U.S. patent application number 15/661925 was filed with the patent office on 2017-11-09 for high molecular weight polylactide and polycaprolactone copolymer and blends for bioresorbable vascular scaffolds.
The applicant listed for this patent is Abbott Cardiovascular Systems Inc.. Invention is credited to Manish Gada, Mary Beth Kossuth, Xiao Ma, James P. Oberhauser, Stephen D. Pacetti.
Application Number | 20170319363 15/661925 |
Document ID | / |
Family ID | 54251710 |
Filed Date | 2017-11-09 |
United States Patent
Application |
20170319363 |
Kind Code |
A1 |
Ma; Xiao ; et al. |
November 9, 2017 |
HIGH MOLECULAR WEIGHT POLYLACTIDE AND POLYCAPROLACTONE COPOLYMER
AND BLENDS FOR BIORESORBABLE VASCULAR SCAFFOLDS
Abstract
Bioresorbable polymer vascular scaffolds made of combinations of
polylactide and polycaprolactone having a high molecular weight
polymer, thin struts in a selected range and sufficient radial
strength to support a vessel upon deployment. The scaffolds have
degradation behavior of molecular weight, radial strength, and mass
that are conducive to healing of a vessel including providing
patency to a vessel, reduction of radial strength, breaking up, and
resorbing to allow return of the vessel to a natural state.
Inventors: |
Ma; Xiao; (Santa Clara,
CA) ; Kossuth; Mary Beth; (San Jose, CA) ;
Oberhauser; James P.; (Saratoga, CA) ; Pacetti;
Stephen D.; (San Jose, CA) ; Gada; Manish;
(Santa Clara, CA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Abbott Cardiovascular Systems Inc. |
Santa Clara |
CA |
US |
|
|
Family ID: |
54251710 |
Appl. No.: |
15/661925 |
Filed: |
July 27, 2017 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
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14307440 |
Jun 17, 2014 |
9750622 |
|
|
15661925 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
A61L 31/041 20130101;
A61L 31/06 20130101; A61F 2/91 20130101; A61F 2/90 20130101; A61L
31/148 20130101 |
International
Class: |
A61F 2/90 20130101
A61F002/90; A61L 31/06 20060101 A61L031/06; A61L 31/04 20060101
A61L031/04; A61L 31/14 20060101 A61L031/14; A61F 2/91 20130101
A61F002/91 |
Claims
1. (canceled)
2. A stent comprising: a bioresorbable polymer scaffold comprising
a polymer combination including a polylactide polymer and
polycaprolactone, wherein the polymer combination is a random
copolymer of poly(L-lactide)(PLLA) and polycaprolactone, wherein
the scaffold includes a plurality of interconnected struts, wherein
the scaffold is made of at least 95 wt % of the polymer
combination, wherein a thickess of the struts is 80 to 100 microns
and a width of the struts is 200 to 250 microns, and wherein the
scaffold has a crimped state and a deployed state and a radial
strength of the scaffold when expanded from the crimped state to
the deployed state in saline or bodily fluid at 37.degree. C. is at
least 650 mm Hg.
3. The stent of claim 2, wherein the polymer combination comprises
1 to 5 wt % of caprolactone units.
4. The stent of claim 2, wherein a hoop ultimate tensile strength
is at least 25% greater than an axial ultimate tensile strength of
the scaffold.
5. The stent of claim 2, wherein a concentration of unreacted
lactide monomer in the scaffold is 0.5 to 1 wt %.
7. The stent of claim 2, wherein a number average molecular weight
(Mn) of the polymer combination is less than 60 kDa at 1 year of
exposure of the scaffold to saline or bodily fluids at 37.degree.
C.
8. The stent of claim 2, wherein a crystallinity of the scaffold is
25% to 50%.
9. The stent of claim 2, wherein a change in retardance as measured
by polarized light microscopy (PLM) from an inner diameter to 50%
of the thickness to the outer diameter of the scaffold is less than
50%.
10. The stent of claim 2, wherein a number average molecular weight
(Mn) of a polymer of the polymer combination is greater than 110
kDa.
11. A stent comprising: a bioresorbable polymer scaffold comprising
a polymer formulation including a random copolymer of
poly(L-lactide) (PLLA) polymer and polycaprolactone (PCL), wherein
the scaffold includes a plurality of interconnected struts, wherein
a thickess of the struts is 80 to 100 microns and a width of the
struts is 200 to 250 microns, wherein a number average molecular
weight (Mn) of the blend is greater than 60 kDa, and wherein the
scaffold has a crimped state and a deployed state and a radial
strength of the scaffold when expanded from the crimped state to
the deployed state in saline or bodily fluid at 37.degree. C. is at
least 650 mm Hg.
12. The stent of claim 11, wherein the polymer formulation
comprises 1 to 5 wt % of caprolactone units.
Description
[0001] This application is a continuation of U.S. application Ser.
No. 14/307,440 filed Jun. 17, 2014 which is incorporated herein by
reference.
BACKGROUND OF THE INVENTION
Field of the Invention
[0002] This invention relates polymeric medical devices, in
particular, bioresorbable stents or stent scaffoldings
Description of the State of the Art
[0003] This invention relates to radially expandable endoprostheses
that 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. A stent is an example of such an
endoprosthesis. Stents are generally cylindrically shaped devices
that 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 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.
[0004] Stents are typically composed of a scaffold or scaffolding
that includes a pattern or network of interconnecting structural
elements or struts, formed from wires, tubes, or sheets of material
rolled into a cylindrical shape. This scaffolding gets its name
because it possibly physically holds open and, if desired, expands
the wall of the passageway. Typically, stents are capable of being
compressed or crimped onto a catheter so that they can be delivered
to and deployed at a treatment site.
[0005] Delivery includes inserting the stent through small lumens
using a catheter and transporting it to the treatment site.
Deployment includes expanding the stent to a larger diameter once
it is at the desired location. Mechanical intervention with stents
has reduced the rate of restenosis as compared to balloon
angioplasty. Yet, restenosis remains a significant problem. When
restenosis does occur in the stented segment, its treatment can be
challenging, as clinical options are more limited than for those
lesions that were treated solely with a balloon.
[0006] Stents are used not only for mechanical intervention but
also as vehicles for providing biological therapy. Biological
therapy uses medicated stents to locally administer a therapeutic
substance. A medicated stent may be fabricated by coating the
surface of either a metallic or polymeric scaffold with a polymeric
carrier that includes an active or bioactive agent or drug.
Polymeric scaffolds may also serve as a carrier of an active agent
or drug. An active agent or drug may also be included on a scaffold
without being incorporated into a polymeric carrier.
[0007] Stents are generally made to withstand the structural loads,
namely radial compressive forces, imposed on the scaffold as it
supports the walls of a vessel. Therefore, a stent must possess
adequate radial strength if its function is to support a vessel at
an increased diameter. Radial strength, which is the ability of a
stent to resist radial compressive forces, relates to a stent's
radial yield strength and radial stiffness around a circumferential
direction of the stent. A stent's "radial yield strength" or
"radial strength" (for purposes of this application) may be
understood as the compressive loading or pressure, which if
exceeded, creates a yield stress condition resulting in the stent
diameter not returning to its unloaded diameter, i.e., there is
irrecoverable deformation of the stent. See, T. W. Duerig et al.,
Min Invas Ther & Allied Technol 2000: 9(3/4) 235-246. Stiffness
is a measure of the elastic response of a device to an applied load
and thus will reflect the effectiveness of the stent in resisting
diameter loss due to vessel recoil and other mechanical events.
Radial stiffness can be defined for a tubular device such as stent
as the hoop force per unit length (of the device) required to
elastically change its diameter. The inverse or reciprocal of
radial stiffness may be referred to as the compliance. See, T. W.
Duerig et al., Min Invas Ther & Allied Technol 2000: 9(3/4)
235-246.
[0008] When the radial yield strength is exceeded, the stent is
expected to yield more severely and only a minimal force is
required to cause major deformation. Radial strength is measured
either by applying a compressive load to a stent between flat
plates or by applying an inwardly-directed radial load to the
stent.
[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. In addition, the stent must
possess sufficient flexibility to allow for crimping, expansion,
and cyclic loading.
[0010] Some treatments with stents require its presence for only a
limited period of time. Once treatment is complete, which may
include structural tissue support and/or drug delivery, it may be
desirable for the stent to be removed or disappear from the
treatment location. One way of having a stent disappear may be by
fabricating a stent in whole or in part from materials that erode
or disintegrate through exposure to conditions within the body.
Stents fabricated from biodegradable, bioabsorbable, bioresorbable,
and/or bioerodable materials such as bioabsorbable polymers can be
designed to completely erode only after the clinical need for them
has ended.
[0011] In addition to high radial strength, a vascular scaffold
must have sufficient resistance to fracture or sufficient
toughness. A vascular scaffold is subjected to a large deformation
during use, in particular, when it is crimped to a delivery
diameter and when it is deployed. A scaffold may be susceptible to
fracture when in use which can negatively impact performance and
even lead to device failure. Fabricating a polymer-based scaffold
that has sufficiently high radial strength as well as resistance to
fracture is a challenge.
[0012] It is advantageous for vascular scaffolds to have thin
struts while maintaining adequate radial strength. Thin struts lead
to a lower profile device in the crimped state for better
deliverability. After implantation, neointima proliferates until
stent struts are covered. Consequently, thinner struts have less
neointimal formation and less area obstruction of the vessel.
Lastly, thin struts disturb blood flow less and are less
thrombogenic. However, polymer based materials can be orders of
magnitude lower in strength in terms of ultimate strength and
stiffness compared to metallic alloys. Fabricating a polymer-based
scaffold that has sufficiently high radial strength at strut
thicknesses comparable to current metallic stents is therefore a
challenge.
[0013] Additionally, treating peripheral vascular disease
percutaneously in the lower limbs is a challenge with current
technologies. Long term results are sub-optimal due to chronic
injury caused by the constant motions of the vessel and the implant
as part of everyday life situations. To reduce the chronic injury,
a bioresorbable scaffold for the superficial femoral artery (SFA)
and/or the popliteal artery can be used so that the scaffold
disappears before it causes any significant long term damage.
However, one of the challenges with the development of a femoral
scaffold and especially a longer length scaffold (4-25 cm) to be
exposed to the distal femoral artery and potentially the popliteal
artery is the presence of fatigue motions that may lead to chronic
recoil and strut fractures especially in the superficial femoral
artery, prior to the intended bioresorption time especially when
implanted in the superficial femoral artery.
