U.S. patent application number 14/470858 was filed with the patent office on 2014-12-18 for stent formed from crosslinked bioabsorbable polymer.
The applicant listed for this patent is Abbott Cardiovascular Systems Inc.. Invention is credited to Lothar W. Kleiner.
Application Number | 20140371403 14/470858 |
Document ID | / |
Family ID | 42934995 |
Filed Date | 2014-12-18 |
United States Patent
Application |
20140371403 |
Kind Code |
A1 |
Kleiner; Lothar W. |
December 18, 2014 |
STENT FORMED FROM CROSSLINKED BIOABSORBABLE POLYMER
Abstract
A stent having a stent body made from a crosslinked
bioabsorbable polymer is disclosed. A method of making the stent
including exposing a tube formed from a bioabsorbable polymer to
radiation to crosslink the bioabsorbable polymer and forming a
stent body from the exposed tube is disclosed. The tube can include
a crosslinking agent which induces crosslinking upon radiation
exposure. Additionally or alternatively, the bioabsorbable polymer
can be a copolymer that crosslinks upon exposure to radiation in
the absence of a crosslinking agent.
Inventors: |
Kleiner; Lothar W.; (Los
Altos, CA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Abbott Cardiovascular Systems Inc. |
Santa Clara |
CA |
US |
|
|
Family ID: |
42934995 |
Appl. No.: |
14/470858 |
Filed: |
August 27, 2014 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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13399143 |
Feb 17, 2012 |
8846071 |
|
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14470858 |
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12422143 |
Apr 10, 2009 |
8147744 |
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13399143 |
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Current U.S.
Class: |
525/445 |
Current CPC
Class: |
A61L 31/14 20130101;
C08G 63/46 20130101; A61L 31/06 20130101; A61L 31/148 20130101 |
Class at
Publication: |
525/445 |
International
Class: |
A61L 31/06 20060101
A61L031/06; A61L 31/14 20060101 A61L031/14; C08G 63/46 20060101
C08G063/46 |
Claims
1. A stent comprising a stent body for supporting a vascular lumen,
wherein the stent body comprises a crosslinked bioabsorbable
polymer, wherein the crosslinked bioabsorbable polymer is formed
through crosslinking of a copolymer of degradable functional groups
and highly reactive functional groups, wherein the degradable
functional groups are derived from monomers that form a
biodegradable polymer, wherein the highly reactive functional
groups are derived from a lactone functionalized with an alkene or
an alkyne group, wherein the crosslinking is between the degradable
functional groups and the highly reactive functional groups.
2. The stent of claim 8, wherein the degradable functional groups
are derived from L-lactic acid.
3. The stent of claim 8, wherein the degradable functional groups
are derived from monomers selected from the group consisting of
glycolic acid, caprolactone, dioxanone, D-lactic acid, mandelic
acid, trimethylene carbonate, 4-hydroxy butyrate, and butylene
succinate.
4. The stent of claim 8, wherein the lactone is
allyl-.delta.-caprolactone or a-diallyl-.delta.-caprolactone
5. The stent of claim 8, wherein the lactone is
a-alkyne-.delta.-valerolactone or
a-alkyne-.delta.-caprolactone.
6. The stent of claim 8, wherein the copolymer of the form AxBy,
wherein A is L-lactic acid and B is an alkene or alkyne
copolymerized with A, and wherein x is the mole % of A and y is the
mole % of B in the copolymer.
7. The stent of claim 6, wherein x is 90-96 wt % and y is 4-10 wt
%.
8. The stent of claim 6, wherein B is selected from the group
consisting of a-allyl-.delta.-valerolactone,
a-diallyl-.delta.-valerolactone, a-alkyne-.delta.-valerolactone,
a-allyl-.delta.-caprolactone, a-diallyl-.delta.-caprolactone, and
a-alkyne-.delta.-caprolactone.
Description
[0001] This application is a divisional application of U.S.
application Ser. No. 13/399,143 filed on Feb. 17, 2012 which is a
divisional application of U.S. application Ser. No. 12/422,143
filed on Apr. 10, 2009, now U.S. Pat. No. 8,147,744, each of which
is incorporated herein by reference.
BACKGROUND OF THE INVENTION
[0002] 1. Field of the Invention
[0003] This invention relates to methods of manufacturing polymeric
medical devices, in particular, stents.
[0004] 2. Description of the State of the Art
[0005] 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.
[0006] 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. Stent are also used widely in
endovascular applications, such as in the popliteal artery.
[0007] Stents are typically composed of 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
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.
[0008] 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.
[0009] 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. Effective concentrations at the treated site require
systemic drug administration which often produces adverse or even
toxic side effects. Local delivery is a preferred treatment method
because it administers smaller total medication levels than
systemic methods, but concentrates the drug at a specific site.
Local delivery thus produces fewer side effects and achieves better
results.
[0010] A medicated stent may be fabricated by coating the surface
of either a metallic or polymeric scaffolding with a polymeric
carrier that includes an active or bioactive agent or drug.
Polymeric scaffolding may also serve as a carrier of an active
agent or drug.
[0011] The stent must be able to satisfy a number of mechanical
requirements. The stent must be capable of withstanding the
structural loads, namely radial compressive forces, imposed on the
stent as it supports the walls of a vessel. Therefore, a stent must
possess adequate radial strength. Radial strength describes the
external pressure that a stent is able to withstand without
incurring clinically significant damage. Additionally, a stent
should be sufficiently rigid to 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. Furthermore, the stent should
possess sufficient toughness or resistance to fracture from stress
arising from crimping, expansion, and cyclic loading.
[0012] Some treatments with implantable medical devices require the
presence of the device only for 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 device disappear may be by fabricating the device in whole or in
part from materials that erode or disintegrate through exposure to
conditions within the body. Thus, erodible portions of the device
can disappear or substantially disappear from the implant region
after the treatment regimen is completed. After the process of
disintegration has been completed, no portion of the device, or an
erodible portion of the device will remain. In some embodiments,
very negligible traces or residue may be left behind. Stents
fabricated from biodegradable, bioabsorbable, and/or bioerodable
materials such as bioabsorbable polymers can be designed to
completely erode only after the clinical need for them has
ended.
