U.S. patent application number 14/211742 was filed with the patent office on 2014-10-30 for core-sheath fibers and methods of making and using same.
The applicant listed for this patent is Abby Deleault, Toby Freyman, Joseph Lomakin, Quynh Pham, Xuri Ray Yan, Gregory T. Zugates. Invention is credited to Abby Deleault, Toby Freyman, Joseph Lomakin, Quynh Pham, Xuri Ray Yan, Gregory T. Zugates.
Application Number | 20140322512 14/211742 |
Document ID | / |
Family ID | 51789480 |
Filed Date | 2014-10-30 |
United States Patent
Application |
20140322512 |
Kind Code |
A1 |
Pham; Quynh ; et
al. |
October 30, 2014 |
CORE-SHEATH FIBERS AND METHODS OF MAKING AND USING SAME
Abstract
According to one aspect of the invention, multicomponent fiber
are provided, which comprise (a) a polymeric core that comprises a
core-forming polymer and (b) a polymeric sheath that comprises a
sheath-forming polymer that is different than the core-forming
polymer. Examples of core-forming polymers include, for instance,
crosslinked polysiloxanes and thermoplastic polymers, among others.
Examples of sheath-forming polymers include, for instance,
solvent-soluble polymers, degradable polymers and hydrogel-forming
polymers, among others. Other aspects of the present invention
pertain to methods of forming such multicomponent fibers. For
example, in certain preferred embodiments, the multicomponent
fibers are formed using coaxial electrospinning techniques. Still
other aspects of the present invention pertain to meshes and other
articles that are formed using the multicomponent fibers.
Inventors: |
Pham; Quynh; (Methuen,
MA) ; Yan; Xuri Ray; (Brighton, MA) ;
Deleault; Abby; (Boston, MA) ; Freyman; Toby;
(Waltham, MA) ; Lomakin; Joseph; (Cambridge,
MA) ; Zugates; Gregory T.; (Chelmsford, MA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Pham; Quynh
Yan; Xuri Ray
Deleault; Abby
Freyman; Toby
Lomakin; Joseph
Zugates; Gregory T. |
Methuen
Brighton
Boston
Waltham
Cambridge
Chelmsford |
MA
MA
MA
MA
MA
MA |
US
US
US
US
US
US |
|
|
Family ID: |
51789480 |
Appl. No.: |
14/211742 |
Filed: |
March 14, 2014 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
61852224 |
Mar 15, 2013 |
|
|
|
61861629 |
Aug 2, 2013 |
|
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Current U.S.
Class: |
428/220 ;
264/465; 428/373; 442/51 |
Current CPC
Class: |
D01F 1/103 20130101;
Y10T 442/186 20150401; Y10T 428/2929 20150115; D01D 5/0007
20130101; D01F 8/16 20130101; D04H 3/005 20130101 |
Class at
Publication: |
428/220 ;
428/373; 442/51; 264/465 |
International
Class: |
D01F 8/16 20060101
D01F008/16; D01D 5/00 20060101 D01D005/00; D04H 3/005 20060101
D04H003/005 |
Goverment Interests
STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH
[0002] This invention was made with government support under
Technology Innovation Program Award Number: 70NANB11H004 awarded by
the National Institute of Standards and Technology (NIST). The
government has certain rights in the invention.
Claims
1. A multicomponent fiber comprising (a) a polymeric core that
comprises a core-forming polymer and (b) a polymeric sheath that
comprises a hydrophilic polymer, wherein said core-forming fiber is
more hydrophobic than said hydrophilic polymer.
2. The multicomponent fiber of claim 1, wherein said multicomponent
fiber is formed by a core-sheath electrospinning process.
3. The multicomponent fiber of claim 1, wherein the multicomponent
fiber ranges from 0.1 to 20 microns in diameter.
4. The multicomponent fiber of claim 1, wherein the ratio of sheath
volume to core volume in the multicomponent fiber ranges from 100:1
to 1:1.
5. The multicomponent fiber of claim 1, wherein the hydrophilic
polymer is covalently crosslinked.
6. The multicomponent fiber of claim 1, wherein the hydrophilic
polymer is selected from polyvinylpyrrolidone, poly(acrylic acid),
poly(vinyl alcohol), poly(ethylene glycol), poly(propylene glycol),
poly(acrylamide), poly(methacrylates), polysaccharides, celluloses,
chitosans, alginates, carrageenan, hyaluronan, gelatin and
collagen.
7. The multicomponent fiber of claim 1, wherein the hydrophilic
polymer is a hydrophilic polyurethane.
8. The multicomponent fiber of claim 7, wherein the hydrophilic
polyurethane is an aliphatic, polyether-based polyurethane.
9. The multicomponent fiber of claim 1, wherein the core-forming
polymer is a thermoplastic polymer.
10. The multicomponent fiber of claim 1, wherein the core-forming
polymer is an aliphatic polyether-based thermoplastic
polyurethane.
11. The multicomponent fiber of claim 1, wherein the core-forming
polymer is a crosslinked polysiloxane.
12. The multicomponent fiber of claim 1, wherein the polysiloxane
is polydimethylsiloxane.
13. A nonwoven mesh formed by the multicomponent fiber of claim
1.
14. The mesh of claim 13, wherein the mesh ranges from 10 to 5000
microns in thickness and the multicomponent fiber ranges from 0.1
to 20 microns in diameter.
15. The mesh of claim 13, wherein the mesh has a modulus wet
tensile strength of at least 0.005 MPa.
16. The mesh of claim 13, wherein upon immersion in aqueous medium
at 25.degree. C. for one hour, the mesh has an absorbency of at
least 10%.
17. The mesh of claim 13, wherein the porosity of the mesh is less
than 99%.
18. A medical article comprising the mesh of claim 13.
19. A method for forming the multicomponent fiber of claim 1,
comprising electrospinning said multicomponent fiber from a first
solution comprising said hydrophilic polymer and a second solution
comprising said core-forming polymer.
20. A multicomponent fiber comprising (a) a polymeric core that
comprises a crosslinked polysiloxane and (b) a polymeric sheath
that comprises a removable sheath-forming polymer.
21. The multicomponent fiber of claim 20, wherein the
multicomponent fiber ranges from 0.1 to 20 microns in diameter.
22. The multicomponent fiber of claim 20, wherein the polysiloxane
is polydimethylsiloxane.
23. The multicomponent fiber of claim 20, wherein the
sheath-forming polymer is a dissolvable or degradable polymer.