[0014] Fabricating a polymer-based scaffold for treating the SFA is
even more challenging than for coronary applications. A scaffold in
the SFA and/or the popliteal artery is subjected to various
non-pulsatile forces, such as radial compression, torsion, flexion,
and axial extension and compression. These forces place a high
demand on the scaffold mechanical performance and can make the
scaffold more susceptible to fracture than less demanding
anatomies. Stents or scaffolds for peripheral vessels such as the
SFA, require a high degree of crush recovery. The term "crush
recovery" is used to describe how the scaffold recovers from a
pinch or crush load, while the term "crush resistance" is used to
describe the force required to cause a permanent deformation of a
scaffold. It has been believed that a requirement of a stent for
SFA treatment is a radial strength high enough to maintain a vessel
at an expanded diameter. A stent which combines such high radial
strength, high crush recovery, and high resistance to fracture is a
great challenge.
INCORPORATION BY REFERENCE
[0015] All publications, patents, and patent applications mentioned
in this specification are herein incorporated by reference to the
same extent as if each individual publication, patent, or patent
application was specifically and individually indicated to be
incorporated by reference, and as if each said individual
publication, patent, or patent application was fully set forth,
including any figures, herein.
SUMMARY OF THE INVENTION
[0016] A first set of embodiments of the present invention includes
a stent comprising: a bioresorbable polymer scaffold comprising a
polymer combination including a polylactide polymer and
polycaprolactone, wherein scaffold includes a plurality of
interconnected struts and a thickess of the struts is less than 120
microns, wherein a number average molecular weight (Mn) of the
polymer combination or a polymer of the polymer combination is
greater than 110 kDa, and wherein the scaffold has a crimped state
and a deployed state and a radial strength of the scaffold when
expanded from the crimped state to the deployed state in saline or
bodily fluid at 37.degree. C. is at least 350 mm Hg.
[0017] The first set of embodiments may have one or more, or any
combination of the following aspects (1) to (6): (1) wherein the
polymer combination comprises a random copolymer of PLA and PCL
random copolymer comprising a 1 to 5 mol % of caprolactone units;
(2) wherein the polymer combination comprises a block copolymer of
PLA polymer blocks and PCL polymer blocks including 1 to 5 wt % of
PCL blocks; (3) wherein the polymer combination comprises a blend
of a PLA homopolymer with PCL homopolymer; (4) wherein the PCL
homopolymer is 1 to 5 wt % of the blend; (5) wherein the polymer
combination comprises a blend of PLA polymer and a PLA and PCL
copolymer; (6) wherein caprolactone units of the copolymer are 1 to
5 wt % of the blend.
[0018] A second set of embodiments of the present invention
includes a stent comprising: a bioresorbable polymer scaffold
comprising polymer formulation including a blend of PLA polymer and
a PLA and PCL copolymer, wherein the scaffold includes a plurality
of interconnected struts and a thickess of the struts is less than
120 microns, wherein a number average molecular weight (Mn) of the
blend is greater than 60 kDa, and wherein the scaffold has a
crimped state and a deployed state and a radial strength of the
scaffold when expanded from the crimped state to the deployed state
in saline or bodily fluid at 37.degree. C. is at least 350 mm
Hg.
[0019] The second set of embodiments may have one or more, or any
combination of the following aspects (1) to (2): (1) wherein
caprolactone units are 1 to 5 wt % of the formulation; (2) wherein
the Mn of the blend is 100 to 250 kDa.
[0020] A third set of embodiments of the present invention includes
a method of fabricating a stent including a bioresorbable scaffold,
comprising: providing a polylactide (PLA) polymer resin having an
intrinsic viscosity of 5 to 8 dL/g and a PLA and polycaprolactone
(PCL) copolymer resin; forming a tube by melt processing the PLA
resin and the copolymer comprising a blend of the PLA polymer and
the copolymer; processing the formed tube to increase the
crystallinity to at least 20%; and forming a scaffold from the
processed tube comprising a plurality of struts having a thickness
of less than 120 microns.
[0021] The third set of embodiments may have one or more, or any
combination of the following aspects (1) to (3): (1) wherein the
processing comprises radially expanding the formed tube to an
expanded diameter and forming the scaffold from the tube at the
expanded diameter; (2) wherein a percent radial expansion is at
least 400%; (3) further comprising adding unreacted lactide monomer
to the PLA polymer resin and the copolymer resin during the melt
processing, wherein the scaffold comprises at least 0.5 wt %
unreacted monomer content.
[0022] A fourth set of embodiments of the present invention
includes a method of fabricating a stent including a bioresorbable
scaffold, comprising: providing a tube comprising a blend of a PLA
polymer and PLA and PCL copolymer formed from melt processing a PLA
polymer resin having an intrinsic viscosity of 5 to 8 dL/g and a
PLA/PCL copolymer resin; radially expanding the tube at least by
400%; and forming a scaffold from the expanded tube comprising a
plurality of struts having a thickness of less than 120 microns,
wherein the scaffold has a crimped state and a deployed state and a
radial strength of the scaffold when expanded from the crimped
state to the deployed state in saline or bodily fluid at 37.degree.
C. is at least 350 mm Hg.
[0023] The fourth set of embodiments may have one or more, or any
combination of the following aspects (1) to (3): (1) wherein an Mn
of the blend after sterlization of the scaffold is 100 kDa to 250
kDa; (2) wherein a size of a majority of the crystalline domains in
the scaffold are 10 nm to 50 nm; (3) wherein a change in retardance
as measured by polarized light microscopy (PLM) from an inner
diameter to 50% of the thickness to the outer diameter of the
scaffold is less than 50%.
BRIEF DESCRIPTION OF THE DRAWINGS
[0024] FIG. 1A depicts a view of an exemplary scaffold.
[0025] FIG. 1B show a cross-selection of a strut of the scaffold of
FIG. 1A.
[0026] FIG. 2 depicts the radial strength dependence on strut
thickness of a scaffold made from poly(L-lactide).
[0027] FIG. 3 depicts an embodiment of a scaffold pattern.
[0028] FIG. 4 depicts another embodiment of a scaffold pattern.
DETAILED DESCRIPTION OF THE INVENTION
[0029] In many treatment applications using stents, stents expand
and hold open narrowed portions of blood vessels. As indicated, to
achieve this, the stent must possess a radial strength in an
expanded state that is sufficiently high and sustainable to
maintain the expanded vessel size for a period of weeks or months.
This generally requires a high strength and rigid material. In the
case of bioresorbable polymer stents or scaffolds, bioresorbable
polymers that are stiff and rigid have been proposed and used in
stents for coronary intervention. Such polymers are stiff or rigid
under physiological conditions within a human body. These polymers
tend to be semicrystalline polymers that have a glass transition
temperature (Tg) in a dry state sufficiently above human body
temperature (approximately 37.degree. C.) that the polymer is stiff
or rigid at these conditions. Polylactide and polylactide based
polymers such as poly(L-lactide) are examples of such
semicrystalline polymers that have been proposed and used as a
stent or scaffold materials.
[0030] Fabricating a vascular scaffold from such materials with
sufficient fracture toughness or fracture resistance is challenging
due to their brittle nature. Vascular scaffolds are subjected to
deformation and stress during manufacture when crimped to a
delivery diameter, when deployed or expanded from a delivery
diameter to a deployment diameter, and during use after deployment.
As a result, vascular scaffolds are susceptible to fracture during
manufacture (particularly during crimping), deployment, and use.
The fracture toughness is important in reducing material-level
damage during crimping and in vitro/in vivo deployment of a
bioresorbable scaffold. The reduced damage allows achievement of a
sufficiently high radial strength with a reduced strut thickness
and cross-section.
[0031] It is a continuing challenge to develop new materials and
processing methods for vascular scaffolds that improve the
resistance to fracture with sufficiently high radial strength,
particularly during crimping and deployment or expansion.
[0032] Another challenge in making a bioabsorbable polymer scaffold
relates to the lower strength to weight ratio of polymers compared
to metals. The strength of a scaffold material is proportional to
the radial strength of the scaffold. Therefore, polymeric scaffolds
require thicker struts than a metallic stent to achieve the radial
strength required to provide patency to a blood vessel. Exemplary
coronary polymer scaffolds have wall thicknesses from about 150 to
170 microns while coronary metallic stents have strut thicknesses
of 60 to 100 microns. It is desirable to have a scaffold profile as
low as possible. Thus, making a scaffold with a smaller form
factor, i.e., with thinner struts, that provides sufficient radial
strength is a challenge.
[0033] FIG. 1A depicts a view of an exemplary scaffold 100 which
includes a pattern or network of interconnecting structural
elements 105. FIG. 1A illustrates features that are typical to many
stent patterns including cylindrical rings 107 connected by linking
elements 110. The cylindrical rings are load bearing in that they
provide radially directed force in response to an inward force on
the scaffold. The linking elements generally function to hold the
cylindrical rings together. Exemplary scaffolds are disclosed in US
2008/0275537, US 2011/0190872, and US 2011/0190871. Any of the
patterns disclosed in these references are applicable to the
inventive scaffolds.
[0034] FIG. 1B show a cross-selection of a strut 2 showing the
polymer scaffold body, polymer backbone, or core of the strut
surrounded by a drug/polymer coating or matrix 16. The
cross-section of the strut has an abluminal or outer surface or
side 12 that faces the vessel wall and a luminal or inner surface
or side 14 that faces the lumen of the vessel. The strut
cross-section shown is to be rectangular with a width (W) and
thickness (T). The scaffold cross-section may be rectangular or
approximately rectangular. The slight curvature at the inner and
outer surfaces due to the tubular geometry is not shown. The
present invention is not limited to this scaffold pattern or type
of pattern and is applicable to any pattern.
[0035] The challenge of obtaining a sufficiently high deployed
scaffold radial strength for a polymer scaffold having
significantly thinner struts than the 150 to 170 micron range is
shown by FIG. 2. FIG. 2 depicts the radial strength dependence on
strut thickness of a scaffold made from PLLA. The number average
molecular weight (Mn) of the PLLA of the scaffold is less than 100
kg/mol. FIG. 2 shows the radial strength for two scaffolds of the
same design with different strut thicknesses, 115 microns and 150
microns. The two scaffolds were in a crimped or reduced profile of
about 0.055 in and then deployed in saline solution at 37.degree.
C. with a balloon to about 3.5 mm OD. The radial strength was
measured by the MSI RX550 Radial Force Tester obtained from MSI of
Flagstaff, Ariz. As shown, the scaffold with 150 micron thickness
has a radial strength of about 1173 mm Hg and the scaffold with 115
micron thickness has a radial strength of about 650 mm Hg, showing
the strong dependence of radial strength on dimensions.
[0036] In addition to resistance to fracture and reduced form
factor, vascular scaffolds should possess degradation behavior that
is favorable to treatment of vascular lesions. The degradation
behavior refers to the temporal degradation profile of molecular
weight, radial strength, and mass. Upon implantation, a
bioresorbable scaffold should maintain its radial strength for a
period of months to provide patency to the vessel while the vessel
wall heals at the increased diameter. The desired minimum radial
strength for coronary applications is 350 mm Hg. In addition,
neointima grows over the scaffold which eventually covers all or
most of the scaffold. After about three to six months the radial
strength decreases significantly followed by breaking up of the
scaffold and resorption of the scaffold material. This allows the
vessel to regain a healthy unrestricted natural state which
includes further expansion and resumption of vasomotion. The
scaffold should completely resorb from the vessel within 18 to 36
months.