[0013] However, there are potential shortcomings in the use of
polymers as a material for stents. For example, the mechanical
properties and other properties are susceptible to degradation
during processing.
SUMMARY OF THE INVENTION
[0014] Various embodiments of the present invention include a stent
comprising a stent body for supporting a vascular lumen, wherein
the stent body comprises a crosslinked bioabsorbable polymer formed
from crosslinking a bioabsorbable polymer, wherein the
bioabsorbable polymer is a polymer formed through a
transesterification reaction between a degradable polyester and a
diol or a triol followed by a chain extension conducted with the
degradable polymer and an alkene or alkyne.
[0015] Additional embodiments of the present invention include a
stent comprising a stent body for supporting a vascular lumen,
wherein the stent body comprises a crosslinked bioabsorbable
polymer, wherein the crosslinked bioabsorbable polymer is formed
through crosslinking of a copolymer of a monomer that forms a
biodegradable polymer and a lactone functionalized with an alkene
or an alkyne group.
[0016] Further embodiments of the present invention include a stent
comprising a stent body for supporting a vascular lumen, wherein
the stent body comprises a crosslinked bioabsorbable polymer formed
from crosslinking a bioabsorbable polymer, wherein the crosslinks
are formed from a crosslinking agent that links functional groups
of the bioabsorbable polymer when exposed to radiation.
BRIEF DESCRIPTION OF THE DRAWINGS
[0017] FIG. 1 depicts a stent.
DETAILED DESCRIPTION OF THE INVENTION
[0018] Various embodiments of the present invention relate to
implantable medical devices, such as a stents, made from
crosslinked bioabsorbable polymers. The embodiments further relate
to methods of making the devices that include crosslinking the
bioabsorbable polymer through radiation exposure. The embodiments
are generally applicable to any tubular polymeric implantable
medical device. In particular, the methods can be applied to
tubular implantable medical devices such as self-expandable stents,
balloon-expandable stents, and stent-grafts.
[0019] A stent may include a pattern or network of interconnecting
structural elements or struts. FIG. 1 depicts a view of a stent
100. In some embodiments, a stent may include a body, backbone, or
scaffolding having a pattern or network of interconnecting struts
or structural elements 105. Stent 100 may be formed from a tube
(not shown). The structural pattern of the device can be of
virtually any design. The embodiments disclosed herein are not
limited to stents or to the stent pattern illustrated in FIG. 1.
The embodiments are easily applicable to other patterns and other
devices. The variations in the structure of patterns are virtually
unlimited. A stent such as stent 100 may be fabricated from a tube
by forming a pattern with a technique such as laser machining or
chemical etching.
[0020] A stent such as stent 100 may be fabricated from a polymeric
tube or a sheet by rolling and bonding the sheet to form the tube.
A tube or sheet for making a stent is conventionally formed by
extrusion or injection molding. A stent pattern, such as the one
pictured in FIG. 1, can be formed in a tube or sheet with a
technique such as laser cutting or chemical etching. The stent can
then be crimped on to a balloon or catheter for delivery into a
bodily lumen.
[0021] An implantable medical device can be made partially or
completely from a biodegradable, bioabsorbable, or biostable
polymer. A polymer for use in fabricating an implantable medical
device can be biostable, bioabsorbable, biodegradable or
bioerodable. Biostable refers to polymers that are not
biodegradable. The terms biodegradable, bioabsorbable, and
bioerodable are used interchangeably and refer to polymers that are
capable of being completely degraded and/or eroded 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,
hydrolysis and metabolic processes.
[0022] A stent made from a biodegradable polymer is intended to
remain in the body for a duration of time until its intended
function of, for example, maintaining vascular patency and/or drug
delivery is accomplished. After the process of degradation,
erosion, absorption, and/or resorption has been completed, no
portion of the biodegradable stent, or a biodegradable portion of
the stent will remain. In some embodiments, very negligible traces
or residue may be left behind.
[0023] The duration of a treatment period depends on the bodily
disorder that is being treated. In treatments of coronary heart
disease involving use of stents in diseased vessels, the duration
can be in a range from about a month to a few years. However, the
duration is typically up to about six months, twelve months,
eighteen months, or two years. In some situations, the treatment
period can extend beyond two years. The stent is expected to be
completely degraded away from the vessel at the end of the
treatment period.
[0024] As indicated above, a stent has certain mechanical
requirements such as high radial strength, high modulus, high
fracture toughness, and high fatigue resistance. A stent that meets
such requirements greatly facilitates the delivery, deployment, and
treatment of a diseased vessel. A polymeric stent with inadequate
mechanical properties can result in mechanical failure or recoil
inward after implantation into a vessel.
[0025] With respect to radial strength, the strength to weight
ratio of polymers is usually smaller than that of metals. To
compensate for this, a polymeric stent can require significantly
thicker struts than a metallic stent, which can result in an
undesirably large profile.
[0026] Additionally, polymers that are sufficiently rigid to
support a lumen at conditions within the human body may also have
low fracture toughness since they may exhibit a brittle fracture
mechanism. For example, these include polymers that have a glass
transition temperature (Tg) above human body temperature (Tbody),
which is approximately 37.degree. C. Such polymers may exhibit
little or no plastic deformation prior to failure. It is important
for a stent to be resistant to fracture throughout the range of use
of a stent, i.e., crimping, delivery, deployment, and during a
desired treatment period. PLLA is but one example of the class of
semicrystalline polymers for which the above description is true.
The Tg of PLLA has been reported to vary between approximately 55
and 65.degree. C. Medical Plastics and Biomaterials Magazine, March
1998.
[0027] Certain embodiments of the present invention include a stent
having a stent body made from a crosslinked bioabsorbable polymer.
Embodiments can include generally a device body made from the
crosslinked bioabsorbable polymer. Additionally, embodiments of the
present invention further include making the stent body, or more
generally, the device body. Embodiments of the method can include
forming a construct, such as a tube that includes a bioabsorbable
polymer. The bioabsorbable polymer may be uncrosslinked.
Alternatively, the bioabsorbable polymer may already have some
crosslinks. The bioabsorbable polymer of the construct becomes
crosslinked when exposed to radiation. The bioabsorbable polymer
may be exposed to radiation sufficient to crosslink the
bioabsorbable polymer.
