of crimper 154 and FIG. 14
shows fibers 152 after applying pressure with crimper 154. FIG. 14
shows that fibers 152 have been flattened by crimper 154.
[0193] As indicated above, there are difficulties associated with
manufacturing stents with small radiopaque markers. Various
embodiments of stents and methods of making stents that include
metallic films coupled and/or embedded within polymeric stents are
disclosed herein. The metallic film may be sufficiently radiopaque
to allow the stent to be visualized during use.
[0194] The stents may be formed by coupling and/or embedding
metallic film in and/or on a polymeric tube. A stent may be
fabricated by forming a pattern of interconnecting structural
elements in the tube with the metallic film using, for example, a
laser. Forming the pattern may include removal of some of the
metallic film in or on the polymer in addition to removal of
polymer.
[0195] In certain embodiments, the metallic film may include a
biostable metal, a bioerodible metal, or a combination of a
biostable and bioerodible metal. As indicated above, representative
examples of bioerodible metals that may be used to fabricate an
implantable medical device may include, but are not limited to,
magnesium, zinc, and iron. Representative examples of biostable
metals that may be used to fabricate an implantable medical device
may include, but are not limited to, gold or platinum.
[0196] In certain embodiments, a stent may include metallic film
coupled to a plurality of portions of a surface of the stent. Some
embodiments of a method of making the stent may include coupling a
metallic film to at least a portion of a surface of a polymeric
tube. In one embodiment, the metallic film may include a band
circumferentially aligned around a surface of the tube. A length of
a band along a longitudinal axis of the stent may be less than or
equal to the length of the tube. In another embodiment, the
metallic film includes a longitudinal strip longitudinally aligned
along the surface of the tube.
[0197] As an illustration, FIG. 15 depicts a polymeric tube 170
with a circumferentially aligned metallic band 174 coupled or
adhered to the surface of tube 170. In addition, FIG. 16 depicts a
polymeric tube 178 with a longitudinally aligned strip of metallic
film 182 coupled or adhered to the surface of tube 178.
[0198] In an alternate embodiment, a method of making a stent may
include coupling the metallic film to at least a portion of a
surface of a polymeric sheet. The sheet may then be rolled and
bonded to form a tube.
[0199] FIG. 17 depicts a portion of a stent 200 fabricated from
tube 170 in FIG. 15. Stent 200 has circumferentially aligned
metallic film markers 204 coupled to structural elements 208. Line
A--A corresponds to the longitudinal axis of the stent. FIG. 18
depicts a stent 220 fabricated from tube 178 in FIG. 16. Stent 220
has longitudinally aligned metallic markers 224 coupled to
structural elements 228. Line A--A corresponds to the longitudinal
axis of the stent.
[0200] In some embodiments, the metallic film may be coupled to the
polymeric tube using any suitable biocompatible adhesive. In one
embodiment, the adhesive may include a solvent. The solvent may
dissolve the polymer of the polymeric tube to allow the metal film
to be coupled to the tube. In another embodiment, the adhesive may
include a solvent mixed with a polymer. The solvent or the
solvent-polymer mixture may be applied to the tube followed by
application of the metallic film. The solvent may then be removed
by heating the tube, for example, in an oven.
[0201] Representative examples of solvents may include, but are not
limited to, chloroform, acetone, chlorobenzene, ethyl acetate,
1,4-dioxane, ethylene dichloride, 2-ethyhexanol, and combinations
thereof. Representative polymers may include biostable and
biodegradable polymers disclosed herein that may be dissolved by
the selected solvent.
[0202] In other embodiments, adhesives may include, but are not
limited to, thermosets such as, for example, epoxies, polyesters
and phenolics; thermoplastics such as, for example, polyamides,
polyesters and ethyl vinyl acetate (EVA) copolymers; and elastomers
such as, for example, natural rubber, styrene-isoprene-styrene
block copolymers, and polyisobutylene. Other adhesives include, but
are not limited to, proteins; cellulose; starch; poly(ethylene
glycol); fibrin glue; and derivatives and combinations thereof.
[0203] Mixtures of solvents and another substance can be used to
form adhesives. In some embodiments, mixtures of water and sugar
such as, for example, mixtures of water and sucrose, can be used as
an adhesive. In other embodiments, mixtures of PEG, or derivatives
thereof, can be mixed with a suitable solvent to form an adhesive.
Suitable solvents for PEG, or derivatives thereof, include, but are
not limited to, water, ethanol, chloroform, acetone, and the
like.
[0204] In other embodiments, the method may further include forming
a coating above an outer surface of the stent with a metallic film
coupled to a surface of the stent. The coating may be above at
least a portion of the metallic film on the surface of the stent.
In one embodiment, the coating may include a biostable or
biodegradable polymer. In one embodiment, the coating may include
an active agent or drug. The coating may be formed by applying a
mixture of a polymer and a solvent, followed by removal of the
solvent. In an embodiment, the polymer coating may inhibit or
prevent detachment of the metallic film from the stent prior to
substantial or complete biodegradation of a biodegradable
coating.
[0205] As an illustration, FIG. 19 depicts a cross-sectional view
of a sidewall of a portion 230 of a structural element of a stent.