[0037] Embodiments of the present invention are directed to
implantable medical devices such as bioresorbable vascular
scaffolds including a high molecular weight polymer having thin
struts in a selected range and sufficient radial strength to
support a vessel upon deployment. The inventive scaffolds may
further have degradation behavior of molecular weight, radial
strength, and mass that are conducive to healing of a vessel, as
described herein, including providing patency to a vessel,
reduction of radial strength, breaking up, and resorbing to allow
return of the vessel to a natural state.
[0038] Selected ranges of stmt thickness include less than 150
microns, less than 140 microns, less than 130 microns, about 100
micron, 80 to 100 microns, 80 to 120 microns, 90 to 100 microns, 90
to 110 microns, 110 to 120 microns, or 95 to 105 microns. The
thickness may refer to a thickness of a scaffold that is formed by
laser cutting a tube. The thickness may further refer to the
thickness of the scaffold formed from laser cutting plus a
thickness of a coating over the laser cut scaffold. All or a
majority of the struts of the scaffold may have a thickness in the
selected range. An aspect ratio of strut width divided by strut
thickness may be defined. Selected ranges of this aspect ratio
include less than 3, less than 2, less than 1, less than 0.5, 0.75
to 2, or 0.9 to 1.5.
[0039] The radial strength of the scaffold can be high enough to
provide mechanical support to a vessel after expanding the vessel
to an increased diameter or prevent or reduce a decrease in the
diameter of the vessel. The scaffold has a crimped state and a
deployed state and a radial strength of the scaffold may refer to a
radial strength when expanded from the crimped state to the
deployed state in saline or bodily fluid at 37.degree. C. The
radial strength may be at least the value required to support a
vessel at a reference vessel diameter, which is the healthy
diameter of a vessel at an implant site. The radial strength is at
least 350 mm Hg, at least 500 mm Hg, at least 650 mm Hg, at least
800 mm Hg, at least 1000 mm Hg, 400 to 600 mm Hg, 500 to 1200 mm
Hg, 700 to 900 mm Hg, or 800 to 1300 mm Hg.
[0040] The high molecular weight polymer, the polymer formulation
of the scaffold, and processing to modify morphology combine to
provide sufficiently high radial strength for the thin stmt
scaffold. The vascular devices may further be resistant to fracture
when crimped to a reduced diameter and when expanded to a
deployment diameter, which also helps provide the high radial
strength. The polymer formulation includes a polylactide polymer
component and a polycaprolactone component as a homopolymer,
blocks, or as part of a random copolymer. The high molecular weight
is provided by starting with a polymer resin having an intrinsic
viscosity (IV) of 4 to 8 dL/g and processing that results in a
finished product number average molecular weight (Mn) of 70 to 250
kDa or 100 to 250 kDa. Finished product may refer to the stent
after sterilization.
[0041] The degradation behavior may be characterized in terms of
the time dependent molecular weight. The molecular weight (Mn) of
the high molecular weight polymer may be less than 100 kDa, 90 kDa,
80 kDa, 70 kDa, 60 kDa, 60 to 100 kDa, 60 to 80 kDa, or 80 to 100
kDa at 1 year of exposure of the scaffold to saline or bodily
fluids at 37.degree. C.
[0042] Embodiments of the inventive scaffold having features
described above include a scaffold material including formulations
or combinations of polylactide (PLA) polymers and polycaprolactone.
A polylactide polymer is one which contains L-lactide or L-lactic
acid in the polymer backbone and may optionally have other
bioresorbable monomers. The polymer combination includes a polymer
having a high molecular, as defined herein. The polycaprolactone
component and the high molecular weight help provide sufficient
radial strength and high fracture toughness of the scaffold.
[0043] The polymer combinations can include a blend, a random
copolymer, or a block copolymer, of a PLA polymer and
polycaprolactone (PCL). The stent body, scaffold, or substrate made
partially or completely made of polymer combination. The stent body
may also include a coating that includes a therapeutic agent.
[0044] The polymer combinations include: (1) PLA and PCL random
copolymer; (2) block copolymer including PLA polymer blocks and PCL
polymer blocks; (3) a blend of a PLA polymer with PCL homopolymer;
(4) blend of a PLA homopolymer blended and a PLA and PCL copolymer;
and (5) a blend of a PCL homopolymer and a PLA and PCL
copolymer.
[0045] Embodiments of the invention include a scaffold made
substantially or completely of the polymer combination.
"Substantially" may correspondent to greater than 90 wt %, greater
than 95 wt %, or greater than 99 wt %. The scaffold may have a
composition of 90 to 95% or 95 to 99% of the polymer combination.
The scaffold may include other components that include, but are not
limited to, fillers, plasticizers, visualization materials (e.g.,
radiopaque), or therapeutic agents.
[0046] The PLA polymer of the combination may include
poly(L-lactide) (PLLA), poly(D,L-lactide) having a constitutional
unit weight-to-weight (wt/wt) ratio of about 96/4,
poly(lactide-co-glycolide), poly(L-lactide-co-glycolide),
poly(D,L-lactide-co-glycolide), poly(D,L-lactide) made from
meso-lactide, and poly(D,L-lactide) made from polymerization of a
racemic mixture of L- and D-lactides. A PLA polymer can include a
PLA with a D-lactide content greater than 0 mol % and less than 15
mol %, or more narrowly, 1 to 15 mol %, 1 to 5 mol %, 5 to 10%, or
10 to 15 mol %. The PLA polymer includes poly(D,L-lactide) having a
constitutional unit weight-to-weight (wt/wt) ratio of about 93/7,
about 94/6, about 95/5, about 96/4, about 97/3, about 98/2, or
about 99/1. The term "unit" or "constitutional unit" refers to the
composition of a monomer as it appears in a polymer.
[0047] Embodiments of the invention include a scaffold including a
PLA and PCL random copolymer. The scaffold may be made
substantially or completely of the copolymer. The copolymer may
include poly(L-lactide-co-caprolactone),
poly(D,L-lactide-co-caprolactone),
poly(L-lactide-co-glycolide-co-caprolactone), and
poly(DL-lactide-co-glycolide-co-caprolactone). The copolymer with
D,L-lactide may be made from a racemic mixture of L- and D-lactide
or may include 1 to 15% of D constitutional units. The scaffold may
be made substantially or completely of the copolymer. In some
embodiments, the scaffold may include no PLA homopolymer, PCL
homopolymer, or less than 20%, 10%, 5%, or less than 1% of either
homopolymer.
[0048] The copolymer may include 1 to 5% (wt % or mol %) of
caprolactone units, or more narrowly, 1 to 2%, 2 to 5%, 3 to 5%, or
about 3%. The scaffold may be made from a copolymer resin with an
IV greater than 5 dL/g, greater than 7 dL/g greater than 8 dl/g,
3.8 to 8 Dl/g, 4 to 8 dL/g, 4 to 6 dL/g, 6 to 8 dL/g, or 5 to 7
dL/g. The Mn of the copolymer in a finished scaffold may be 100 to
250 kDa.
[0049] The Tm of copolymer resin or copolymer of the scaffold may
be 165.degree. C. The Tg of the copolymer may be 60 to 65.degree.
C.
[0050] The crystallinity of the copolymer or scaffold made of the
copolymer may be 20 to 50%, 20 to 25%, 25 to 30%, 30 to 35%, 35 to
40%, 40 to 45%, and 45 to 50%.
[0051] Embodiments of the invention include a scaffold including a
block copolymer including PLA polymer blocks and PCL polymer
blocks. The scaffold may be made substantially or completely of the
block copolymer. The block copolymer may be a linear block
copolymer or branched block copolymer such as a star block
copolymer.
[0052] The scaffold may include no PLA homopolymer, PCL
homopolymer, or less than 20%, 10%, 5%, or less than 1% of either
homopolymer. The PLA blocks may include PLLA,
poly(L-lactide-co-glycolide), poly(D,L-lactide-co-glycolide), and
poly(D,L-lactide). Blocks with D,L-lactide may be made from a
racemic mixture of L- and D-lactide or may include 1 to 15% of D
constitutional units. The scaffold may be made substantially or
completely of the block copolymer.
[0053] The block copolymer may include 1 to 5% (wt % or mol %) of
polycaprolactone blocks, or more narrowly, 1 to 2%, 2 to 5%, 3 to
5%, or about 3%. The scaffold may be made from a copolymer resin
with an IV greater than 5 dL/g, greater than 7 dL/g greater than 8
dl/g, 4 to 8 dL/g, 4 to 6 dL/g, 6 to 8 dL/g, or 5 to 7 dL/g. The Mn
of the block copolymer in a finished scaffold may be 100 to 250
kDa.
[0054] The Tm of copolymer resin or copolymer of the scaffold may
be 60 and 150 to 185.degree. C. for PCL and PLA block,
respectively. The Tg of the copolymer may be -60 and 60 to
75.degree. C. for PCL and PLA block, respectively. The
crystallinity of the copolymer or scaffold made of the copolymer
may be 20 to 50%, 20 to 25%, 25 to 30%, 30 to 35%, 35 to 40%, 40 to
45%, and 45 to 50%.
[0055] Embodiments of the invention include a scaffold including a
blend of a PLA polymer with a PCL homopolymer. The scaffold may be
made substantially or completely of the blend. The PLA polymer may
include PLLA, poly(L-lactide-co-glycolide),
poly(D,L-lactide-co-glycolide), and poly(D,L-lactide). PLA polymers
with D,L-lactide may be made from a racemic mixture of L- and
D-lactide or may include 1 to 15% of D constitutional units. The
scaffold may be made substantially or completely of the block
copolymer.
[0056] The blend may include 1 to 5% (wt % or mol %) of PCL
homopolymer, or more narrowly, 1 to 2%, 2 to 5%, 3 to 5%, or about
3%. The scaffold may be made from a PLA resin or resin blend with
an IV greater than 5 dL/g, greater than 7 dL/g greater than 8 dl/g,
4 to 8 dL/g, 4 to 6 dL/g, 6 to 8 dL/g, or 5 to 7 dL/g.
[0057] The Mn of the blend in the finished scaffold may be 100 to
250 kDa.
[0058] The Tm of the blend may be 150 to 185.degree. C. There may
or may not be a Tm of 60.degree. C. that is attributed to the PCL
homopolymer. The Tg of the blend may be 60 to 75.degree. C. There
may or may not be a Tg of -60.degree. C. that is attributed to the
PCL homopolymer.