24. A mesh formed by the multicomponent fiber of claim 20.
25. The mesh of claim 24, wherein the mesh ranges from 10 to 5000
microns in thickness and the multicomponent fiber ranges from 0.1
to 20 microns in diameter.
26. A medical article comprising the mesh of claim 24.
27. A method for forming the multicomponent fiber of claim 20,
comprising electrospinning said multicomponent fiber from a first
solution comprising said removable sheath-forming polymer and a
second solution comprising a polysiloxane pre-polymer and a
crosslinking agent.
28. A method of forming a silicone fiber, comprising: (a) forming a
composite fiber comprising a silicone core and a removable polymer
sheath and (b) removing the polymer sheath.
29. The method of claim 28, wherein the removable polymer is a
dissolvable or degradable polymer.
30. The method of claim 28, wherein the fiber is electrospun into
the form of a mesh prior to removing the polymer sheath.
Description
RELATED APPLICATIONS
[0001] This application claims the benefit of U.S. Provisional
Application No. 61/852,224, filed Mar. 15, 2013, entitled "Systems
and Methods for the Production of Silicone Fibers using Coaxial
Electrospinning" and U.S. Provisional Application No. 61/861,629,
filed Aug. 2, 2013, 2013, entitled "Biocomponent
Elastomeric-Hydrogel Fibers," each of which is incorporated herein
by reference in its entirety.
TECHNICAL FIELD
[0003] The present disclosure relates, among other things, to
core-sheath fibers, to methods of making core-sheath fibers and to
devices and applications associated with core-sheath fibers.
BACKGROUND
[0004] Fibers and collections of fibers have been used as materials
in various industrial applications, including applications in
medicine and surgery ranging from sutures to wound dressings to
skin grafts to arterial grafts, among many others. These
applications are based on the unique properties of fibers as
materials.
SUMMARY OF THE INVENTION
[0005] According to one aspect of the invention, multicomponent
fiber are provided, which comprise (a) a polymeric core that
comprises a core-forming polymer and (b) a polymeric sheath at
least partially surrounding the polymeric core that comprises a
sheath-forming polymer that is different than the core-forming
polymer. Examples of core-forming polymers include, for instance,
crosslinked polysiloxanes and thermoplastic polymers, among others.
Examples of sheath-forming polymers include, for instance,
solvent-soluble polymers, degradable polymers and hydrogel-forming
polymers, among others.
[0006] Other aspects of the present invention pertain to methods of
forming such multicomponent fibers. For example, in various
preferred embodiments, the multicomponent fibers are formed using
coaxial electrospinning techniques.
[0007] Still other aspects of the present invention pertain to
meshes and other articles that are formed using the multicomponent
fibers.
[0008] These and many other aspects and embodiments of the present
invention will become immediately apparent to those of ordinary
skill in the art upon review of the Detailed Description and Claims
to follow.
BRIEF DESCRIPTION OF THE DRAWINGS
[0009] FIG. 1 shows a photomicrograph of a cross-section of the
PLGA/PDMS sheath/core fibers formed in accordance with an
embodiment of the invention.
[0010] FIG. 2 shows the PDMS fibers of FIG. 1 after sheath layer
removal.
[0011] FIGS. 3A-3D show top-down and cross-sectional
photomicrographs of PLGA/PDMS sheath/core fibers formed in
accordance with an embodiment of the invention, both before sheath
removal (FIGS. 3A and 3C) and after sheath removal (FIGS. 3B and
3D).
[0012] FIG. 4 shows an image of water droplet (left) and an oil
droplet (right), placed on a PDMS mesh in accordance with the
present invention.
[0013] FIG. 5 is a stress-strain diagram illustrating mechanical
properties of a PDMS mesh in accordance with the present invention
as compared to a cast PDMS film.
[0014] FIGS. 6A-6C show cross-sectional photomicrographs of
PLGA/PDMS sheath/core fibers that were electrospun at three
differing sheath:core flow rates, in accordance with an embodiment
of the invention.
[0015] FIG. 7A-D shows photomicrographs of PVP/PDMS sheath/core
fibers formed in accordance with an embodiment of the invention,
which show: (A) a cross-section of core-sheath fibers where the PVP
cured at 100.degree. C.; (B) the same fibers as in (A) after they
have undergone water extraction; (c) a cross-section of core-sheath
fibers where the PVP cured at 150.degree. C.; (D) the same fibers
as in (C) after they have undergone water extraction.
[0016] FIG. 8 shows FTIR (Fourier transform infrared spectroscopy)
scans of a pure PDMS film, a pure PVP film and a PVP/PDMS
sheath/core fiber formed in accordance with an embodiment of the
invention (cured at 100.degree. C.), when dry and when wet.
[0017] FIG. 9 shows FTIR scans of a pure PDMS film, a pure PVP film
and a PVP/PDMS sheath/core fiber formed in accordance with an
embodiment of the invention (cured at 150.degree. C.), when dry and
when wet.
[0018] FIG. 10 is a stress-strain diagram illustrating mechanical
properties of PVP/PDMS sheath/core fibers formed in accordance with
an embodiment of the invention (cured at 100.degree. C. and
150.degree. C.), when dry and when wet.
[0019] FIGS. 11A and 11B shows balloon formed from a hydrated
PVP-PDMS fiber mesh cured at 100.degree. C., in accordance with an
embodiment of the invention, at two levels of expansion.
[0020] FIG. 12A-D shows photomicrographs of fibers with a
hydrophilic polyurethane (HLPU) sheath and a more hydrophobic
polyurethane (HBPU) core, also referred to herein as HLPU/HBPU
sheath/core fibers, formed at four HLPU:HBPU ratios, in accordance
with an various embodiment of the invention.
[0021] FIG. 13 shows swelling and tensile strength as a function of
HLPU content for meshes formed from HLPU/HBPU sheath/core fibers
formed in accordance with various embodiments of the invention.
[0022] FIG. 14 shows swelling and shrinkage as a function of HLPU
content for meshes formed from HLPU/HBPU sheath/core fibers formed
in accordance with various embodiments of the invention.
[0023] FIG. 15 shows swelling for meshes formed from four different
HLPU/HBPU sheath/core fibers formed in accordance with the
invention (Formulations A-D), as well as two commercially available
wound dressings.
[0024] FIG. 16 shows wet tensile strength for meshes formed from
four different HLPU/HBPU sheath/core fibers formed in accordance
with the invention (Formulations A-D), as well as two commercially
available wound dressings.