[0028] In exemplary embodiments, the bioabsorbable polymer can be a
homopolymer or a copolymer. The bioabsorbable polymer can also be a
polymer blend of two or more different types of polymer, either a
miscible polymer blend or an immiscible polymer blend. The
crosslinking in a polymer blend may result linking or bonding
between polymers of different types. Alternatively, the
crosslinking can be selective, in that only one type of polymer is
crosslinked or only certain types of polymers are crosslinked in
the blend.
[0029] Exemplary bioabsorbable polymers include poly(L-lactide)
(PLLA), poly(D-lactide) (PDLA), polyglycolide (PGA), polymandelide
(PM), polycaprolactone (PCL), poly(trimethylene carbonate) (PTMC),
polydioxanone (PDO), poly(4-hydroxy butyrate) (PHB), and
poly(butylene succinate) (PBS), poly(DL-lactide) (PDLLA), and
poly(L-lactide-co-glycolide) (PLGA). Polymers that are preferred
for a stent body are those that have thermal stability in the range
at or close to Tbody, since such polymer may be rigid and maintain
a high modulus and compressive strength at Tbody so that the stent
can support a lumen. Such polymers have a Tg above human body
temperature, preferably at least 10, 20, or 30.degree. C. greater
than human body temperature. PLLA and PLGA are examples of such
polymers.
[0030] In additional embodiments, the bioabsorbable polymer can
also be a blend of PLLA and PDLA to create a stereocomplex, which
is expected to further enhance the thermal and mechanical stability
of the polymer. Exemplary blends can have a ratio of PDLA to PLLA
of between 0 and 1, although the blends can have a ratio greater
than one.
[0031] The degree of crosslinking may be characterized by crosslink
or crosslinking density. The crosslink density can be expressed as
the average molecular weight (number average or weight average)
between crosslink sites (Mc). Alternatively, the crosslink density
can be expressed as the mole fraction of monomer units which are
crosslink points (Xc). An Introduction to Plastics, Hans-Georg
Elias 2.sup.nd ed. Wiley (2003). Crosslink density, the molecular
weigh between crosslinks (Mc), can be determined by known methods
such as dynamical mechanical analysis (DMA).
[0032] The crosslink density can further be described as gel
fractions. The gel fraction is calculated by the amount of
insoluble material in solvent, when the crosslinked polymer is
mixed with a solvent for the uncrosslinked polymer, with the
following equation:
Gel fraction(%)=(Wg/W.sub.0)100
where W.sub.o and Wg are the dried weights of the initial polymer
and its remaining weight (the gel component which corresponds to
the crosslinked component) after dissolution in a solvent at room
temperature. When a crosslinked polymer is mixed into a solvent, a
portion which is sufficiently crosslinked swells rather than
dissolves in the solvent. A portion which is not crosslinked or not
sufficiently crosslinked dissolves in the solvent. Therefore, the
crosslinked portion may be separated from the remainder of the
polymer so that a gel fraction provides a measure of the degree of
crosslinking.
[0033] In embodiments of the present invention, the crosslink
density of the crosslinked bioabsorbable polymer of the stent body
can be determined from a number of techniques including equilibrium
swelling (also known as degree of swelling), NMR spectroscopy,
dynamic mechanical analysis, and gel fraction. Gel fraction is
typically used and should be at least 1%. Sometimes, crosslink
density and gel fraction are used interchangeably since gel
fraction is related to crosslink density. For the purposes of this
disclosure, these terms will be used interchangeably. The crosslink
density of the bioabsorbable polymer of the tube after radiation
exposure as determined by its gel fraction is increased by at least
1%. More specifically, the gel fraction is or is increased to
between 1-5%, 5-20%, 20-50%, 50-70%, or greater than 70%.
[0034] Additionally, embodiments of the present invention further
include making the stent body, or more generally, the device body.
Embodiments of the method can include forming a construct, such as
a tube that includes a bioabsorbable polymer. The bioabsorbable
polymer of the construct is crosslinked when exposed to
radiation.
[0035] As described in more detail below, the crosslinking can be
caused or induced partially or completely by the presence of a
crosslinking agent mixed or dispersed within the polymer.
Alternatively or additionally, the crosslinking can be caused or
induced by chemical reaction and bonding between reactive moieties
present on different polymer chains of the bioabsorbable polymer.
In this alternative, the crosslinking is bonding between functional
groups of the polymer without being linked or bonded by a
crosslinking agent that is distinct or separate from the polymer
prior to the crosslinking.
[0036] The crosslinked bioabsorbable polymer can result in a stent
body with high strength sufficient to support a bodily lumen for a
desired time period, with high fracture toughness, and with
acceptable recoil after deployment (less than 10% of the deployed
diameter). The crosslinked bioabsorbable polymer may have a
relatively low crystallinity, for example, between 10-25%. A
distinct advantage of a stent body made from a crosslinked
bioabsorbable polymer is that the polymer can have a relatively low
crystallinity (e.g., 10-25%) while providing sufficient radial
strength to support blood a vessel (e.g., less than 10% recoil for
at least 1-3 months) and yet have relatively high fracture
toughness (e.g., few or no cracked structural elements upon
deployment). The crystallinity can be greater than 25%, however, it
is important that the degree of crosslinking or crosslink density
is not high enough to cause brittle behavior that results in
unacceptable fracture or failure during use of a stent.
[0037] In some embodiments, the tube is exposed to radiation which
causes the crosslinking of the polymer. A stent body is made from
the exposed and crosslinked tube. In other embodiments, a stent
body is fabricated from the tube prior to radiation exposure and
the stent body is exposed to the radiation to crosslink the
bioabsorbable polymer.
[0038] The fabrication of a stent from a tube may include
additional processing steps. In some embodiments, the polymer tube
can be radially deformed or expanded, axially deformed, or both
radially and axially deformed. The stent body can be formed from
the radially deformed, axially deformed, or radially and axially
deformed tube. The deformation tends to increase the strength and
toughness of the polymer. In particular, the radial deformation
tends to increase the radial strength of the tube. The increase in
radial strength is believed to be due to the circumferential
polymer chain orientation and an increase in crystallinity, both
induced by the deformation. Both radial and axial deformation
provide biaxial orientation of the polymer chains.