[0059] The crystallinity of the blend or scaffold made of the blend
may be 20 to 50%, 20 to 25%, 25 to 30%, 30 to 35%, 35 to 40%, 40 to
45%, and 45 to 50%.
[0060] Embodiments of the invention include a scaffold including a
blend of a PLA polymer with a PLA and PCL copolymer. The scaffold
may be made substantially or completely of the blend. The PLA
polymer may include PLLA, poly(L-lactide-co-glycolide),
poly(D,L-lactide-co-glycolide), and poly(D,L-lactide). PLA polymers
with D,L-lactide may be made from a racemic mixture of L- and
D-lactide or may include 1 to 15% of D constitutional units.
[0061] The copolymer may be PLA and PCL random copolymer or a block
copolymer of PLA polymer blocks and PCL homopolymer blocks. The
random copolymer may include any from the list of PLA and PCL
random copolymers provided above. The block copolymer may be linear
block copolymer or branched block copolymer such as a star block
copolymer. The scaffold may be made substantially or completely of
the blend.
[0062] The scaffold may be made from a PLA resin or the resin blend
with an IV greater than 5 dL/g, greater than 7 dL/g greater than 8
dl/g, 4 to 8 dL/g, 4 to 6 dL/g, 6 to 8 dL/g, or 5 to 7 dL/g. The Mn
of the PLA polymer in a finished scaffold may be 150 to 300 kDa.
The Mn of the copolymer in a finished scaffold may be 100 to 250
kDa. The Mn of the blend in the finished scaffold may be 100 to 250
kDa.
[0063] The caprolactone units in either the random or block
copolymer may be 1 to 5% (wt % or mol %) of the blend, or more
narrowly, 1 to 2%, 2 to 5%, 3 to 5%, or about 3% of the blend. The
random copolymer may be 1% to 50% caprolactone units. Exemplary
random copolymers include 95/5 poly(L-lactide-co-caprolactone),
wherein 95/5 refers to 95 mol % L-lactide and 5% caprolactone, and
70/30 poly(L-lactide-co-caprolactone), wherein 70/30 refers to 70
mol % L-lactide and 30 mol % caprolactone. The IV of the copolymer
resin used may be 1.5 g/dL, 3.8 g/dL, or higher.
[0064] The Tm of the blend may be 160 to 185.degree. C. The Tg of
the blend may be 60 to 75.degree. C., and greater than 37.degree.
C. when hydrated.
[0065] The crystallinity of the blend or scaffold made of the blend
may be 20 to 50%, 20 to 25%, 25 to 30%, 30 to 35%, 35 to 40%, 40 to
45%, and 45 to 50%.
[0066] Embodiments of the invention include a scaffold including a
blend of a PCL homopolymer with a PLA and PCL copolymer. The
scaffold may be made substantially or completely of the blend. The
copolymer may be PLA and PCL random copolymer or a block copolymer
of PLA polymer blocks and PCL homopolymer blocks. The random
copolymer may include any from the list of PLA and PCL random
copolymers provided above. The block copolymer may include any from
the list of PLA and PCL block copolymers provided above. The block
copolymer may be linear block copolymer or branched block copolymer
such as a star block copolymer. The scaffold may be made
substantially or completely of the blend.
[0067] The scaffold may be made from a copolymer resin or the resin
blend with an IV greater than 5 dL/g, greater than 7 dL/g greater
than 8 dl/g, 4 to 8 dL/g, 4 to 6 dL/g, 6 to 8 dL/g, or 5 to 7 dL/g.
The Mn of the blend polymer in a finished scaffold may be 100 to
250 kDa. The total caprolactone units in both the copolymer and the
PCL may be 1 to 5% (wt % or mol %) of the blend, or more narrowly,
1 to 2%, 2 to 5%, 3 to 5%, or about 3% of the blend. The PCL
homopolymer may be 0.5% to 4% of the blend. The caprolactone
content of the copolymer may be 0.5% to 4%.
[0068] Exemplary random copolymers include 95/5
poly(L-lactide-co-caprolactone), wherein 95/5 refers to 95 mol %
L-lactide and 5% caprolactone and 70/30
poly(L-lactide-co-caprolactone), where 70/30 refers to 70 mol %
L-lactide and 30% caprolactone. The IV of the copolymer resin used
may be 1.5 dL/g, 3.8 dL/g, or higher.
[0069] The Tm of the blend may be 160 to 185.degree. C. The Tg of
the blend may be 60 to 75.degree. C.
[0070] The crystallinity of the blend or scaffold made of the blend
may be 20 to 50%, 20 to 25%, 25 to 30%, 30 to 35%, 35 to 40%, 40 to
45%, and 45 to 50%.
[0071] The inventive scaffolds may further have degradation
behavior of molecular weight, radial strength, and mass that are
conducive to and promote healing of a vessel, as described herein.
The higher molecular weight of the inventive scaffolds, as compared
to scaffolds with Mn less than about 100 kDa, will result in
degradation times of molecular weight, radial strength and mass
that are the same or not significantly higher. It has been reported
in the literature that in PLLA/PCL combinations, the PCL
accelerates degradation, as compared to PLLA without PCL. For
example, Tsuji, H. et al., Journal of Applied Polymer Science 67,
405-415 (1998) have reported that adding PCL into PLLA system
accelerates degradation as long as the PCL molecular weight is
lower than that of PLLA and the amount of PCL is less than 50%. In
the present invention, the Mn of the PCL in blends of PLLA and PCL
is less than the Mn of PLLA.
[0072] It has also been shown that adding PCL to a PLLA system
accelerates degradation as long as the following conditions are
met: 1) the PCL molecular weight is lower than that of PLLA and 2)
the amount of PCL is less than 50%. (Tsuji, H., Ikada, Y, Journal
of Applied Polymer Science 67, 405-415 (1998)) It has also been
shown that a PLLA/PCL 95/5 copolymer degrades twice as fast as
PLLA.
[0073] Therefore, utilizing a PLLA/PCL resin with an IV of 5 to 7
dL/g will not cause a significantly slower degradation rate
compared to a PLLA scaffold made from a resin with an IV of 3.8
dL/g or less.
[0074] Thus, in general, the molecular weight and amount of PCL can
be adjusted to obtain the degradation properties disclosed
herein.
[0075] The scaffold materials disclosed may be used with a variety
of scaffold patterns. FIG. 3 depicts a first embodiment of a
pattern 200 which includes longitudinally-spaced rings 212 formed
by struts 230. The pattern 200 of FIG. 3, represents a tubular
scaffold structure (for example, as shown in FIG. 1), so that an
axis A-A is parallel to the central or longitudinal axis of the
scaffold. The scaffold structures shown may be in a state prior to
crimping or after deployment.
[0076] In FIG. 3, a ring 212 is connected to an adjacent ring by
several links 234, each of which extends parallel to axis A-A. In
this first embodiment of a scaffold pattern (pattern 200) four
links 234 connect the interior ring 212, which refers to a ring
having a ring to its left and right in FIG. 3, to each of the two
adjacent rings. Thus, ring 212b is connected by four links 234 to
ring 212c and four links 234 to ring 212a. Ring 212d is an end ring
connected to only the ring to its left in FIG. 3. The rings are
undulating and may be approximately zig-zag or sinusoidal in
shape.
[0077] A ring 212 is formed by struts 230 connected at crowns 207,
209 and 210. A link 234 is joined with struts 230 at a crown 209
(W-crown) and at a crown 210 (Y-crown). A crown 207 (free-crown)
does not have a link 234 connected to it. A "W-crown" refers to a
crown where the angle extending between a strut 230 and the link
234 at the crown 210 is an obtuse angle (greater than 90 degrees).
A "Y-crown" refers to a crown where the angle extending between a
strut 230 and the link 234 at the crown 209 is an acute angle (less
than 90 degrees). The same definitions for Y-crown and W-crown also
apply to the cell 304 below.
[0078] Preferably the struts 230 that extend from a crown 207, 209
and 210 at a constant angle from the crown center, i.e., the rings
212 are approximately zig-zag in shape, as opposed to sinusoidal
for pattern 200, although in other embodiments a ring having curved
struts is contemplated.
[0079] As such, in this embodiment a ring 212 height, which is the
longitudinal distance between adjacent crowns 207 and 209/210 may
be derived from the lengths of the two struts 230 connecting at the
crown and a crown angle .theta.. In some embodiments the angle
.theta. at different crowns will vary, depending on whether a link
234 is connected to a free or unconnected crown, W-crown or
Y-crown.
[0080] The zig-zag variation of the rings 212 occurs primarily
about the circumference of the scaffold (i.e., along direction B-B
in FIG. 3). The struts 212 centroidal axes lie primarily at about
the same radial distance from the scaffold's longitudinal axis.
Ideally, substantially all relative movement among struts forming
rings also occurs axially, but not radially, during crimping and
deployment. Although, polymer scaffolds often times do not deform
in this manner due to misalignments and/or uneven radial loads
being applied.
[0081] The rings 212 are capable of being collapsed to a smaller
diameter during crimping and expanded to a larger diameter during
deployment in a vessel. According to one aspect of the disclosure,
the pre-crimp diameter (e.g., the diameter of the tube from which
the scaffold is cut) is always greater than a maximum expanded
scaffold diameter that the delivery balloon can, or is capable of
producing when inflated. According to one embodiment, a pre-crimp
diameter is greater than the scaffold expanded diameter, even when
the delivery balloon is hyper-inflated, or inflated beyond its
maximum use diameter for the balloon-catheter.
[0082] Pattern 200 includes four links 237 (two at each end, only
one end shown in FIG. 3) having structure formed to receive a
radiopaque material in each of a pair of transversely-spaced holes
formed by the link 237. These links are constructed in such a
manner as to avoid interfering with the folding of struts over the
link during crimping, which, as explained in greater detail below,
is necessary for a scaffold capable of being crimped to a diameter
of about at most Dmin or for a scaffold that when crimped has
virtually no space available for a radiopaque marker-holding
structure.
[0083] Links 234b and 234d connect the cell 204 to the right and
left adjacent rings in FIG. 3, respectively. Link 234b connects to
cell 204 at a W-crown 209. Link 234d connects to cell 04 at a
Y-crown 210. There are four crowns 207 for cell 204, which may be
understood as four crowns devoid of a link 234 connected at the
crown. There is only one free crown between each Y-crown and
W-crown for the cell 204. Cell 204 may be referred to as a W closed
cell element since its shape resembles the letter "W", for example,
cell 204 shown by box VB.
[0084] There are four cells 204 formed by each pair of rings 212 in
pattern 200, e.g., four cells 204 are formed by rings 212b and 212c
and the links 234 connecting this ring pair, another four cells 204
are formed by rings 212a and 212b and the links connecting this
ring pair, etc. Cell 204 may be referred to as a W closed cell
element since its shape resembles the letter "W", for example, cell
204 shown by box VB.