[0025] FIG. 17 shows shrinkage for meshes formed from four
different HLPU/HBPU sheath/core fibers formed in accordance with
the invention (Formulations A-D), as well as two commercially
available wound dressings.
[0026] FIGS. 18A and 18B show photomicrographs of a mesh formed
from HLPU/HBPU sheath/core fibers before and after annealing,
respectively, in accordance with an embodiment of the
invention.
[0027] FIG. 19 shows phosphate buffered saline (PBS) retention for
meshes formed from annealed (B Annealed) and non-annealed (B
Normal) HLPU/HBPU sheath/core fibers formed in accordance with the
invention, as well as two commercially available wound
dressings.
[0028] FIG. 20 shows shrinkage/expansion for meshes formed from
annealed (B Annealed) and non-annealed (B Normal) HLPU/HBPU
sheath/core fibers formed in accordance with the invention, as well
as two commercially available wound dressings.
[0029] FIG. 21 is a photomicrograph of HLPU/HBPU sheath/core fibers
with encapsulated silver nanoparticles.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS
[0030] In accordance with one aspect of the present disclosure,
multicomponent fibers are provided which comprise a polymeric core
and a polymeric sheath at least partially surrounding (i.e.,
encapsulating) the core.
[0031] As used herein, "fibers," "microfibers," and "nanofibers"
are used synonymously to refer to elongated structures that differ
only by size (with "microfibers" indicating fibers that have
cross-sectional diameters on the order of microns to hundreds of
microns, "nanofibers" indicating fibers that have cross-sectional
diameters on the order of nanometers to hundreds of nanometers, and
"fibers" indicating fibers of any size).
[0032] Fibers in accordance with the present disclosure can thus be
formed in a wide variety of sizes. Preferred overall fiber
diameters range from 0.05 to 50 microns (.mu.m) (e.g., ranging from
0.05 to 0.1 to 0.25 to 0.5 to 1 to 2.5 to 5 to 10 to 25 to 50
microns), more preferably 0.1 to 20 microns, among other possible
dimensions. Preferred core diameters range from 0.01 to 10 microns
(e.g., ranging from 0.01 to 0.025 to 0.05 to 0.1 to 0.25 to 0.5 to
1 to 2.5 to 5 to 10 microns), among other possible dimensions.
Preferred sheath thicknesses range from 0.02 to 25 microns (e.g.,
ranging from 0.02 to 0.05 to 0.1 to 0.25 to 0.5 to 1 to 2.5 to 5 to
10 to 25 microns), more preferably ranging from 0.2 to 18 microns,
among other possible dimensions.
[0033] The ratio of the sheath volume to core volume can vary
widely. Preferred sheath volume:core volume ratios range, for
example, from 100:1 to 1:100, among other values, for example
ranging from 100:1 to 50:1 to 25:1 to 10:1 to 5:1 to 2:1 to 1:1 to
1:2 to 1:5 to 1:10 to 1:25 to 1:50 to 1:100.
[0034] Multicomponent fibers in accordance with the present
disclosure can be formed using various fiber spinning techniques,
including various melt spinning and solvent spinning methods. Thus,
although solvent spinning techniques, and more particularly,
electrostatic solvent spinning techniques, are detailed herein, the
invention is not limited to such techniques. Further exemplary
techniques for forming multicomponent fibers include hot melt
spinning, melt electrospinning, centrifugal fiber spinning, wet
spinning, dry spinning, gel spinning, gravity spinning, extrusion,
extrusion spinning, and rapid prototyping, among others. Using
these and other techniques, multicomponent fibers may be formed
that comprise (a) a polymeric core that comprises a core-forming
polymer and (b) a polymeric sheath at least partially surrounding
the polymeric core that comprises a sheath-forming polymer that is
different than the core-forming polymer.
[0035] Electrospinning is a process that uses an electrical charge
to draw very fine, typically micro- or nano-scale, fibers from a
liquid. Solvent electrospinning utilizes an electrical force
applied to a polymer solution to induce electrospinning jets. As
streams associated with the jets travel in the air (or other
atmosphere), evaporation of the solvent results in a single long
polymer fibers deposited on a grounded collector. The collected
fibers can result in the formation of a mesh which may be used in
various technologies in medical and non-medical industries
including, for example, drug delivery devices, tissue engineering,
nano-scale sensors, wound dressings, self-healing coatings, and
filters, among many others.
[0036] As used herein, a "mesh" is a structure that is formed by a
collection of one or more fibers interlaced to form a three
dimensional network. Meshes include woven and non-woven meshes.
[0037] Meshes in accordance with the present disclosure can vary
widely in thickness with preferred thicknesses ranging from 10 to
5000 microns (e.g., ranging from 10 to 25 to 50 to 100 to 250 to
500 to 1000 to 2500 to 5000 microns), among other values.
[0038] Meshes in accordance with the present disclosure can vary
widely in porosity. In certain embodiments, the meshes of the
present disclosure have a porosity of 99% or less, for example,
ranging from 99% to 90% to 80% to 70% to 60% to 50% to 40% to 30%
to 20% to 10% or less. Porosity can be measured by determining the
volume of the polymer and dividing that quantity by the volume of
the mesh. In this regard, Polymer volume=Mesh mass/Polymer density;
Mesh volume=Mesh length.times.Mesh width.times.Mesh thickness=Mesh
area.times.Mesh thickness; and Mesh porosity=(Mesh volume-Polymer
volume)/Mesh volume. In various embodiments, the porosity of a
given mesh may be reduced by annealing the mesh at a temperature
and for a time wherein a decrease in mesh porosity is observed.
Electrospinning
[0039] Conventionally, core-sheath electrospinning, also referred
to herein as coaxial electrospinning, uses two concentric needles
to separately deliver two solutions, specifically, an inner core
polymer solution and an outer sheath polymer solution. The core
solution is delivered through the inner needle whereas the sheath
solution is delivered through the outer needle. Upon activation of
an electric field, the two different polymer solutions are ejected
in a continuous stream toward a grounded collector; this forms a
single core-sheath Taylor cone at the needle tip, leading to the
formation of a core-sheath fiber. The creation of core-sheath
fibers using needles, however, has limited throughput.
[0040] In certain embodiments, core-sheath fibers are generated
using a high-throughput core-sheath needleless electrospinning
fixture, which utilizes one or more slits on the surface of a
hollow vessel to co-localize numerous materials to multiple sites
that form Taylor cones, thereby promoting the formation of multiple
electrospinning jets and thus multiple electrospun fibers. The
slits on the surface of the hollow vessel thus may generate
high-throughput production of core-sheath fibers. For further
information, see e.g., U.S. Patent Pub. No. 2012/0193836 to Sharma
et al. and U.S. Patent Pub. No. 2013/0241115 to Sharma et al., the
disclosures of which are hereby incorporated by reference.