[0039] The radial deformation can be accomplished by a blow molding
process. In such a process, the polymer tube is disposed within a
cylindrical mold with a diameter greater than the polymer tube. The
polymer tube is heated, preferably so that its temperature is above
its Tg. The pressure inside of the tube is increased to cause
radial expansion of the tube so the outside surface of the tube
conforms to the inside surface of the mold. The polymer tube can be
axially deformed by a tensile force along the tube axis before,
during, and/or after the radial deformation. The polymer tube is
than cooled below Tg and further processing steps can then be
performed, such as laser machining of the tube to form a stent
pattern.
[0040] The degree of radial expansion or deformation may be
quantified by percent radial expansion:
[ Outside Diameter of Deformed Tube Original outside Diameter of
Tube - 1 ] .times. 100 % ##EQU00001##
[0041] In some embodiments, percent radial expansion can be
200-500%. In an exemplary embodiment, the percent radial expansion
is about 300%. Similarly, the degree of axial deformation may be
quantified by the percent axial elongation:
[0042] Preliminary data suggests that 200 percent radial combined
with 200% axial provides the best results.
[ Length of Deformed Tube Original Length of Tube - 1 ] .times. 100
% ##EQU00002##
[0043] In some embodiments, the tube can be elongated before during
or after the radial expansion. The percent axial elongation can be
30-100%.
[0044] In some embodiments, the tube is radially expanded prior to
the exposing step and the stent body is formed from the expanded
and exposed tube. Exposing the construct in the expanded state may
be preferable since the crosslinking may tend to reduce or inhibit
recoil and improve other physical properties, such as
toughness.
[0045] In other embodiments including the radial expansion step,
the tube can be crosslinked prior to the radial expansion step. In
this embodiment, a stent body fabricated from an expanded tube may
have a greater tendency to recoil toward the diameter of the tube
prior to radial expansion. Recoil in this manner may be desirable
for a treatment in which recoil from a deployed diameter is
acceptable or desirable.
[0046] In another embodiment, there is no radial deformation step.
In this embodiment, the polymer tube is exposed to radiation,
followed by formation of a stent pattern. Alternatively, a stent
pattern is formed in the tube, followed by radiation exposure. In
either case, the stent may then be crimped. The crimped stent may
then have a tendency to self-expand and can be deployed at an
implant site through self-expansion rather than balloon
expansion.
[0047] As indicated above, the stent body can be formed using laser
machining to form a stent pattern in the tube. For example, a
femtosecond laser can be used. Laser machining removes material of
the tube to form the stent pattern. However, material not removed
that is near a machined surface can be modified by energy from the
laser. The modification is generally undesirable since mechanical
properties are adversely effected. The modified region is referred
to as a heat affected zone.
[0048] Another processing step can include forming a therapeutic
coating layer over all or a portion of the stent body surface. The
coating can include a therapeutic substance dispersed in a
polymer.
[0049] A further processing step includes mounting the stent on a
delivery device, for example, over a balloon on a catheter. The
stent can be mounted by reducing the diameter of the stent with a
crimping process so that the stent is secured to the balloon at a
reduced diameter. In some embodiments, the stent body can be
crosslinked at the crimped diameter. However, as mentioned above,
the stent may have a strong tendency to recoil toward the crimped
diameter. Such recoil may be undesirable in many treatment
situations. However, recoil in this manner may be desirable for a
treatment in which recoil from a deployed diameter is acceptable or
desirable.
[0050] A further embodiment may be to treat the tube or stent with
radiation after expansion, then crimp, and finally treat again with
radiation for sterilization. The second sterilization dose may be
low enough so that the crosslink density is not increased
dramatically, so that recoil is not impaired.
[0051] In another embodiment, the tube or stent can be treated with
radiation after expansion, then crimp, and sterilize with EtO.
[0052] In another embodiment, the tube or stent may be treated with
radiation after extrusion, omit expansion, crimp, and finally
sterilize with EtO.
[0053] A medical device, such as a stent, typically undergoes
sterilization to reduce the bioburden of the stent to an acceptable
sterility assurance level (SAL). Bioburden refers generally to the
number of microorganisms with which an object is contaminated. SAL
is a measure of the degree of sterilization and refers to the
probability of a viable microorganism being present on a product
unit after sterilization. There are numerous methods of sterilizing
medical devices such as stents, the most common being ethylene
oxide treatment and treatment with ionization radiation such as
electron beam and gamma radiation. Generally, it is desirable for
the sterilization procedure to have little or no adverse affects on
the material properties of the stent. Stents are typically
sterilized in a crimped state after packaging.
[0054] Ethylene oxide ("EtO") sterilization is performed by
exposing the device to gaseous ethylene oxide mixtures at elevated
temperatures, at high relative humidity for a period of time to
obtain a desired bioburden level. The elevated temperatures (e.g.,
a temperature between 30.degree. C. and (Tg-5.degree. C.)) speeds
up the sterilization of the device and the dissipation of the EtO
from the device. Exposure to the EtO gas mixture can result in
degradation of the mechanical properties of a polymer or distortion
of the fabricated shape if the conditions are not chosen properly.
The temperature, relative humidity, and time need to be chosen
carefully to be compatible with the polymer. If any of the
conditions are too high, the polymer can lose its intended shape,
form, or function as the EtO, temperature, relative humidity, and
time can affect these properties. An acceptable range may be
30.degree. C. to 50.degree. C., 30 to 100% relative humidity, with
the lowest possible temperature and humidity preferred for PLLA. A
radiation crosslinked stent would be more resistant to physical
property, form, function and shape degradation after exposure to
EtO sterilization conditions. Radiation sterilization is well known
to those of ordinary skill the art. Medical devices composed in
whole or in part of polymers can be sterilized by various kinds of
radiation, including, but not limited to, electron beam (e-beam),
gamma ray, ultraviolet, infra-red, ion beam, x-ray, and laser
sterilization. A sterilization dose can be determined by selecting
a dose that provides a required SAL. A sample can be exposed to the
required dose in one or multiple passes.