A coating 232 is above a metallic film marker 234 which is coupled
or adhered to a polymeric substrate 238. Coating 232 tends to
inhibit detachment of marker 234 from substrate 238.
[0206] In other embodiments, structural elements of a stent may
include two radial polymeric layers with metallic film embedded in
a plurality of locations in between the layers. In certain
embodiments, a method of making the stent may include forming a
tube including a metallic film embedded between two polymer layers
and fabricating a stent from the tube. In an embodiment, the
metallic film may be a band circumferentially aligned around the
tube in between the polymeric layers. In another embodiment, the
metallic film may be a longitudinal strip longitudinally aligned
along the tube in between the polymeric layers.
[0207] Alternatively, a sheet may be formed including the metallic
film embedded between two polymer layers. The sheet may then be
rolled and bonded to form a tube.
[0208] As an illustration, FIG. 20 depicts a cross-sectional view
of a sidewall of a portion 240 of a structural element of a stent.
A metallic film marker 244 is embedded between a luminal polymeric
layer 248 and an abluminal polymeric layer 252.
[0209] Some embodiments may include forming the tube by extruding
an outer polymeric tubular layer over an inner tubular polymeric
layer with a metallic film disposed above a surface of the inner
layer. Extruding the outer layer over the inner layer may then
embed the metallic film between the layers. In some embodiments,
the melting temperature of an inner polymeric layer may be higher
than the melting temperature of the outer layer. The outer layer
may be extruded over the inner layer at a temperature above the
melting temperature of the outer polymer layer and below the
melting temperature of the inner polymer layer, allowing the inner
layer to maintain its structural integrity.
[0210] In further embodiments, a method of making a stent may
include elongating a polymeric tube so that a diameter of the stent
decreases. A metallic film in the form of a metallic band may then
be positioned around at least a portion of the elongated tube. The
polymeric tube with the metallic band positioned around the tube
may then be heated. In some embodiments, the method may further
include allowing the heated tube to radially expand so as to couple
the metallic band to the tube. The heated tube may radially expand
to at least a diameter of the metallic band.
[0211] Representative examples of polymers that may be used to
fabricate embodiments of implantable medical devices disclosed
herein include, but are not limited to, poly(N-acetylglucosamine)
(Chitin), Chitosan, poly(3-hydroxyvalerate),
poly(lactide-co-glycolide), poly(3-hydroxybutyrate),
poly(4-hydroxybutyrate),
poly(3-hydroxybutyrate-co-3-hydroxyvalerate), polyorthoester,
polyanhydride, poly(glycolic acid), poly(glycolide), poly(L-lactic
acid), poly(L-lactide), poly(D,L-lactic acid), poly(D,L-lactide),
poly(L-lactide-co-D,L-lactide), poly(caprolactone),
poly(L-lactide-co-caprolactone), poly(D,L-lactide-co-caprolactone),
poly(glycolide-co-caprolactone), poly(trimethylene carbonate),
polyester amide, poly(glycolic acid-co-trimethylene carbonate),
co-poly(ether-esters) (e.g. PEO/PLA), polyphosphazenes,
biomolecules (such as fibrin, fibrinogen, cellulose, starch,
collagen and hyaluronic acid), polyurethanes, silicones,
polyesters, polyolefins, polyisobutylene and ethylene-alphaolefin
copolymers, acrylic polymers and copolymers other than
polyacrylates, vinyl halide polymers and copolymers (such as
polyvinyl chloride), polyvinyl ethers (such as polyvinyl methyl
ether), polyvinylidene halides (such as polyvinylidene chloride),
polyacrylonitrile, polyvinyl ketones, polyvinyl aromatics (such as
polystyrene), polyvinyl esters (such as polyvinyl acetate),
acrylonitrile-styrene copolymers, ABS resins, polyamides (such as
Nylon 66 and polycaprolactam), polycarbonates, polyoxymethylenes,
polyimides, polyethers, polyurethanes, rayon, rayon-triacetate,
cellulose, cellulose acetate, cellulose butyrate, cellulose acetate
butyrate, cellophane, cellulose nitrate, cellulose propionate,
cellulose ethers, and carboxymethyl cellulose. Additional
representative examples of polymers that may be especially well
suited for use in fabricating embodiments of implantable medical
devices disclosed herein include ethylene vinyl alcohol copolymer
(commonly known by the generic name EVOH or by the trade name
EVAL), poly(butyl methacrylate), poly(vinylidene
fluoride-co-hexafluoropropene) (e.g., SOLEF 21508, available from
Solvay Solexis PVDF, Thorofare, N.J.), polyvinylidene fluoride
(otherwise known as KYNAR, available from ATOFINA Chemicals,
Philadelphia, Pa.), ethylene-vinyl acetate copolymers, poly(vinyl
acetate), styrene-isobutylene-styrene triblock copolymers, and
polyethylene glycol.
[0212] While particular embodiments of the present invention have
been shown and described, it will be obvious to those skilled in
the art that changes and modifications can be made without
departing from this invention in its broader aspects. Therefore,
the appended claims are to encompass within their scope all such
changes and modifications as fall within the true spirit and scope
of this invention.
* * * * *