[0041] In electrospinning, each jet that forms thus leads to one
long continuous fiber that gets collected. In a typical operation
of the needleless fixture, there are approximately 10 jets that
form along the length of the slit; the collected mesh is therefore
comprised of approximately 10 very long fibers intertwined with one
another. In contrast, during the operation of open bath free
surface electrospinning used in the high-throughput core-sheath
needleless electrospinning fixture, hundreds of jets form and
disappear with each rotation of the drum. Thus, the resulting mesh
consists of thousands of relatively short fibers.
[0042] The design of the needleless electrospinning fixture takes
into account processing parameters that may enable greater control
over fiber diameter. For example, in addition to the solution
properties, solution flow rates can be manipulated to control fiber
diameter. Furthermore, the number of jets produced can also be
controlled, which may lead to differences in fiber diameter.
[0043] The fibers of any embodiment of the present disclosure may
thus be collected in a non-woven mesh form. However, alternate
embodiments include fibers that are collected as aligned fibers (as
through gap alignment or rotating drum), twisted yarns, ropes, in a
pattern, or any other method of fiber collection known in the art
of electrospinning.
Fibers with Silicone Components
[0044] Various aspects of the invention pertain to multicomponent
fibers that are formed using silicone polymers (also referred to
herein as "silicones", "siloxane polymers" or "polysiloxanes"). For
example, in certain embodiments, multicomponent fibers are formed
that comprise (a) a polymeric core that comprises one or more
silicone polymers and (b) a polymeric sheath at least partially
encapsulating the core that comprises one or more additional
polymers other than silicone, or vice versa.
[0045] The present disclosure is applicable to all siloxanes (i.e.,
compounds with --Si--O--Si linkages), including polysiloxanes,
which are formed from multiple siloxane units,
##STR00001##
where R.sub.1 and R.sub.2 are organic radicals, for example,
linear, branched or cyclic alkyl groups (e.g., methyl groups, ethyl
groups, propyl groups, isopropyl groups, butyl groups, isobutyl
groups, sec-butyl groups, tert-butyl groups, cyclohexyl groups and
so forth), which may be substituted or unsubstituted, as well as
substituted or unsubstituted aryl groups (e.g., phenyl groups, p-,
m- or o-alkyl-substituted phenyl groups, and so forth). R.sub.1 and
R.sub.2 can be the same or different.
[0046] In various embodiments, polysiloxanes including as PDMS can
be functionalized by a variety of mechanisms (e.g. plasma, UV, CVD,
etc.) to modify the surface properties (e.g. hydrophobicity, etc.)
or provide specific chemical interactions (e.g. antibody binding).
Fibers can be functionalized resulting in immobilized biomolecules
on the surface and/or in the bulk. Functionalization can provide
many new properties to the material, including biological effects,
sensor applications. Microfibers and nanofibers further enhance
these benefits by providing high surface areas and small pores, for
example.
[0047] In this regard, functional groups polymerized as pendant
groups attached to the siloxane (e.g., hydrides, hydroxyls, amines,
isocyanates, epoxies, etc.) may be used to add chemical activity
and diversity and to modify mechanical properties, swelling and
solvent resistance, and refractive index, among other properties.
The coaxial electrospinning of polysiloxanes as described herein
may be combined with functionalization to obtain silicone
microfibers and nanofibers with different properties, making them
useful in additional applications. For example, treatments which
make the fibers more hydrophilic will provide elastic, durable
filters which wet more readily. In some embodiments, a
functionalizing moiety for the PDMS is incorporated into the fiber.
Upon curing, the functional moiety in the fiber becomes
incorporated into the PDMS through siloxane chemistry. This allows
for one-step functionalizing of the PDMS. In one specific
embodiment, PDMS surfaces can be functionalized with biotin groups
by adding biotinylated phospholipids to the PDMS prepolymer before
curing, as described in Bo Huang et al., "Phospholipid
biotinylation of polydimethylsiloxane (PDMS) for protein
immobilization," Lab Chip, 2006, 6, 369-373. These biotin groups
can then be further modified with avidin-conjugated to a species of
interest, for example, proteins, antibodies or fragments thereof,
to functionalize the silicone surface. This may be useful, for
example, in removing proteins from a liquid (e.g. protein
separation) or in medical implants where preferential binding of
certain proteins is advantageous (e.g. improved endothelial cell
interactions).
[0048] It is noted that some classes of polymers, including various
siloxane polymers, are difficult to electrospin due to their low
molecular weight and flowability. In this regard, various
polysiloxanes remain flowable until they are crosslinked, which
does not allow for sufficient polymer chain entanglement for fibers
to form.
[0049] For example, polydimethylsiloxane (PDMS) is a silicon-based
organic polymer belonging to a larger group of siloxane polymers as
indicated above, which commonly exhibit properties of elasticity
and durability. The ability to manufacture fibers and constructs
made from PDMS and other siloxane polymers that exhibit such
properties, along with an ability to control the fiber diameter, is
highly advantageous in medical technologies as well various other
applications. Although attempts have been made to electrospin PDMS
fibers, the techniques developed thus far use blended polymer
systems (i.e. not pure PDMS) and there are currently no
electrospinning methods known to the inventors for manufacturing
pure PDMS fiber constructs such as meshes.
[0050] Thus, in some aspects of the present disclosure, core-sheath
electrospinning techniques are provided, which can be used form
fibers that comprise silicone materials that have not been
previously electrospun using known techniques. The fibers formed by
the techniques described herein comprise a silicone material as the
core material, and a different polymer material as the sheath
material. After fiber formation and/or collection, the core-sheath
fibers are typically crosslinked by a suitable mechanism. For
example, the fibers may be cured overnight at room temperature or
for a few hours at temperatures up to 100.degree. C., among other
cros slinking techniques.
[0051] In certain embodiments, the polymeric sheath may be formed
from hydrophilic or hydrogel materials, which are discussed in more
detail below.
[0052] In certain embodiments, the polymeric sheath may be formed
from materials that can be dissolved, degraded or otherwise removed
from the silicone core, leaving behind pure silicone fibers.
Examples of such materials include degradable polymers and
solvent-soluble polymers, including water-soluble polymers.