[0055] However, it is known that radiation can alter the properties
of the polymers being treated by the radiation. Such radiation can
cause a drastic drop in molecular weight of the polymer due to
chains scission and formation of free radicals. This can lead to a
more brittle material prone to cracking during deployment
(expansion). Sterilization occurs after crimping.
[0056] The present invention reduces or prevents the degradation or
drastic drop in molecular weight of the polymer of a device caused
by processing conditions which cause such degradation. In
particular, the reduction in molecular weight and properties caused
by sterilization are reduced or prevented by the crosslinking.
[0057] Additionally, the fracture toughness of the bioabsorbable
polymer can be enhanced by the crosslinking since the crosslinking
reduces the degree of crystallinity. Quynh, Tran et al., European
Polymer Journal 43 1779-1785 (2007). Specifically, fracture
toughness is enhanced if the crosslinking is sufficiently high, but
can be reduced if the crosslinking is to high. Thus, if the
radiation dose is not sufficiently high to induced the sufficient
degree of crosslinking, the fracture toughness will not be
enhanced. Also, if the radiation dose is too high, a degree of
crosslinking can be induced that results in brittle behavior.
Furthermore, the degree of crystallinity of a stent produced by the
several possible crosslinking processes described above is lower
than one that is not crosslinked. Thus, it is expected that
cracking during crimping and deployment of the stent will be
reduced or eliminated. Additionally, the crosslinking is expected
to reduce the physical aging of the polymer since the degrees of
freedom in the mobility of the polymer chains in reduced in the
amorphous regions.
[0058] Furthermore, the crosslinking is also expected to enhance
the thermal and mechanical stability of the polymer. The
crosslinking tends to increase the Tg and increase the modulus.
Thus, the recoil at Tbody is reduced.
[0059] Additionally, the stent may be sterilized using EtO, which
may be more desirable than radiation since EtO may better preserve
the properties of the finished good, i.e., the stent and catheter
assembly. It has been observed in practice that EtO sterilization
cycles tend to cause bioabsorbable stents to fracture upon
deployment if the stents are not crosslinked.
[0060] As used herein, crosslinks refer generally to chemical
covalent bonds that link one polymer chain to another. A
crosslinked polymer includes crosslinks throughout a polymer
material sample. When polymer chains are linked together by
crosslinks, they lose some of their ability to move as individual
polymer chains, thus stabilizing the polymer.
[0061] Crosslinks can be formed by chemical reactions that are
initiated by heat, pressure, crosslinking agent, or radiation. The
radiation can include, but is not limited to, electron beam, gamma,
or UV light. The crosslinking induced by radiation can be caused by
or facilitated by a crosslinking agent. A crosslinking agent is a
substance or compound that promotes or regulates intermolecular
covalent bonding between polymer chains, linking them together to
create a more rigid structure. The crosslinking agent is a compound
that is separate and distinct from the polymer chains prior to the
crosslinking between which it promotes or regulates bonding. In its
role in promoting or regulating covalent boding, the crosslinking
agent becomes covalently bonded to the polymer chains. Therefore,
the crosslinking agent can become incorporated into the crosslinked
polymer.
[0062] Radiation crosslinking of polylactic acids with crosslinking
agents has been described, for example, in Mitomo, Hiroshi et al.,
Polymer 46 4695-4703 (2005); Quynh, Tran et al., European Polymer
Journal 43 1779-1785 (2007); and Quynh, J. of Applied Polymer
Science, 110, 2358-2365 (2008), which are all incorporated by
reference herein. The radiation dose, type, and mole or weight
percent of a crosslinking agent can also influence the crosslink
density. The radiation dose is directly proportional to the
crosslink density.
[0063] In certain embodiments of the invention, the polymer
construct that is to be irradiated, such as a tube, can include a
crosslinking agent. The crosslinking agent can be mixed or
dispersed within the bioabsorbable polymer of the tube. When the
tube is exposed to radiation, the crosslinking agent induces
crosslinking of the bioabsorbable polymer.
[0064] As already indicated, the degree of crosslinking depends on
the weight percent of the crosslinking agent and the radiation
dose. The tube may include an amount of crosslinking agent
sufficient to provide a desired crosslink density or gel fraction.
In exemplary embodiments, the tube includes less than 1 wt %, 1-3
wt %, 3-5 wt %, or greater than 5 wt % crosslinking agent. The
remaining material of the tube can be the bioabsorbable polymer or
consist essentially of the bioabsorbable polymer. The tube can also
include a filler material mixed with the bioabsorbable polymer and
crosslinking agent.
[0065] A limiting factor on the radiation dose and amount of
crosslinking agents is that the crosslink density should not be so
high that the bioabsorbable exhibits brittle fracture behavior
during use of the stent, e.g., during crimping and deployment.
Additionally, the concentration of crosslinking agent in the
polymer can become so high that the performance and properties of
bioabsorbable polymer are compromised. Thus, in general, the weight
percent of crosslinking agent is preferably below 5 wt %. However,
there may be polymer, crosslinking agent, and radiation dose
combinations in which concentrations above 5 wt % that would be
favorable.
[0066] The amount of crosslinking agent and radiation dose can be
varied to obtain a desired crosslink density and gel fraction.
Also, the amount of crosslinking agent and radiation dose can be
varied to obtain desire mechanical properties such as high radial
strength and high fracture toughness. The radiation dose can be
10-100 kGy, 30-40 kGy, or more narrowly 25-30 kGy. Exemplary
crosslinking agents include triallyl isocyanurate (TAIC),
trimethally isocyanurate (TMAIC), and trimethylolpropane
triacrylate (TMPTA), however, other crosslinking agents may be
used.
[0067] A crosslinking agent can be mixed or dispersed into the
bioabsorbable polymer of a tube using melt processing. For example,
the crosslinking agent can be fed into an extruder that in the
manufacture of the tube. Alternatively, the crosslinking agent can
be mixed with a polymer melt in batch and fed into a extruder or
injection molder to make the tube. For example, the structure of
TAIC is:
##STR00001##
[0068] Although other crosslinking mechanisms are possible, the
mechanism of the crosslinking can include chain scission induced by
radiation between a C--C bond in the bioabsorbable polymer, such as
PLLA, which contains a carbonyl and methyl group, but can occur
elsewhere as well. During chain scission, a free radical is formed.