[0085] FIG. 4 depicts another embodiment of a scaffold pattern 300.
Like the pattern 200, the pattern 300 includes
longitudinally-spaced rings 312 formed by struts 330. A ring 312 is
connected to an adjacent ring by several links 334, each of which
extends parallel to axis A-A. The description of the structure
associated with rings 212, struts 230, links 234, and crowns 207,
209, 210 in connection with FIG. 3, above, also applies to the
respective rings 312, struts 330, links 334 and crowns 307, 309 and
310 of the second embodiment, except that in the second embodiment
there are only three struts 334 connecting each adjacent pair of
rings, rather than four. Thus, in the second embodiment the ring
312b is connected to the ring 312c by only three links 334 and to
the ring 312a by only three links 334. A link formed to receive a
radiopaque marker, similar to link 237, may be included between
312c and ring 312d. In contrast to pattern 200, there are three
cells 304 formed by a ring pair and their connecting links in
pattern 300.
[0086] Links 334b and 334d connect the cell 304 to the right and
left adjacent ring in FIG. 4, respectively. Link 334b connects to
cell 304 at a W-crown 309. Link 334d connects to cell 304 at a
Y-crown 310. There are eight connected or free crowns 307 for cell
304, which may be understood as eight crowns devoid of a link 334
connected at the crown. There are one or three free crowns between
a Y-crown and W-crown for the cell 304. Cell 304 may be thought of
as a W-V closed cell element since its shape resembles the letters
"W" and "V", for example, cell 304 shown by box VA.
[0087] Comparing FIGS. 3 to 4, one can appreciate that the W cell
204 is symmetric about the axes B-B and A-A whereas the W-V cell
304 is asymmetric about both of these axes. The W cell 204 is
characterized as having no more than one crown 207 between links
234. Thus, a Y-crown crown or W-crown is always between each crown
207 for each closed cell of pattern 200. In this sense, pattern 200
may be understood as having repeating closed cell patterns, each
having no more than one crown that is not supported by a link 234.
In contrast, the W-V cell 304 has three unsupported crowns 307
between a W-crown and a Y-crown. As can be appreciated from FIG.
4A, there are three unsupported crowns 307 to the left of link 334d
and three unsupported crowns 307 to the right of link 334b.
[0088] Another embodiment of a pattern includes a repeating pattern
of W-W cells. The sequence of crests starting at a W-crown and
going around a circumference of a ring is: W-crown, 3-free crowns,
Y-crown, 2 free crowns, W-crown, etc. Thus, there are either 2 or 3
free crowns between a W-crown and Y-crown.
[0089] Crown angle .theta. in any of the patterns may be greater
than 70.degree., greater than 80.degree., greater than 90.degree.,
greater than 100.degree., 70.degree. to 80.degree., 80.degree. to
90.degree., 90.degree. to 100.degree., 100.degree. to 120.degree.,
100.degree. to 130.degree., 120.degree. to 130.degree., 120.degree.
to 140.degree., or 130.degree. to 140.degree..
[0090] The fabrication of the inventive scaffold may include the
following processes or steps: forming a hollow, thin-walled
polymeric tube (i.e., pre-cut tube), preferably with no holes in
the walls; processing that increases the strength of the polymer of
the scaffold body and also the radial strength of the scaffold;
forming a stent scaffolding made up of thin struts from the tube by
laser machining a stent pattern in the tube; optionally forming a
therapeutic coating over the scaffolding; crimping the scaffold
over a delivery balloon, and sterilization of the scaffold using
radiation, an ethylene oxide process, or some other sterilization
process. Detailed discussion of the manufacturing processes of a
bioabsorbable stent can be found elsewhere, e.g., U.S. Patent
Publication Nos. 2007/0283552 and 2012/0073733.
[0091] A pre-cut tube can be formed by a melt processing method
such as extrusion or injection molding. In extrusion, for example,
a polymer resin is fed into an extruder inlet and conveyed through
the extruder barrel as a melt above the melting temperature (Tm) of
the polymer. For example, the temperature of the melt in the
extruder may be 180 to 250.degree. C. At the end of the extruder
barrel, the polymer melt is forced through a die to form a tubular
film which is longitudinaly drawn and cooled to form the tube.
[0092] The degree of crystallinity of the tube formed from the melt
processing may be 0%, less than 5%, less than 10%, 5 to 10%, or 10
to 15%.
[0093] A polymer resin is the raw material used for the melt
processing for forming the polymeric tube. In order to provide the
high molecular weight of the finished sterilized product, the resin
has a much higher molecular weight than the finished product. The
molecular weight of the resin may be expressed in terms of the
intrinsic viscosity (IV) in dL/g. The IV of a polymer resin may be
higher than 5 dL/g, greater than 7 dL/g greater than 8 dl/g, 4 to 8
dL/g, 5 to 8 dL/g, 4 to 6 dL/g, 6 to 8 dL/g, or 5 to 7 dL/g.
[0094] The polymer of inventive scaffold after sterilization has a
number average molecular weight (Mn) of 100 to 250 kDa. The
molecular weight of the polymer decreases during the processing
steps. Most of the decrease occurs during the melt processing of
the resin and during sterilization if radiation sterilization is
used.
[0095] In addition to the type of polymer(s) and their relative
composition, the strength of the scaffold material and the radial
strength the scaffold also depend on the morphology of the scaffold
polymer. Morphology includes crystallinity, crystal domain size,
and polymer chain alignment in crystalline and amorphous domains.
Thus, the strength and radial strength can further be modified by
additional processing that modifies the morphology of the polymer,
which increases the strength of the scaffold material and the
radial strength of the scaffold.
[0096] The additional processing may increase the crystallinity of
the scaffold material which increases the strength and stiffness of
the scaffold material and the radial strength and radial stiffness
of the scaffold. Additional processing may also be performed that
increases the alignment of the scaffold polymer chains in the
circumferential or hoop direction, axial direction, or both which
increases the strength of the scaffold material and radial strength
of the scaffold. The processing can be performed prior to laser
cutting, after laser cutting, or both. Preferably, the processing
is performed prior to laser cutting.
[0097] The additional processing can include annealing the pre-cut
tube and/or the scaffold at a temperature and for a time sufficient
to increase the crystallinity to a desired level. The annealing can
be performed prior to laser cutting, after laser cutting, or both.
Preferably, the processing is performed prior to laser cutting. The
temperature may be between the glass transition temperature (Tg) of
the scaffold polymer and the melting temperature (Tm) of the
scaffold polymer. The annealing process can include heating and
maintaining a polymer construct in a temperature range for a
selected period of time. The annealing process may increase the
crystallinity from the initial crystallinity to 20 to 30%, 20 to
25%, 30 to 40%, 40 to 45%, 45 to 50%, and greater than 50%. The
annealing temperature may be any temperature between the Tg to the
Tm of the polymer or a polymer of the scaffold. More narrowly, the
temperature may be Tg+5.degree. C., Tg+5.degree. C. to
Tg+10.degree. C., Tg+10.degree. C. to Tg+15.degree. C.,
Tg+15.degree. C. to Tg+20.degree. C., Tg+20.degree. C. to
Tg+25.degree. C., Tg+25.degree. C. to Tg+30.degree. C., or greater
than Tg+30.degree. C. The annealing time may be 1 min to 10 days,
or more narrowly, 1 min to 30 min, 30 min to 1 hr, 1 hr to 3 hr, 3
hr to 10 hr, 10 hr to 1 day, 1 day to 5 days, or 5 to 10 days.
[0098] Additionally or alternatively, the processing can include
radially deforming the pre-cut tube to increase the radial strength
of the tube. The radially expanded tube may then be laser cut to
form a scaffold. The radial expansion increases the radial strength
both through an increase in crystallinity and induced polymer chain
and crystal alignment in the circumferential or hoop direction. The
radial expansion process may be performed by several processes
including blow molding (e.g., US 2011/0066222) or by expanding over
a mandrel (e.g., WO 2014/045068). In blow molding, the pre-cut tube
is disposed within a mold and heated to a temperature between Tg
and Tm and expanded by increasing a pressure inside of the
tube.
[0099] In embodiments of additional processing without radial
expansion, a tube may be formed by melt processing having a target
thin scaffold thickness. The formed tube may also have a target
diameter of a finished scaffold or target scaffold diameter. The
tube may then be annealed to increase the crystallinity to a
desired level, as disclosed herein. In some embodiments, the tube
may be annealed at a fixed diameter which may be performed by
annealing over a tubular mandrel having an outside diameter the
same as the inside diameter of the scaffold or outside diameter
slightly smaller to allow a friction fit of the tube over the
mandrel. Alternatively, the formed tube may have a diameter larger
than the target scaffold diameter, for example, 1 to 10%, or more
narrowly, 5 to 10% larger. The formed tube may then be annealed
over a mandrel having an outside diameter equal to the target
diameter of the finished scaffold. The formed tube may then be
shrunk to fit over the mandrel when annealed so that it has the
target scaffold diameter after the annealing. After any of these
annealing alternatives, the annealed scaffold may then be cut to
form the scaffold.
[0100] In embodiments of additional processing including radial
expansion, a tube may be formed by melt processing having a formed
tube thickness greater than the target thin scaffold thickness and
a tube diameter less than the target scaffold diameter. In such
embodiments, the formed tube may be radially expanded so that the
radially expanded tube has the target scaffold diameter and the
target thin scaffold thickness. The tube may also be axially
elongated during the radial expansion. The radially expanded tube
may then be cut to form the scaffold.
[0101] The degree of radial expansion may be quantified by the
radial expansion ratio (RE ratio): ID.sub.expanded/ID.sub.initial
or the percent expansion (% RE)=(RE ratio -1).times.100%. The % RE
may be 200 to 400%, 400 to 500%, 500 to 550%, 550 to 600%, or
greater than 500%. Similarly, the degree of axial elongation, may
be quantified by an axial elongation (AE) ratio,
L.sub.elonagated/L.sub.original or the percent Axial extension (%
AE)=(AE ratio -1).times.100%. The % AE may be 20% to 50%, 50% to
100%, 100% to 200%, or greater than 200%.
[0102] Exemplary embodiments include a formed tube with a thickness
of 75 microns to 150 microns and outer diameter of 2 mm to 5.0
mm.
[0103] In alternative embodiments, the formed tube may be radially
expanded so that the radially expanded tube has the target thin
scaffold thickness, but with a diameter slightly larger (e.g., 1 to
10% larger) than the target scaffold diameter. The radially
expanded tube may then be annealed, as described above, to shrink
fit the tube over a mandrel so that the annealed radially expanded
tube has the target scaffold diameter.