[0053] Examples of degradable polymers include one or more of the
following, among others: (a) polyester homopolymers and copolymers
such as polyglycolide (PGA) (also referred to as polyglycolic
acid), polylactide (PLA) (also referred to as polylactic acid)
including poly-L-lactide, poly-D-lactide and poly-D,L-lactide,
poly(lactide-co-glycolide) (PLGA), polycaprolactone,
polyvalerolactone, poly(beta-hydroxybutyrate), polygluconate
including poly-D-gluconate, poly-L-gluconate, poly-D,L-gluconate,
poly(p-dioxanone), poly(lactide-co-delta-valerolactone),
poly(lactide-co-epsilon-caprolactone), poly(lactide-co-beta-malic
acid), poly(beta-hydroxybutyrate-co-beta-hydroxyvalerate), among
others, (b) polycarbonate homopolymers and copolymers such as
poly(trimethylene carbonate), poly(lactide-co-trimethylene
carbonate) and poly(glycolide-co-trimethylene carbonate), among
others, (c) poly(ortho ester) homopolymers and copolymers such as
those synthesized by copolymerization of various diketene acetals
and diols, among others, (d) polyanhydride homopolymers and
copolymers such as poly(adipic anhydride), poly(suberic anhydride),
poly(sebacic anhydride), poly(dodecanedioic anhydride), poly(maleic
anhydride) and poly[1,3-bis(p-carboxyphenoxy)methane anhydride],
among others, (e) polyphosphazenes such as aminated and alkoxy
substituted polyphosphazenes, among others and (f) amino-acid-based
polymers.
[0054] Examples of water-soluble polymers include non-crosslinked
hydrophilic polymers, which may be selected from homopolymers and
copolymers formed from one or more of the following monomers, among
others: ethylene oxide, vinyl pyrrolidone, vinyl alcohol, vinyl
acetate, vinyl pyridine, methyl vinyl ether, acrylic acid and salts
thereof, methacrylic acid and salts thereof, hydroxyethyl
methacrylate, acrylamide, N,N-dimethyl acrylamide, N-hydroxymethyl
acrylamide, alkyl oxazolines, saccharide monomers (e.g.,
polysaccharides such as dextran, alginate, etc.), and amino acids
(e.g., hydrophilic polypeptides and proteins such as gelatin,
etc.). When crosslinked, the preceding hydrophilic polymers are
useful as hydrogels.
[0055] For normal nonwoven materials, microarchitecture is highly
dependent upon fiber diameter. Accordingly, an advantage of this
core-sheath manufacturing process in which the sheath is
subsequently removed is the ability to obtain pore sizes,
porosities and other microarchitectural features. Using the
high-throughput core-sheath needleless electrospinning fixture
(see, e.g., U.S. Patent Pub. No. 2012/0193836 and U.S. Patent Pub.
No. 2013/0241115 to Sharma et al.), the ratio of sheath-to-core
thickness can be varied to provide larger pore sizes with smaller
fibers or higher porosities with smaller fibers than can be
obtained with other fabrication techniques.
Fibers with Hydrogel Components and Components of Varying
Hydrophilicity/Hydrophobicity
[0056] Various aspects of the invention pertain to multicomponent
fibers that are formed using hydrogels. For example, in certain
embodiments, multicomponent fibers are formed that comprise (a) a
polymeric core that comprises one or more core-forming polymers and
(b) a polymeric sheath that comprises one or more hydrophilic or
hydrogel-forming polymers.
[0057] Various aspects of the invention pertain to multicomponent
fibers that comprise (a) a polymeric sheath that comprises one or
more hydrophilic polymers and (b) a polymeric core that comprises
one or more polymers that are more hydrophobic than the one or more
hydrophilic polymers. Conversely, other aspects of the invention
pertain to multicomponent fibers that comprise (a) a polymeric core
that comprises one or more hydrophilic polymers and (b) a polymeric
sheath that comprises one or more polymers that are more
hydrophobic than the one or more hydrophilic polymers.
[0058] Polymers for use as core and/or sheath polymers include
those that, upon immersion in an aqueous medium (e.g., water, PBS,
etc.) at 25.degree. C. for one hour have water absorption values
ranging anywhere from 0% to 1000% or more water, calculated as (wet
weight-dry weight)/dry weight (.times.100), for example ranging
from 0% to 1% to 2.5% to 5% to 10% to 25% to 50% to 100% to 250% to
500% to 1000% or more. As used herein, a "hydrophilic polymer" is
one that has a water absorption value ranging from 5-1000% or more
water. A "more hydrophobic" polymer, also referred to herein as a
"less hydrophilic" polymer, is defined as a polymer that absorbs
less water than a given polymer to which it is being compared.
[0059] In some embodiments, core and sheath polymers are selected
such that the ratio of the sheath polymer water absorption value
relative to the core polymer water absorption value ranges from 2:1
to 100:1 (for example ranging from 2:1 to 5:1 to 10:1 to 20:1 to
50:1 to 100:1), among other possible values, preferably 5:1 to 20:1
in certain embodiments. By way of example, the water absorption
value of the sheath polymer in Example 4 below is 500% whereas the
water absorption value of the sheath polymer is 50%, yielding a
sheath:core water absorption ratio of 10:1.
[0060] Hydrogels comprise a three dimensional crosslinked network
of hydrophilic polymers which have the ability to absorb
substantial amounts of water. Hydrogels have long been used in in
many applications in the medical field, ranging from drug delivery
to tissue engineering scaffolds. Despite many potential
applications, hydrogels have limited utility in healthcare or other
fields due to a lack of structural control and a poor understanding
of hydrogel mechanical properties. Others in the field have looked
into reinforcing hydrogels with a variety of additives. Still
others have aimed to reinforce hydrogels by making a polymeric
fiber or polymeric fiber construct (e.g. a mesh) and then
submersing it in a hydrogel or hydrogel-forming polymer before
cross-linking the polymer. Such methods and structures have been
generally ineffective, and there remains a need for hydrogel
structures with desired properties.
[0061] In certain aspects of the present disclosure,
electrospinning is used to form a fiber core that comprises one or
more fiber-forming polymers at least partially surrounded by a
sheath that comprises one or more hydrogel-forming polymers. The
resulting composite fiber may be optionally subjected to a
crosslinking step (e.g., by application of energy such as heat,
visible light or ultraviolet light, by application of a
crosslinking agent, etc.) to crosslink the hydrogel-forming
polymers, the core-forming polymers, or both. The result is a
composite fiber that has mechanical and hydration properties that
differ from either material alone. These composite fibers can be
gathered, formed or processed into various shapes (e.g., tube,
mesh, yarns, etc.) for use as medical devices or other
products.