The site of the free radical can then react with the C's of the
double bond of the TAIC to form a saturated C--C crosslink.
[0069] The most effective crosslinking agent based on the magnitude
of the gel fraction can be determined by measuring gel formation at
various radiation dose levels. It has been found that TAIC is the
most effective crosslinking agent in PLLA since it exhibits the
highest gel fraction for radiation doses between 20 and 100 kGy at
3 wt % for each crosslinking agent. Mitomo, Hiroshi et al., Polymer
46 4695-4703 (2005). Likewise, the most effective concentration for
a given crosslinking agent based on the magnitude of the gel
fraction can be determined by measuring gel formation at various
concentration levels and radiation doses. It has been found that a
concentration of 3 wt % TAIC is the most effective in PLLA since it
provides the highest gel fraction based on a comparison of
concentrations between 0.5 wt % and 5 wt %. Mitomo, Hiroshi et al.,
Polymer 46 4695-4703 (2005)
[0070] Additionally, the concentration of crosslinking agent and
radiation dose can be selected based on mechanical properties such
as tensile strength and elongation at break (a measure of the
fracture toughness of the polymer). It has been shown for
irradiated PLLA containing TAIC samples that the tensile strength
was the highest for 30 kGy dose compared to other doses between 0
and 50 kGy. Id. Additionally, it has been found that for a
PLLA/PDLA blend with a 1:1 ratio of the polymers containing TAIC,
the blend with the concentration of 3 wt % TAIC showed the best
mechanical properties over a range of doses between 30 and 100
kGy.
[0071] A given radiation dose results in a corresponding degree of
crosslinking. It may be desirable to minimize the residual
unreacted crosslinking agent present in the finished stent.
Therefore, in some embodiments, the amount of crosslinking agent in
the tube prior to crosslinking can be selected so that all or
substantially all of the crosslinking agents reacts to form the
crosslinks.
[0072] Alternatively, the amount of crosslinking agent can be
selected so that unreacted crosslinking agent remains in the
bioabsorbable polymer after radiation exposure. When the stent is
exposed to radiation in a sterilization step, additional
crosslinking will occur which can further reduce or completely
deplete the remaining crosslinking agent. This can be advantageous
since the crosslinking reaction inhibits undesirable chains
scission caused by radiation. Quynh, J. of Applied Polymer Science,
110, 2358-2365 (2008).
[0073] In further embodiments of the invention, the polymer
construct that is to be irradiated, such as a tube, can be composed
in whole or in part of a polymer that is crosslinkable due to
formation of links or bonds between different moieties or
functional groups of the polymer, referred to as a
self-crosslinkable polymer, when exposed to radiation. In such
embodiments, the polymer crosslinks form in the absence of a
crosslinking agent. In some embodiments, the polymer is free of
crosslinking agents that are not chemically bound to the polymer
chains. The crosslinking can be due entirely to crosslinking
between the moieties or functional groups of the polymer. In other
embodiments, the polymer can additionally include a crosslinking
agent so that crosslinking is due to the crosslinking agent and the
reaction between the moieties or functional groups of the polymer
without the aid of the crosslinking agent.
[0074] In some embodiments, the self-crosslinkable polymer can be a
copolymer that includes reactive functional groups, for example,
alkenes or alkynes, and functional groups that form biodegradable
polymers when polymerized or copolymerized. The latter functional
groups (referred to as degradable functional groups) are derived
from monomers that include, but are not limited to L-lactic acid,
glycolic acid, caprolactone, dioxanone, D-lactic acid, mandelic
acid, trimethylene carbonate, 4-hydroxy butyrate, and butylene
succinate. "Reactive" refers to upon exposure of the polymer to
radiation, crosslinking is induced at the reactive functional
groups.
[0075] The self-crosslinkable polymer can be formed through
copolymerization of compounds that have the reactive functional
groups and a monomer, such as lactic acid, to form a biodegradable,
crosslinkable polymer. The self-crosslinkable copolymer can be a
random or alternating copolymer.
[0076] Although other crosslinking mechanisms are possible, the
mechanism of the crosslinking of the self-crosslinkable polymer can
include chain scission induced by radiation between a C--C bond in
a degradable functional groups, such as lactic acid, which contains
carbonyl and the methyl group, but can occur elsewhere as well.
During chain scission, a free radical is formed. The site of the
free radical can then react with a reactive site of the highly
reactive functional groups to form a C--C crosslink.
[0077] Although other types of reactive functional groups are
possible, in some embodiments, the reactive functional groups
include alkenes or alkynes. The double or triple bonds of the
alkene or alkyne, respectively, act as the reactive sites that form
crosslinks with the degradable functional groups. In particular,
the site of the free radical can then react with the alkene to form
a saturated C--C crosslink or react with the alkyne to form a C--C
crosslink with a double bond.
[0078] In such embodiments, the bioabsorbable polymer is a
copolymer formed through copolymerization of one or more monomers,
such as lactic acid, with an alkene or alkyne. A general form of
such a self-crosslinkable copolymer is AxBy, where A is a moiety
such as lactic acid, and B is an alkene or alkyne that can
copolymerize with A, and where x is the mole % of A and y is the
mole % of B in the copolymer. Exemplary compositions can include x
as 90-96 wt % and y as 4-10 wt %, although x and y can be outside
these ranges.
[0079] For example, a self-crosslinkable polymer formed from
L-lactide and a-allyl-.delta.-valerolactone is:
##STR00002##
Additionally, a self-crosslinkable polymer formed from L-lactide
and a-alkyne-.delta.-valerolactone is:
##STR00003##
[0080] In exemplary embodiments, the reactive functional groups can
include lactones with alkene or alkyne groups. Exemplary alkene
monomers that can be copolymerized with a monomers such as lactic
acid to form a self-crosslinkable polymer include, but are not
limited to, a-allyl-.delta.-valerolactone (AVL) or
a-allyl-.epsilon.-caprolactone. Exemplary alkyne monomers that can
be copolymerized with a monomer such as lactic acid to form a
self-crosslinkable polymer including, but are not limited to,
a-alkyne-.delta.-valerolactone or a-alkyne-.epsilon.-caprolactone.