[0104] The size of the crystalline domains may also influence the
properties of the polymer and scaffold. It has been found that a
larger number of smaller crystalline domains improve fracture
toughness and thus improve radial strength. The temperature of the
additional processing that increases crystallinity (annealing,
radial expansion) influences the size of the crystalline domains
generated. It has found that lower temperatures closer to Tg favor
smaller crystalline domains, for example, Tg to Tg+30 or Tg+10 to
Tg+30. The scaffold may include crystalline domain sizes of less
than 10 nm, 10 nm to 50 nm, 10 to 20 nm, 10 to 30 nm, 20 to 40 nm,
40 to 50 nm, or greater than 50 nm. The disclosed range may
correspond to the average crystalline domain size or a majority of
the crystalline domain sizes.
[0105] It is believed that the high molecular weight of the polymer
tube may provide improved polymer orientation from radial
expansion, and thus, improved radial strength over a thicker target
tube thickness. The inventors have found that radial expansion of
lower molecular weight PLLA tubes made from a resin of 3.8 dL/g, an
orientation gradient results between the inside diameter (ID) and
the outside diameter (OD) of the expanded tubes and scaffolds made
from the tubes. The tubes were expanded using blow molding. The
scaffold after sterilization had an Mn of 70 to 100 kDa. The tubes
had an initial wall thickness of 0.0215 in and were expanded 400%
from an outer diameter of 0.068 in to an outer diameter of 0.1365
in with a thickness of about 0.0062 inches.
[0106] Studies using polarized light microcopy (PLM) of a radial
section of the scaffold have shown that the degree of orientation
of polymer chains or crystals decreases between the ID and the OD
of the scaffold. Polarized light microscopy refers to optical
microscopy techniques involving illumination of sample with
polarized light. PLM is most commonly used on birefringent samples
where the polarized light interacts strongly with the sample and so
generates contrast with the background. Birefringence refers to the
optical property of a material having a refractive index that
depends on the polarization and propagation and direction of light.
Such materials are optically anisotropic and are said to be
birefringent (or birefractive). The birefringence is often
quantified as the maximum difference between refractive indices
exhibited by the material. Crystals with asymmetric crystal
structures and plastics under mechanical stress are often
birefringent.
[0107] Optical isotropy means having the same optical properties in
all directions. An optically isotropic material may have
crystallites that are smaller than a resolution limit, or have
crystallites that are randomly oriented relative to each other and
therefore have no measurable difference in orientation.
[0108] Polarized light microscopy is capable of distinguishing
between isotropic and anisotropic substances. There are two
polarizing filters in a polarizing microscope termed the polarizer
and analyzer. A Michel-Levy Chart arises when polarized white light
is passed through a birefringent sample. The Michel-Levy chart
includes interference colors that describe optical retardance due
to crystallite orientation. Retardance refers to the difference in
phase shift between two characteristic polarizations of light upon
reflection from an interface. Silver at the far left of the chart
indicates very little orientation and the sequence of colors from
right to left reveals increasing orientation.
[0109] Specifically, the PLM studies of the tubes showed that from
the ID to about 50 microns from the ID, there is high induced
polymer orientation. The polarized light micrographs of thin
sections progressing from OD to ID show that that the outermost 40
to 50 microns of the expanded tube has a low anisotropy, as shown
by a first order silver Michel-Levy color, an optical path
difference (OPD) ca. 280 nm. However, a more strongly oriented
region is observed in the innermost 50 to 70 microns exhibiting a
first order gold to first order red, OPD of 420 to 560 nm. Thus, at
about 50 to 70 microns from the ID, a transition was observed in
the direction form inner to outer from high induced orientation to
low orientation. The radial section from about 50 to 70 microns
from the ID to the OD had little or no orientation. A
semi-quantitative comparison of the magnitude of the gradient in
orientation or anisotropy may be given as the change in retardance
divided by the distance over which the change occurs, from the
inner diameter to the outer diameter, which is 100% change over 50
microns.
[0110] It is believed that the gradient in orientation may be due
in part to the significant difference in radial strain experienced
between ID and OD of the extruded tubing during expansion. The
degree of strain of the wall material decreases from the ID to the
OD. Additionally, it is believed that the longer heat exposure of
the outer surface and section to the heated glass mold results in
faster relaxation of polymer chains in the radial section that
causes a loss of induced orientation.
[0111] It is expected that with a high molecular weight polymer
disclosed herein as compared the polymer in the above cited study,
the relaxation time of the polymer chains is much longer and hence
the orientation of the outer section will be better preserved.
[0112] Furthermore, higher expansion ratios for both hoop
(>400%) and axial (>200%) directions may be achieved without
resulting in the scaffolds being too brittle. The inventors have
also found that when PLLA scaffolds also fabricated from resin with
IV of 3.8 dL/g was expanded with expansion ratio of 500% at the
hoop direction, cracks and fractures were seen. Higher molecular
weight polymer creates additional toughness and strength through
the effective transfer of load and dispersion of stress across
multiple chains. This provides the capability to process the
material into expanded tubing and lased scaffolds with higher
orientation at both hoop and axial directions, rendering higher
strength.
[0113] The expanded tube including the disclosed materials or the
inventive scaffolds may have a high radial uniformity of polymer
and crystal orientation through their thickness. The tubes or
scaffolds may have a change in retardance as measured by PLM from
the inner diameter to a selected distance to the outer diameter of
less than 100%, less than 80%, less than 50%, less than 30%, less
than 10%, 10 to 30%, 30 to 60%, or 60 to 80%. The selected distance
may be 50%, 60%, 70%, 80%, 90%, 100%, 50 to 70%, 60 to 80%, 80 to
90%, or 90 to 100% of the thickness of the scaffold. The polymer
composition, molecular weight of the tube, the radial expansion
conditions, or any combination may be adjusted to obtain any of
these ranges in changes in retardance.
[0114] The degree of crystallinity of the pre-cut tube or scaffold
prior to the processing may be less than 5%, 1 to 5%, 5 to 10%,
less than 10%, 10 to 15%, less than 30%, or 15 to 30%. In an
embodiment, the crystallinity prior to processing can be between
10-25%. The degree crystallinity of the processed tube, cut
scaffold, crimped scaffold, sterilized scaffold, may be 20 to 30%,
20 to 25%, 30 to 40%, 40 to 45%, 45 to 50%, and greater than
50%.
[0115] The polymer of a scaffold may have a Young's modulus greater
than 500 MPa, or more narrowly, 500 to 600 MPa, 600 to 700 MPa, or
700 to 1000 MPa. The polymer of a scaffold may have a flexular
modulus of greater than 2.5 GPa, or more narrowly, 2.5 to 3 GPa, 3
to 5 GPa, 5 to 6 GPa, 6 to 10 GPa, 6 to 8 GPa, 8 to 10 GPa, or
greater than 10 GPa. The properties of the scaffold can be adjusted
with enhanced processing that are disclosed herein. The properties
disclosed for the scaffolds disclosed herein may refer to the
properties of the scaffold in a finished state, before or after
sterilization.
[0116] The various embodiments of the device may be configured to
eventually completely absorb from an implant site. The device may
provide drug delivery once implanted, provide mechanical support to
the vessel, and then gradually completely absorb away. The device
may also be configured to provide no mechanical support to a vessel
and serve primarily as a drug delivery vehicle. The device may be
configured to completely erode away within 6 months, 6 to 12
months, 12 to 18 months, 18 months to 2 years, or greater than 2
years.
[0117] A completely bioresorbable device may still include some
nonbiodegradable elements such as radiopaque markers or particulate
additives. The polymers of the device can be biostable,
bioresorbable, bioabsorbable, biodegradable, or bioerodable.
Biostable refers to polymers that are not biodegradable. The terms
biodegradable, bioresorbable, bioabsorbable, and bioerodable are
used interchangeably and refer to polymers that are capable of
being completely degraded and/or eroded into different degrees of
molecular levels when exposed to bodily fluids such as blood and
can be gradually resorbed, absorbed, and/or eliminated by the body.
The processes of breaking down and absorption of the polymer can be
caused by, for example, by hydrolysis and metabolic processes.
[0118] A scaffold may have a tendency to decrease in diameter or
recoil (e.g., 2 to 10%) right after implantation (i.e., less than
about 30 minutes post-implantation) as well as over a period of
days, weeks, or months. Once implanted, the device may not have
radial strength sufficient to reduce or prevent the immediate or
long-term recoil.
[0119] The mechanical properties of the scaffold material disclosed
herein may include elongation at break (ultimate elongation),
tensile modulus, and strength. The scaffold polymer or material may
have an elongation at break less than 5%, 5 to 10%, 10 to 25%, 25
to 50%, 50 to 100%, 100 to 200%, 200 to 400%, or greater than 400%
at 25 deg C., 37.degree. C., or in a range of 25 to 37.degree. C.
in a dry state or in a wet state. The scaffold polymer or material
may have a tensile modulus less than 100 MPa, 100 to 2600 Mpa, 100
to 200 MPa, 200 to 400 MPa, 400 to 600 MPa, 600 to 800 MPa, 800 to
1000 MPa, 1000 to 1200 MPa, 1200 to 1400 MPa, 1400 to 1600 MPa,
1600 to 1800 MPa, 1800 to 2000 MPa, 2000 to 2200 MPa, 2200 to 2400
MPa, 2400 to 2600 MPa, or greater than 2600 MPa at 25 deg C.,
37.degree. C., or in a range of 25 to 37.degree. C. in a dry state
or in a wet state. The wet state may correspond to soaking the
material for at least 2 minutes in a simulated body fluid such as a
phosphate buffered saline solution.
[0120] Drug delivery from the device can be provided from a coating
on a surface of the stent body of the device. The coating may be in
the form a neat drug. Alternatively, the coating may include a
polymer matrix with the drug mixed or dissolved in the polymer. The
polymer matrix can be bioresorbable. Suitable polymers for the drug
delivery polymer can include any PLA-based polymer disclosed
herein, any other polymers disclosed herein, and copolymers and
blends thereof in any combination.
[0121] The coating can be formed by mixing the polymer and the drug
in a solvent and applying the solution to the surface of the
device. The drug release rate may be controlled by adjusting the
ratio of drug and polymeric coating material. The drug may be
released from the coating over a period of one to two weeks, up to
one month, one to three months, one to four months, up to three
months, or up to four months after implantation. Thickness of the
coating on the device body may 1 to 20 microns, 1 to 2 microns, 1
to 5 microns, 2 to 5 microns, 3 to 5 microns, 5 to 10 microns, or
10 to 20 microns. In some embodiments, the stent body of the device
includes a drug release coating and the body is free of drug, aside
from any incidental migration of drug into the body from the
coating. The Mn of the coating polymer may be less than 40 kDa, 40
to 60 kDa, 60 to 80 kDa, 80 to 100 kDa.