[0062] Polyurethanes may be employed as core and/or sheath polymers
in various embodiments. Polyurethanes are generally formed from
diisocyanates and long-chain diols and, typically, chain extenders.
Aromatic diisocyanates may be selected from suitable members of the
following, among others: methylenediphenyl diisocyanate (MDI),
toluene diisocyanate (TDI), naphthalene diisocyanate (NDI),
para-phenylene diisocyanate (PPDI), 3,3'-tolidene-4,4'-diisocyanate
and 3,3'-dimethyl-diphenylmethane-4,4'-diisocyanate. Non-aromatic
(aliphatic) diisocyanates may be selected from suitable members of
the following, among others: hexamethylene diisocyanate (HDI),
dicyclohexylmethane diisocyanate (H12MDI), isophorone diisocyanate
(IPDI), cyclohexane diisocyanate (CHDI),
2,2,4-trimethyl-1,6-hexamethylene diisocyanate (TMDI), and
meta-tetramethylxylyene diisocyanate (TMXDI), among others. Long
chain diols include polyether diols (e.g., polyethylene glycol,
polyoxypropylene glycol, polytetramethylene ether glycol, etc.),
polyester diols (e.g., polybutane diol adipate, polyethylene
adipate, polycaprolactone diol, etc.), and polycarbonate diols.
Other long-chain diols include diol versions for the hydrophilic
polymers listed above. Chain extenders include short chain diols
such as 1,4 butane diol, among others.
[0063] Polyurethanes other than those described in the prior
paragraph, may also be employed as core and/or sheath polymers in
various embodiments
[0064] Hydrogels for use in the present disclosure include those
formed from hydrophilic polymers which are crosslinked via a
suitable mechanism, for example, covalently crosslinked and/or
non-covalently crosslinked (e.g., by ionic crosslinking, physical
crosslinking, etc.).
[0065] Examples of hydrophilic polymers which may be crosslinked
include various hydrophilic polymers such as those set forth above.
Further examples of hydrophilic polymers include hydrophilic
polyurethanes (e.g., polyurethanes having hydrophilic segments),
which may be physically crosslinked (e.g., via hard segments
present in the polyurethanes). Specific hydrophilic polyurethanes
include aliphatic, polyether-based polyurethanes and aromatic,
polyether-based polyurethanes, among others. It is further noted
that the hydrophilic polymers set forth above may be employed as
hydrophilic segments in polyurethanes in certain embodiments.
[0066] Examples of core-forming polymers, which include
thermoplastic polymers and polymers of varying
hydrophilicity/hydrophobicity in many embodiments, include
silicones (polysiloxanes) such as those described above,
thermoplastic polyurethanes such as aliphatic, polyether-based
polyurethanes and aromatic, polyether-based polyurethanes, among
others, and polyamides (e.g., nylon-6,6, nylon-6, nylon-6,9,
nylon-6,10, nylon-6,12, nylon-11, nylon-12, nylon-4,6, etc.), among
others. Examples of core-forming polymers further include
homopolymers and copolymers (including block copolymers) comprising
one or more of the following monomers, among others: (a)
unsaturated hydrocarbon monomers (e.g., ethylene, propylene,
isobutylene, 1-butene, 4-methyl pentene, 1-octene and other
alpha-olefins, isoprene, butadiene, etc.); (b) halogenated
unsaturated hydrocarbon monomers (e.g., tetrafluoroethylene,
vinylidene chloride, vinylidene fluoride, chlorobutadiene, vinyl
chloride, vinyl fluoride, etc.); (c) vinyl aromatic monomers
including unsubstituted vinyl aromatic monomers (e.g., styrene,
2-vinyl naphthalene, etc.) and vinyl substituted aromatic monomers
(e.g., alpha-methyl styrene), ring-substituted vinyl aromatic
monomers; and (d) relatively hydrophobic (meth)acrylic monomers,
including alkyl (meth)acrylates (e.g., isopropyl acrylate, butyl
acrylate, sec-butyl acrylate, isobutyl acrylate, cyclohexyl
acrylate, tert-butyl acrylate, hexyl acrylate, 2-ethylhexyl
acrylate, dodecyl acrylate, hexadecyl acrylate, and isobornyl
acrylate, isopropyl methacrylate, isobutyl methacrylate, t-butyl
methacrylate, cyclohexyl methacrylate, 2-ethylhexyl methacrylate,
octyl methacrylate, dodecyl methacrylate, hexadecyl methacrylate,
octadecyl methacrylate, isobornyl methacrylate, etc.), arylalkyl
(meth)acrylates (e.g., benzyl acrylate, benzyl methacrylate, etc.),
and halo-alkyl (meth)acrylates (e.g., 2,2,2-trifluoroethyl
acrylate). It is noted that many of the preceding polymers can be
employed as segments in polyurethanes in some embodiments.
[0067] Advantages associated with providing multi-component fibers
with a hydrogel sheath and a core material that differs from the
sheath material is that fibers, meshes and other constructions can
be formed which have good water absorption and retention properties
(as a result of the hydrogel material) coupled with desirable
mechanical properties such as strength, elasticity, durability and
shrinkage (as a result of the core material).
Fibers with Silicone Core and Removable Sheath
[0068] As previously noted, certain aspects of the present
disclosure pertain to multicomponent fibers that comprise (a) a
polymeric core that comprises one or more silicone polymers and (b)
a polymeric sheath that comprises one or more additional polymers
other than silicone. In certain embodiments, the polymeric sheath
may be formed from materials that can be dissolved, degraded or
otherwise removed from the silicone core, leaving behind pure
silicone fibers. Examples of such materials include degradable
polymers and solvent-soluble polymers (including water soluble
polymers) such as those set forth above, among others. As elsewhere
herein, the fibers can be formed or processed into various shapes
(e.g., tube, mesh, yarns) for use as medical devices or other
products.
[0069] In some embodiments, a silicon core-forming polymer is
co-electrospun with a removable (e.g., dissolvable or degradable)
sheath-forming polymer to create novel composite fibers. The
electrospinning may achieved by needleless electrospinning, coaxial
electrospinning, slit-surface electrospinning, or any other
suitable technique known in the art of fiber spinning.