.delta.-valerolactone can be converted to
a-alkene-.delta.-valerolactone or a-alkyne-.delta.-valerolactone.
Similarly, .delta.-caprolactone can be converted to
a-alkene-.delta.-valerolactone or a-alkyne-.delta.-valerolactone.
The copolymers named above are all random copolymers.
[0081] Schemes for functionalizing lactones with allyl and alkyne
groups have been disclosed. Parrish, Bryan et al., J. of Polymer
Science: Part A: Polymer Chemistry, Vol. 40, 1983-1990 (2002) and
Parrish, Bryan et al., J. Am. Chem. Soc., 127, 7404-7410 (2005),
which are incorporated by reference herein. Lactones can be
functionalized with allyl groups .alpha. to the carbonyl as
intermediates. This .alpha. methylene group is susceptible to
functionalization under anionic conditions because of the enhanced
acidity of its protons relative to the other protons on the ring.
For example, .delta.-valerolactone can be functionalized with an
allyl to form a-allyl-.delta.-valerolactone according to the
following scheme (Id.):
##STR00004##
The .delta.-valerolactone is quenched with allyl bromide/HMPA.
Similarly, .delta.-valerolactone can be functionalized with an
alkyne to form a-alkyne-.delta.-valerolactone according to the
following scheme (Id.):
##STR00005##
Both the alkene and alkyne have high thermal stability, with the
alkene being more thermally stable than the alkyne. The high
thermal stability allows processing with reduced or no degradation
through conventional melt processing, such as extrusion.
[0082] A copolymer of the functionalized lactone and a degradable
functional group, such as L-lactic acid, may be synthesized
according to the following exemplary reaction between L-lactic acid
and a-allyl-.delta.-valerolactone to form
poly(L-lactide-co-a-allyl-.delta.-valerolactone)
(poly(LLA-co-AVL):
##STR00006##
[0083] Additionally, the self-crosslinkable polymer can be a
copolymer of .alpha.,.alpha.-diallyl-.delta.-valerolactone (DAVL)
and a degradable functional group, such as L-lactic acid. DAVL can
be synthesized as follows:
##STR00007##
[0084] L-lactic acid and DAVL are copolymerized to form
poly(LLA-co-DAVL) as follows:
##STR00008##
[0085] In other embodiments, a self-crosslinkable polymer can be
made through a scheme including a transesterification reaction
between a degradable polyester, such as PLLA, and a diol like PEG
or a triol such as 1,1,1-tris(hydroxymethyl)ethane. Then a chain
extension is conducted with the degradable polymer and an alkyne
such as glycidyl propargyl ether or alkyne valerolactone. For
example, a transesterification of PLLA and a diol like
1,1,1-tris(hydroxymethyl)ethane can be performed as follows:
##STR00009##
The chain extension with the PLLA and an alkyne such as glycidyl
propargyl ether or alkyne valerolactone is as follows:
##STR00010##
In the structure above, the pendant groups containing the dots
represent alkyne or alkene functional groups. The alkyne monomers,
glycidyl propargyl ether and alkyne valerolactone, respectively,
are shown below:
##STR00011##
[0086] The body, scaffolding, or substrate of a stent may be
primarily responsible for providing mechanical support to walls of
a bodily lumen once the stent is deployed therein. A stent body,
scaffolding, or substrate, for example, as pictured in FIG. 1, can
refer to a stent structure with an outer surface to which no
coating or layer of material different from that of which the
structure is manufactured. If the body is manufactured by a coating
process, the stent body can refer to a state prior to application
of additional coating layers of different material. "Outer surface"
refers to any surface however spatially oriented that is in contact
with bodily tissue or fluids. A stent body, scaffolding, or
substrate can refer to a stent structure formed by laser cutting a
pattern into a tube or a sheet that has been rolled into a
cylindrical shape.
[0087] In some embodiments, the stent body, scaffolding, struts, or
structural elements of the present invention may be nonporous or
substantially nonporous. Substantially nonporous refers to a
porosity of less than 0.1 percent. Alternatively, the stent body,
scaffolding, struts, or structural elements of the present
invention may be porous. Additionally, the surface of the stent
body, scaffolding, struts, or structural elements of the present
invention may have cavities or alternatively, be cavity-free.
[0088] For the purposes of the present invention, the following
terms and definitions apply:
[0089] "Molecular weight" can refer to the molecular weight of
individual segments, blocks, or polymer chains. "Molecular weight"
can also refer to weight average molecular weight or number average
molecular weight of types of segments, blocks, or polymer
chains.
[0090] The number average molecular weight (Mn) is the common,
mean, average of the molecular weights of the individual segments,
blocks, or polymer chains. It is determined by measuring the
molecular weight of N polymer molecules, summing the weights, and
dividing by N:
M _ n = i N i M i i N i ##EQU00003##
where Ni is the number of polymer molecules with molecular weight
Mi. The weight average molecular weight is given by
M _ w = i N i M i 2 i N i M i ##EQU00004##
where Ni is the number of molecules of molecular weight Mi.
[0091] "Ambient temperature" can be any temperature including and
between 20.degree. C. and 30.degree. C.
[0092] 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 semicrystalline polymer is
exposed to an increasing temperature, the coefficient of expansion
and the heat capacity of the polymer both increase as the
temperature is raised, indicating increased molecular motion. As
the temperature is raised the actual molecular volume in the sample
remains constant, and so a higher coefficient of expansion points
to an increase in free volume associated with the system and
therefore increased freedom for the molecules to move. The
increasing heat capacity corresponds to an increase in heat
dissipation through movement. Tg of a given polymer can be
dependent on the heating rate and can be influenced by the thermal
history of the polymer. Furthermore, the chemical structure of the
polymer heavily influences the glass transition by affecting
mobility.
[0093] "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,
Pa., 1989).