[0122] Alternatively or additionally, the drug can also be embedded
or dispersed into the body of device, and be slowly released up to
months (e.g., one to three months or three to six months after
implantation) and while the device is degrading. In this case, the
drug can be included with the polymer when the tube is formed that
is used to form the device. For example, the drug can be included
in the polymer melt during extrusion or injection molding or in a
solution when the tube is formed from dipping or spraying or
casting.
[0123] The final device can be balloon expandable or self
expandable. In the case of a balloon expandable device, the
geometry of the device can be an open-cell structure similar to the
stent patterns disclosed herein or closed cell structure, each
formed through laser cutting a hollow thin-walled tube.
[0124] In a balloon expandable device, when the device is crimped
from a fabricated diameter to a crimped or delivery diameter onto a
balloon, structural elements plastically deform. The device may
have minimal recoil outward so the delivery diameter may different
slightly from the crimped diameter. Aside from this minimal recoil,
the device retains a crimped or delivery diameter without an inward
force on the balloon due to the plastically deformed structural
elements.
[0125] The device is radially expandable at, for example,
37.degree. C. in body fluid or simulated body fluid. When the
device is expanded by a balloon, the structural elements
plastically deform. The device is expanded to an intended expansion
or deployment diameter and retains the intended expansion diameter
or a diameter slightly less due to acute recoil inward due to
inward pressure from the vessel during the about the first 30
minutes. The diameter may vary slightly after the acute period due
to biological interactions with the vessel, stress relaxation, or
both. At the final expanded diameter, the device does not exert any
chronic outward force, which is a radial outward force exerted by
the device in excess of the radial inward force exerted by the
vessel on device.
[0126] In the case of a self-expandable device, when the device is
compressed from a fabricated diameter to a delivery diameter on a
balloon, the structural elements deform elastically. Therefore, to
retain the device at the delivery diameter, the device is
restrained in some manner with an inward force, for example with a
sheath or a band. The compressed device is expanded to an intended
expansion or deployment diameter by removing the inward restraining
force which allows the device to self-expand to the intended
deployment diameter. The structural elements deform elastically as
the device self-expands. If the final expansion diameter is the
same as the fabricated diameter, the device does not exert any
chronic outward force. If the final expansion diameter is less than
the fabricated diameter, the device does exert a chronic outward
force.
[0127] The geometric structure of the device is not limited to any
particular stent pattern or geometry. The device can have the form
of a tubular scaffold structure that is composed of a plurality of
ring struts and link struts. The ring struts form a plurality of
cylindrical rings arranged about the cylindrical axis. The rings
are connected by the link struts. The scaffold comprises an open
framework of struts and links that define a generally tubular body
with gaps in the body defined by the rings and struts.
[0128] This open framework of struts and links may be formed from a
thin-walled cylindrical tube by a laser cutting device that cuts
such a pattern into the thin-walled tube that may initially have no
gaps in the tube wall. The scaffold may also be fabricated from a
sheet by rolling and bonding the sheet to form the tube.
[0129] A stent or scaffold may have lengths of between 8 and 18 mm,
18 and 36 mm, 36 and 40 mm or even between 40 and 200 mm as
fabricated or when implanted in an artery. Exemplary lengths
include 12 mm, 14 mm, 18 mm, 24 mm, or 48 mm. The scaffold may have
a pre-crimping or as-fabricated diameter of 2 to 3 mm, 2.5 to 3.5
mm, 3 to 4 mm, 3 to 5 mm, 5 to 10 mm, 6 to 8 mm, or any value
between and including these endpoints. Diameter may refer to the
inner diameter or outer diameter of the scaffold. Exemplary
diameters include 2.5 mm, 3.0 mm, 3.25 mm, 3.5 mm, 4 mm, 5 mm, or 6
mm. The struts of the scaffold may have a radial wall thickness or
width of 150 microns, 80 to 100 microns, 100 to 150 microns, 150 to
200 microns, 200 to 250 microns, 250 to 300 microns, 300 to 350
microns, 350 to 400 microns, or greater than 400 microns. Any
combination of these ranges for radial wall thickness and width may
be used.
[0130] The scaffold may be configured for being deployed by a
non-compliant or semi-compliant balloon from a delivery diameter of
0.8 to 1 mm, 1 to 1.2 mm, 1.2 to 1.4 mm, 1.4 to 1.6 mm, 1.6 to 1.8
mm, and 1.8 to 2.2 mm, 1 mm, 1.2 mm, 1.3 mm, 1.4, mm, 1.6 mm, 1.8
mm, or 2 mm. Exemplary balloon sizes include 2.5 mm, 3 mm, 3.5 mm,
4 mm, 5.5 mm, 5 mm, 5.5 mm, 6 mm, 6.5 mm, 7 mm, or 8 mm, where the
balloon size refers to a nominal inflated or deployment diameter of
the balloon. The scaffold may be deployed to a diameter of between
2.5 mm and 3 mm, 3 mm and 3.5 mm, 3.5 mm and 4 mm, 4 mm and 10 mm,
7 and 9 mm, or any value between and including the endpoints.
Embodiments of the invention include the scaffold in a crimped or
delivery diameter over and in contact with a deflated catheter
balloon.
[0131] The intended deployment diameter may correspond to, but is
not limited to, the nominal deployment diameter of a catheter
balloon which is configured to expand the scaffold. A device
scaffold may be laser cut from a tube (i.e., a pre-cut tube) that
is less than an intended deployment diameter. In this case, the
pre-cut tube diameter may be 0.5 to 1 times the intended deployment
diameter or any value in between and including the endpoints.
[0132] A device scaffold may be laser cut from a tube (i.e., a
pre-cut tube) that is greater than an intended deployment diameter.
In this case, the pre-cut tube diameter may be 1 to 1.5 times the
intended deployment diameter, or any value in between and including
the endpoints.
[0133] The device of the present invention may have a selected high
crush recovery and crush resistance. Crush recovery describes the
recovery of a tubular device subjected to a pinch or crush load.
Scaffolds having a high crush recovery are particularly useful for
treatment of the superficial femoral artery since upon implantation
a scaffold is subjected to high crushing forces. The crush recovery
can be described as the percent recovery to the device pre-crush
shape or diameter from a certain percent crushed shape or diameter.
Crush resistance is the minimum force required to cause a permanent
deformation of a scaffold. The crush recovery and crush resistance
can be based on a pre-crush shape or diameter of an as-fabricated
device prior to crimping and expansion or a device after it has
been crimped and expanded to an intended deployment diameter. The
crush recovery of the device can be such that the device attains
greater than about 70%, 80% or 90% of its diameter after being
crushed to at least 50% of its pre-crush diameter.
[0134] The crush recovery and crush resistance of a balloon
expandable scaffold that undergoes plastic deformation when crimped
and deployed depend both on the scaffold material and scaffold
pattern. Exemplary crush recoverable balloon expandable scaffold
patterns can be found in US 2011/0190872 and US 2014/0067044.
[0135] A coating may be formed over the scaffold by mixing a
coating polymer (e.g., a PLA polymer) and a drug (e.g., a
macrocyclic drug) in a solvent and applying the solution to the
surface of the scaffold. The application may be performed by
spraying, dipping, ink-jet printing, or rolling the scaffold in the
solution. The coating may be formed as a series of layers by
spraying or dipping followed by a step to remove all or most of
residual solvent via, for example, evaporation by heating. The
steps may then be repeated until a desired coating thickness is
achieved.
[0136] The drug release rate may be controlled by adjusting the
ratio of drug and polymeric coating material. The drug to polymer
ratio may be between 5:1 to 1:5. The drug may be released from the
coating over a period of one to two weeks, up to one month, or up
to three months after implantation. Thickness or average thickness
of the coating on the device body may be less than 4 microns, 3
microns, 2.5 microns, 1 to 20 microns, 1 to 2 microns, 2 to 3
microns, 2 to 2.9 microns, 2 to 2.5 microns, 1 to 5 microns, 2 to 5
microns, 3 to 5 microns, 5 to 10 microns, or 10 to 20 microns. The
coating may be over part of the surface or the entire surface of a
scaffold substrate. In some embodiments, the body of the device
includes a drug release coating and the body is free of drug, aside
from any incidental migration of drug into the body from the
coating.
[0137] In some embodiments, the coating may include a primer layer
between the scaffold body or structure and a drug delivery coating
layer to enhance the adhesion of the drug coating to the scaffold.
Alternatively, the coating may have no primer layer and only a drug
delivery coating layer.
[0138] The coated scaffold may then be crimping over a delivery
balloon. The crimped scaffold may then be packaged and then
sterilized with radiation such as electron-beam (E-Beam) radiation
or a low temperature ethylene oxide process (see e.g., US
2013/0032967). The range of E-beam exposure may be between 20 and
30 kGy, 25 to 35 kGy, or 25 to 30 kGy.
[0139] The device body may include or may be coated with one or
more therapeutic agents, including an antiproliferative,
anti-inflammatory or immune modulating, anti-migratory,
anti-thrombotic or other pro-healing agent or a combination
thereof. The anti-proliferative agent can be a natural proteineous
agent such as a cytotoxin or a synthetic molecule or other
substances such as actinomycin D, or derivatives and analogs
thereof (manufactured by Sigma-Aldrich 1001 West Saint Paul Avenue,
Milwaukee, Wis. 53233; or COSMEGEN available from Merck) (synonyms
of actinomycin D include dactinomycin, actinomycin IV, actinomycin
I1, actinomycin X1, and actinomycin C1), all taxoids such as
taxols, docetaxel, and paclitaxel, paclitaxel derivatives, all
olimus drugs such as macrolide antibiotics, rapamycin, everolimus,
novolimus, myolimus, deforolimus, umirolimus, biolimus, merilimus,
temsirolimus structural derivatives and functional analogues of
rapamycin, structural derivatives and functional analogues of
everolimus, FKBP-12 mediated mTOR inhibitors, biolimus,
perfenidone, prodrugs thereof, co-drugs thereof, and combinations
thereof. Representative rapamycin derivatives include
40-O-(3-hydroxy)propyl-rapamycin,
40-O-[2-(2-hydroxy)ethoxy]ethyl-rapamycin, or
40-O-tetrazole-rapamycin, 40-epi-(N1-tetrazolyl)-rapamycin (ABT-578
manufactured by Abbott Laboratories, Abbott Park, Ill.), prodrugs
thereof, co-drugs thereof, and combinations thereof.