[0070] In one preferred embodiment, detailed in Examples 1 and 2
below, fibers are formed with a PDMS core and a biodegradable
polymer sheath. Cross-linking of PDMS is performed using a two-part
system by mixing the pre-polymer and a cross-linking agent which
initiates the cross-linking reaction (exposure to heat accelerates
this reaction). As used herein, a "pre-polymer"is a polymer
material that is subjected to a cross-linking or other curing
process to create a crosslinked polymer. In other embodiments,
two-part PDMS systems can be cured by exposure to UV-light. In
still other embodiments, two-part PDMS systems can be crosslinked
into elastomers through free radical, condensation, or addition
reactions. Alternatively, one-part PDMS systems may be used which
cure upon exposure to moisture in the atmosphere or photo-curing,
among other techniques. Any of these variations in PDMS
chemistries, or other polymers that require physical or chemical
cross-linking to become a solid, may be used in the fibers and
methods described herein.
[0071] Thus, although a polysiloxane (i.e., PDMS) is exemplified as
a preferred embodiment, other embodiments may use polymers (e.g.,
thermosetting polymers, etc.) that require cross-linking to become
solid. Examples include other polysiloxanes and certain types of
polyesters, polyurethanes, polyimides, epoxies, etc.
[0072] Although degradable polymers (i.e.,
poly(lactide-co-glycolides)) are exemplified as preferred
embodiments, other embodiments may use other degradable polymers or
a solvent-soluble polymer sheath (e.g., formed from a water-soluble
sheath material such as uncrosslinked PEO, PVA, gelatin, dextran,
carbohydrates, etc.), which may be subsequently removed by
dissolution. Embodiments employing aqueous solvents as dissolution
agents generally do not result in swelling of PDMS fibers.
[0073] In some embodiments, the sheath is etched away using an
acid.
[0074] Depending on the mechanical properties of the sheath
polymer, mechanical disruption may be used to break apart the
sheath. Any combination of the described methods, or other suitable
means, may be employed to remove the sheath from the underlying
core.
[0075] In some embodiments, therapeutic agents such as small
molecule drugs, anesthetics, procoagulants, anticoagulants,
antimicrobials, biologics, RNAi, genetic material, genetic vectors,
vaccines, or particles such as silver nanoparticles are within the
polysiloxane core.
[0076] In some embodiments, a porogen (e.g., selected from salts,
sugars, etc.) is incorporated within the polysiloxane core. Upon
subsequent sheath removal, the porogen is also removed. This will
leave behind a fiber with porosity or rough surface features that
may improve hydrophobicity, among other properties. Alternatively,
a porogen may be incorporated into the sheath such that after fiber
formation, a certain percentage of the porogen is located at the
interface of the core and sheath. Upon sheath removal, there is a
negative imprint of the porogen on the polysiloxane fiber surface.
The surface of polysiloxane fibers can also be roughened by a
suitable etching process (e.g., laser etching) or mechanical
means.
[0077] Additionally, porosity can be introduced to the fibers of
the present disclosure as a product of the cross-linking reaction
that forms the fiber. For example, isocyanate functionalized PDMS
can react with water to form porous foam fibers. Another example is
acetoxy functionalized PDMS formulated with sodium bicarbonate. The
acetic acid byproduct of the cross-linking reaction can react with
sodium bicarbonate can generate gas and porosity, thus allowing for
the formation of porous foam fibers.
[0078] Manipulation of fiber size can yield different fiber
properties. For example, in filtration applications, smaller fibers
with larger pores or higher porosity can increase the permeability
and surface area. Polysiloxane materials (e.g., PDMS) as described
herein provide high durability, thermal and oxidative stability and
flexibility in combination with small pore size and high
permeability. Additionally, polysiloxane fiber meshes formed in
accordance with the present disclosure have high surface area due
to the small size fibers, which can promote adhesion and wetting
where desired. In some embodiments, these same properties may be
useful in medical applications where cell infiltration into fibers
is desired. In particular, smaller fiber diameters generally
facilitate cellular interaction, ingrowth and proliferation while
larger pores and higher permeability generally facilitate nutrient,
cytokine and gas exchange while also improving cell migration.
[0079] In additional embodiments, the sheath is left on the core in
order to form a composite fiber that contains a PDMS core and
polymer sheath (e.g., nylon, polyethylene, polystyrene,
polycarbonates, etc.) that possesses unique properties. In some
embodiments, upon immersion in water, the outer sheath may form a
hydrogel to fill the porosity of a PDMS fiber mesh.
[0080] In further embodiments, core and sheath polymers are
reversed, and a polysiloxane is used as the sheath polymer that
coats a core polymer. This allows the formation of bi-component
fibers with the polysiloxane on the outside. Additionally, removal
of the core polymer results in polysiloxane hollow fibers.
[0081] Small diameter fiber meshes can provide higher surface area,
higher permeabilities and lower pore sizes than meshes made from
larger diameter fibers. The present disclosure thus provides
materials which combine the benefits of polysiloxanes such as PDMS
and small-diameter fiber meshes. For example, solvent-resistant
filters or elastomeric, biocompatible microfiber or nanofiber
medical device components (e.g., heart valve leaflets, vascular
grafts, stent graft coverings) may be formed.
[0082] In this regard, in some embodiments, silicone meshes may be
used in heart valve leaflets. Replacement heart valves, in some
cases, use synthetic materials to recreate the native leaflets.
Native leaflets are thin, highly flexible and durable. In addition
to these properties the leaflets need to be nonthrombogenic.
Encouraging endothelialization is one of the best ways to provide
nonthrombogenic implants. The microfiber architecture provided by
electrospun silicone is thought to encourage endothelial cell
growth. However, this same porosity may lead to blood passing
through the pores of the mesh and reduced blood flow control by the
valve. This phenomenon is expected to be temporary, however, as
proteins and cells become trapped in the pores. In a preferred
embodiment, microfiber meshes of silicone are electrospun to a
thickness of between 100 to 1000 microns (um). Target fiber
diameters are between 500 nm to 10 um. These meshes are then cut
into appropriate shapes and attached to a main body which will be
implanted via open or minimally-invasive surgery. Alternate
embodiments include: providing a membrane (e.g., silicone, PLGA)
either on one side of the mesh or sandwiched between two meshes to
prevent blood flow through the mesh; functionalizing the silicone
with proteins or antibodies (e.g., CD34, VEGF) to encourage tissue
ingrowth and reendothelialization; electrospinning onto a frame
(e.g., polymer fiber, metal wire, contoured conductive mesh) to
help shape the leaflet and/or provide an attachment to the main
body; electrospinning onto a biocompatible fiber structure which
will create a composite implant (e.g., fibers provide additional
mechanical strength or varying stiffness across the leaflet); and
coating or functionalizing the fibers to decrease thrombogenicity
(e.g., heparin).