[0094] The underlying structure or substrate of an implantable
medical device, such as a stent can be completely or at least in
part made from a biodegradable polymer or combination of
biodegradable polymers, a biostable polymer or combination of
biostable polymers, or a combination of biodegradable and biostable
polymers. Additionally, a polymer-based coating for a surface of a
device can be a biodegradable polymer or combination of
biodegradable polymers, a biostable polymer or combination of
biostable polymers, or a combination of biodegradable and biostable
polymers.
[0095] It is understood that after the process of degradation,
erosion, absorption, and/or resorption has been completed, no part
of the stent will remain or in the case of coating applications on
a biostable scaffolding, no polymer will remain on the device. In
some embodiments, very negligible traces or residue may be left
behind. For stents made from a biodegradable polymer, the stent is
intended to remain in the body for a duration of time until its
intended function of, for example, maintaining vascular patency
and/or drug delivery is accomplished.
[0096] Other representative examples of polymers that may be used
to fabricate an implantable medical device include, but are not
limited to, poly(N-acetylglucosamine) (Chitin), Chitosan,
poly(hydroxybutyrate), poly(hydroxybutyrate-co-valerate),
polyorthoester, polyanhydride, polyester amide, poly(glycolic
acid-co-trimethylene carbonate), co-poly(ether-esters) (e.g.
PEO/PLA), polyphosphazenes, biomolecules (such as fibrin,
fibrinogen, cellulose, starch, collagen and hyaluronic acid),
polyurethanes, silicones, polyesters, polyolefins, polyisobutylene
and ethylene-alphaolefin copolymers, acrylic polymers and
copolymers other than polyacrylates, vinyl halide polymers and
copolymers (such as polyvinyl chloride), polyvinyl ethers (such as
polyvinyl methyl ether), polyvinylidene halides (such as
polyvinylidene chloride), polyacrylonitrile, polyvinyl ketones,
polyvinyl aromatics (such as polystyrene), polyvinyl esters (such
as polyvinyl acetate), acrylonitrile-styrene copolymers, ABS
resins, polyamides (such as Nylon 66 and polycaprolactam),
polycarbonates, polyoxymethylenes, polyimides, polyethers,
polyurethanes, rayon, rayon-triacetate, cellulose, cellulose
acetate, cellulose butyrate, cellulose acetate butyrate,
cellophane, cellulose nitrate, cellulose propionate, cellulose
ethers, and carboxymethyl cellulose. Another type of polymer based
on poly(lactic acid) that can be used includes graft copolymers,
and block copolymers, such as AB block-copolymers
("diblock-copolymers") or ABA block-copolymers
("triblock-copolymers"), or mixtures thereof.
[0097] Additional representative examples of polymers that may be
especially well suited for use in fabricating or coating an
implantable medical device include ethylene vinyl alcohol copolymer
(commonly known by the generic name EVOH or by the trade name
EVAL), poly(butyl methacrylate), poly(vinylidene
fluoride-co-hexafluororpropene) (e.g., SOLEF 21508, available from
Solvay Solexis PVDF, Thorofare, N.J.), polyvinylidene fluoride
(otherwise known as KYNAR, available from ATOFINA Chemicals,
Philadelphia, Pa.), ethylene-vinyl acetate copolymers, and
polyethylene glycol.
EXAMPLES
[0098] Some embodiments of the present invention are illustrated by
the following examples. The examples are being given by way of
illustration only and not by way of limitation. The parameters and
data are not be construed to unduly limit the scope of the
embodiments of the invention.
Example 1
[0099] Synthesis of
poly(L-lactide-co-a-allyl-.delta.-valerolactone) (poly(LLA-co-AVL)
was performed according to the reaction scheme described above. All
reactions were performed with 0.04 mol % Sn(Oct)2 catalyst. Details
of the synthesis and molecular weight data are provided in Table
1.
TABLE-US-00001 TABLE 1 Poly(LLA-co-AVL) composition and molecular
weight analysis. Feed Incorp. Reaction [mol %] AVL.sup.a Time [h]
Yield [%] AVL:LLA [mol %] Mn.sup.b Mw.sup.b PDI.sup.b 1a 17 38
0:100 -- 67k 109k 1.63 1b 18 37 0:100 -- 122k 221k 1.81 2 15 68
5:95 1.4 74k 125k 1.70 3 15 75 10:90 2.6 61k 135k 2.20 4a 20 48
20:80 3.9 69k 115k 1.65 4b 20 73 20:80 3.7 129k 194k 1.51 5 20 55
30:70 5.6 82k 124k 1.43 .sup.aDetermined by 1H-NMR,
.sup.bDetermined by GPC in CHCl3 using Polystyrene standards,
Molecular weight in g/mole, PDI = polydispersity index.
[0100] Samples of poly(LLA-co-AVL) were prepared with different
mole percentages of AVL. The samples were analyzed with dynamic
scanning calorimetry (DSC) to determine the Tg and the melting
temperature (Tm). The results of the DSC analysis are shown in
Table 2.
TABLE-US-00002 TABLE 2 Tg and Tm from DSC for different mole
percentages of AVL in poly(LLA-co-AVL). Mol % AVL in PLLA 0 1.4 2.6
3.9 5.6 Tg [.degree. C.] 61 59 58 52 48 Tm [.degree. C.] 175 160
164 159 150
Example 2
[0101] Synthesis of poly(L-lactide-co-a
.alpha.,.alpha.-diallyl-.delta.-valerolactone) (poly(LLA-co-DAVL)
were performed according the to reaction scheme described above.
All reactions were performed with 0.04 mol % Sn(Oct)2 catalyst. The
reactions for samples 1 and 2 were performed without an initiator.
Details of the synthesis and molecular weight data are provided in
Table 3.
TABLE-US-00003 TABLE 3 Poly(LLA-co-AVL) composition and molecular
weight Analysis. Feed Incorp. Reaction [mol %] DAVL Time [h] Yield
[%] DAVL:LLA [mol %] Mn.sup.b Mw.sup.b PDI.sup.b 1 18 57 10:90 0.6
82k 143k 1.73 2 18 39 20:80 0.6 60k 111k 1.87 3 17 45 30:70 -- 49k
81k 1.64 .sup.aDetermined by 1H-NMR, .sup.bDetermined by GPC in
CHCl3 using polystyrene standards.
[0102] 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.
* * * * *