[0140] The anti-inflammatory agent can be a steroidal
anti-inflammatory agent, a nonsteroidal anti-inflammatory agent, or
a combination thereof. In some embodiments, anti-inflammatory drugs
include, but are not limited to, novolimus, myolimus, alclofenac,
alclometasone dipropionate, algestone acetonide, alpha amylase,
amcinafal, amcinafide, amfenac sodium, amiprilose hydrochloride,
anakinra, anirolac, anitrazafen, apazone, balsalazide disodium,
bendazac, benoxaprofen, benzydamine hydrochloride, bromelains,
broperamole, budesonide, carprofen, cicloprofen, cintazone,
cliprofen, clobetasol propionate, clobetasone butyrate, clopirac,
cloticasone propionate, cormethasone acetate, cortodoxone,
deflazacort, desonide, desoximetasone, dexamethasone dipropionate,
diclofenac potassium, diclofenac sodium, diflorasone diacetate,
diflumidone sodium, diflunisal, difluprednate, diftalone, dimethyl
sulfoxide, drocinonide, endrysone, enlimomab, enolicam sodium,
epirizole, etodolac, etofenamate, felbinac, fenamole, fenbufen,
fenclofenac, fenclorac, fendosal, fenpipalone, fentiazac,
flazalone, fluazacort, flufenamic acid, flumizole, flunisolide
acetate, flunixin, flunixin meglumine, fluocortin butyl,
fluorometholone acetate, fluquazone, flurbiprofen, fluretofen,
fluticasone propionate, furaprofen, furobufen, halcinonide,
halobetasol propionate, halopredone acetate, ibufenac, ibuprofen,
ibuprofen aluminum, ibuprofen piconol, ilonidap, indomethacin,
indomethacin sodium, indoprofen, indoxole, intrazole, isoflupredone
acetate, isoxepac, isoxicam, ketoprofen, lofemizole hydrochloride,
lomoxicam, loteprednol etabonate, meclofenamate sodium,
meclofenamic acid, meclorisone dibutyrate, mefenamic acid,
mesalamine, meseclazone, methylprednisolone suleptanate,
momiflumate, nabumetone, naproxen, naproxen sodium, naproxol,
nimazone, olsalazine sodium, orgotein, orpanoxin, oxaprozin,
oxyphenbutazone, paranyline hydrochloride, pentosan polysulfate
sodium, phenbutazone sodium glycerate, pirfenidone, piroxicam,
piroxicam cinnamate, piroxicam olamine, pirprofen, prednazate,
prifelone, prodolic acid, proquazone, proxazole, proxazole citrate,
rimexolone, romazarit, salcolex, salnacedin, salsalate,
sanguinarium chloride, seclazone, sermetacin, sudoxicam, sulindac,
suprofen, talmetacin, talniflumate, talosalate, tebufelone,
tenidap, tenidap sodium, tenoxicam, tesicam, tesimide, tetrydamine,
tiopinac, tixocortol pivalate, tolmetin, tolmetin sodium,
triclonide, triflumidate, zidometacin, zomepirac sodium, aspirin
(acetylsalicylic acid), salicylic acid, corticosteroids,
glucocorticoids, tacrolimus, pimecorlimus, prodrugs thereof,
co-drugs thereof, and combinations thereof.
[0141] These agents can also have anti-proliferative and/or
anti-inflammatory properties or can have other properties such as
antineoplastic, antiplatelet, anti-coagulant, anti-fibrin,
antithrombonic, antimitotic, antibiotic, antiallergic, antioxidant
as well as cystostatic agents. Examples of suitable therapeutic and
prophylactic agents include synthetic inorganic and organic
compounds, proteins and peptides, polysaccharides and other sugars,
lipids, and DNA and RNA nucleic acid sequences having therapeutic,
prophylactic or diagnostic activities. Nucleic acid sequences
include genes, antisense molecules which bind to complementary DNA
to inhibit transcription, and ribozymes. Some other examples of
other bioactive agents include antibodies, receptor ligands,
enzymes, adhesion peptides, blood clotting factors, inhibitors or
clot dissolving agents such as streptokinase and tissue plasminogen
activator, antigens for immunization, hormones and growth factors,
oligonucleotides such as antisense oligonucleotides and ribozymes
and retroviral vectors for use in gene therapy. Examples of
antineoplastics and/or antimitotics include methotrexate,
azathioprine, vincristine, vinblastine, fluorouracil, doxorubicin
hydrochloride (e.g. Adriamycin.RTM. from Pharmacia & Upjohn,
Peapack N.J.), and mitomycin (e.g. Mutamycin.RTM. from
Bristol-Myers Squibb Co., Stamford, Conn.). Examples of such
antiplatelets, anticoagulants, antifibrin, and antithrombins
include sodium heparin, low molecular weight heparins, heparinoids,
hirudin, argatroban, forskolin, vapiprost, prostacyclin and
prostacyclin analogues, dextran, D-phe-pro-arg-chloromethylketone
(synthetic antithrombin), dipyridamole, glycoprotein IIb/IIIa
platelet membrane receptor antagonist antibody, recombinant
hirudin, thrombin inhibitors such as Angiomax a (Biogen, Inc.,
Cambridge, Mass.), calcium channel blockers (such as nifedipine),
colchicine, fibroblast growth factor (FGF) antagonists, fish oil
(omega 3-fatty acid), histamine antagonists, lovastatin (an
inhibitor of HMG-CoA reductase, a cholesterol lowering drug, brand
name Mevacor.RTM. from Merck & Co., Inc., Whitehouse Station,
N.J.), monoclonal antibodies (such as those specific for
Platelet-Derived Growth Factor (PDGF) receptors), nitroprusside,
phosphodiesterase inhibitors, prostaglandin inhibitors, suramin,
serotonin blockers, steroids, thioprotease inhibitors,
triazolopyrimidine (a PDGF antagonist), nitric oxide or nitric
oxide donors, super oxide dismutases, super oxide dismutase
mimetic, 4-amino-2,2,6,6-tetramethylpiperidine-1-oxyl
(4-amino-TEMPO), estradiol, anticancer agents, dietary supplements
such as various vitamins, and a combination thereof. Examples of
such cytostatic substance include angiopeptin, angiotensin
converting enzyme inhibitors such as captopril (e.g. Capoten.RTM.
and Capozide.RTM. from Bristol-Myers Squibb Co., Stamford, Conn.),
cilazapril or lisinopril (e.g. Prinivil.RTM. and Prinzide.RTM. from
Merck & Co., Inc., Whitehouse Station, N.J.). An example of an
antiallergic agent is permirolast potassium. Other therapeutic
substances or agents which may be appropriate include
alpha-interferon, and genetically engineered epithelial cells. The
foregoing substances are listed by way of example and are not meant
to be limiting. Other active agents which are currently available
or that may be developed in the future are equally applicable.
[0142] "Molecular weight" refers to either number average molecular
weight (Mn) or weight average molecular weight (Mw). References to
molecular weight (MW) herein refer to either Mn or Mw, unless
otherwise specified. The Mn may be as measured by Gel Permeation
Chromatography with refractive index detection relative to
polystyrene standards. Suitable mobile phase solvents are acetone,
tetrahydrofuran, chloroform, 1,1,1-trichloroethane,
2,2,2-trifluoroethanol, and hexafluoro-2-propanol,
[0143] "Semi-crystalline polymer" and other terms relating to
crystalline polymer may be as defined in Pure Appl. Chem., Vol. 83,
No. 10, pp. 1831-1871, 2011. Semi-crystalline polymer refers to a
polymer that has or can have regions of crystalline molecular
structure and amorphous regions. The crystalline regions may be
referred to as crystallites, lamella, or spherulites which can be
dispersed or embedded within amorphous regions.
[0144] The "degree of crystallinity" may be expressed in terms of,
w.sub.c (mass fraction), .phi..sub.c (volume fraction) and refers
to mass fraction or volume fraction of crystalline phase in a
sample of polymer. The mass-fraction and the volume-fraction
degrees of crystallinity are related by the equation,
w.sub.c=.phi..sub.c .rho./.rho..sub.c, where .rho. and .rho..sub.c
are the mass concentrations (mass densities) of the entire sample
and of the crystalline phase, respectively. The degree of
crystallinity can be determined by several experimental techniques.
Among the most commonly used are: (i) x-ray diffraction, (ii)
calorimetry, (iii) mass density measurements, (iv) infrared
spectroscopy (IR), (v) solid-state NMR spectroscopy, and (vi) vapor
permeability.
[0145] The "glass transition temperature," Tg, 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 Tg corresponds to the
temperature where the onset of segmental motion in the chains of
the polymer occurs. When an amorphous or semi-crystalline 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 increased, the heat capacity increases. The
increasing heat capacity corresponds to an increase in heat
dissipation through movement. Tg of a given polymer can be
dependent on the heating rate and can be influenced by the thermal
history of the polymer as well as its degree of crystallinity.
Furthermore, the chemical structure of the polymer heavily
influences the glass transition by affecting mobility.
[0146] The Tg can be determined as the approximate midpoint of a
temperature range over which the glass transition takes place.
[ASTM D883-90]. The most frequently used definition of Tg uses the
energy release on heating in differential scanning calorimetry
(DSC). As used herein, the Tg refers to a glass transition
temperature as measured by differential scanning calorimetry (DSC)
at a 20.degree. C./min heating rate.
[0147] The "melting temperature" (Tm) is the temperature at which a
material changes from solid to liquid state. In polymers, Tm is the
peak temperature at which a semicrystalline phase melts into an
amorphous state. Such a melting process usually takes place within
a relative narrow range (<20.degree. C.), thus it is acceptable
to report Tm as a single value.
[0148] "Elastic deformation" refers to deformation of a body in
which the applied stress is small enough so that the object
retains, substantially retains, or moves towards its original
dimensions once the stress is released.
[0149] The term "plastic deformation" refers to permanent
deformation that occurs in a material under stress after elastic
limits have been exceeded.
[0150] "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 a 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.
[0151] "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.
[0152] "Strength" refers to the maximum stress along an axis which
a material will withstand prior to fracture. The ultimate strength
is calculated from the maximum load applied during the test divided
by the original cross-sectional area.
[0153] "Modulus" and "stiffness" 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. The modulus or the stiffness typically is
the initial slope of a stress--strain curve at low strain in the
linear region. For example, a material has both a tensile and a
compressive modulus.
[0154] The tensile stress on a material may be increased until it
reaches a "tensile strength" which refers to the maximum tensile
stress which a material will withstand prior to fracture. The
ultimate tensile strength is calculated from the maximum load
applied during a test divided by the original cross-sectional area.
Similarly, "compressive strength" is the capacity of a material to
withstand axially directed pushing forces. When the limit of
compressive strength is reached, a material is crushed.
[0155] "Elongation at break" or "ultimate elongation" is the
elongation recorded at the moment of rupture of a specimen in a
tensile elongation test, expressed as a percentage of the original
length or the strain.
[0156] "Toughness" is the amount of energy absorbed prior to
fracture, or equivalently, the amount of work required to fracture
a material. One measure of toughness is the area under a
stress-strain curve from zero strain to the strain at fracture. The
units of toughness in this case are in energy per unit volume of
material. See, e.g., L. H. Van Vlack, "Elements of Materials
Science and Engineering," pp. 270-271, Addison-Wesley (Reading, P
A, 1989).
[0157] 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.
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