[0083] In some embodiments, silicone meshes may be used in stent
graft coverings in a method similar to the heart valve leaflet,
except that the silicone fibers are electrospun onto a tubular
collector to form a tube of silicone microfibers or nanofibers.
Preferred mesh thickness is between 100 and 1000 microns. Target
fiber diameters are between 500 nm to 10 um. This tube can then be
attached to a stent to form the stent graft. Alternatively, the
fibers may be electrospun directly onto the stent. Alternate
embodiments described for the heart valve are applicable here as
well. Advantages include the fact that silicone microfibers and or
nanofibers will encourage cellular ingrowth while providing an
elastic, biocompatible, durable implant. In some embodiments,
silicone meshes may be used in vascular grafts similar to the stent
graft design except that the tube is not attached to a stent and
the preferred mesh thickness range is larger (100 to 5000
microns).
[0084] In some embodiments, silicone meshes may be used in
bioengineered blood vessels. Much like the vascular graft above,
the silicone mesh may be fashioned into a tube and seeded with
cells ex vivo. These cells, typically fibroblasts, smooth muscle
cells and endothelial cells, are incubated under various conditions
(e.g., pulsatile flow, steady flow, no flow) in nutrient-rich
environments to grow tissue on the graft material. The silicone
microfibers and nanofibers provide advantages in encouraging cell
infiltration and growth as well as provide an elastic character
typical of blood vessels. The silicone tube may be used alone or in
combination with other natural (e.g., collagen) or synthetic (e.g.,
PTFE, ePTFE, polyurethane) materials. In other embodiments, the
graft is seeded with cells and implanted without significant
incubation or implanted without cell seeding. In the latter case,
cells from the host will infiltrate and populate the graft.
[0085] In some embodiments, silicone meshes may be used in
arteriovenous (AV) grafts and shunts. These grafts are used in
hemodialysis patients to provide better needle access for repeated
dialysis. Silicone microfiber or nanofiber meshes will provide a
robust set of mechanical properties as well as encourage cellular
ingrowth. The elasticity, durability, biocompatibility and low
thrombogenicity of silicone will improve the performance of these
grafts. In one embodiment, a silicone microfiber or nanofiber mesh
is fashioned into a tube and implanted. This tube may be
pre-treated by functionalization or coating with other materials
(e.g., heparin, collagen, gelatin, growth factors) to improve
integration and cell ingrowth. In other embodiments, the silicone
mesh may be combined with other natural or synthetic materials as
sheets or meshes to form a composite, layered structure. This
layered structure may improve the mechanical properties, the
ability to contain blood immediately after implantation or long
term durability or performance.
[0086] In some embodiments, silicone fibers electrospun into a flat
mesh configuration of thickness 500 to 5000 microns may be used in
hernia meshes. To improve mechanical properties, a composite may be
formed with biocompatible polymer fibers by electrospinning
directly onto those fibers in the desired configuration. These
fibers may also be provided in a configuration that improves
suture-ability of the mesh. In alternate embodiments, the mesh may
be functionalized, using the various methods described above, to
improve tissue ingrowth or integration.
[0087] In some embodiments, silicone meshes may be used in dural
covering. In neurosurgical procedures where the dura are
compromised, it is desirable to provide a covering to re-seal the
membrane. A silicone microfiber or nanofiber mesh, optionally
combined with a polymer membrane (e.g., silicone, PLGA, collagen)
can be used for this purpose.
[0088] In other embodiments, silicone meshes may be used in wound
dressing. Challenges for wound dressings include adherence to the
wound and permeability to air and water (wound exudates). In one
embodiment, silicone is electrospun into a mesh between 100 and
5000 microns thick. Preferred fiber diameters are between 500 nm
and 10 microns. The electrospun silicone is non-adherent to the
wound and provides high permeability and will be used a wound
contacting layer in a dressing. In another embodiment, the silicone
dressing is supplied separately and medical staff may place
additional gauze or other bandages in layers on top of the silicone
dressing. In another embodiment, the silicone mesh is combined with
a gauze or other backing material as part of the finished product
to absorb fluid and protect the wound. In still other embodiments,
the silicone can be fabricated with therapeutic agents such as
antibiotics, antifungals, topical pain relievers, disinfectants
(e.g., iodine) or the like. Another embodiment provides a silicone
mesh that has been treated with or manufactured with a hydrogel
sheath (e.g., PEG) to provide moisture to the wound bed. The
advantage of the silicone mesh in this case is the high porosity
can contain the hydrogel material while aiding in removal when the
dressing needs to be removed. In still other embodiments, the
silicone mesh is fabricated for use with negative pressure wound
therapy. In this case, the mesh is sized to be compatible with
these devices and is placed on the wound bed as negative pressure
is applied. The high permeability and porosity allow exudate
removal as well as a non-adherent dressing when it must be removed.
For application in negative pressure wound therapy, the silicone
fibers may be electrospun onto a collector with a shape and
topography similar to the intended treatment site (e.g., face,
hand, etc.). In this way, the dressing can improve the therapy by
improved conformance to the wounded tissue.
[0089] In some embodiments, silicone meshes may be used in
hemostatic applications. For hemostatic applications, the device is
configured much like the wound dressings, but the silicone
microfibers or nanofibers are fabricated or surface modified with a
prothrombotic or procoagulant agent (e.g., thrombin, kaolin,
chitosan, fibrin, etc.). The silicone provides a non-adherent
dressing that can be removed easily. In addition, the high
permeability and porosity allows the blood to penetrate and contact
a large amount of the surface area with the prothrombotic agent.
This open structure also allows for coagulation factor diffusion
back into the wound promoting clot formation. This material may
also be integrated as a non-adherent layer on other dressings
(e.g., Combat Gauze); for this application the fibers may or may
not be manufactured with a prothrombotic or procoagulant agent.
[0090] In other embodiments, silicone meshes may be used in
filtration applications. Silicone meshes may be used as filters or
as part of a filter for air, other gases, liquids, slurries or
particles. The high solvent resistance and durability provide
advantages over other microfiber and nanofiber filters. In
particular, the low pore size and high permeability of electrospun,
microfiber nanofiber meshes are desirable for filters. In addition,
the elastomeric nature provides a way to clean the filter. Simply
stretching the material biaxially, circumferentially or otherwise
will increase the pore size. Then, backflow of gas or liquid will
provide a method to clear the pores of debris or other material. In
a similar manner, cake which forms on the intake side of a filter
may be easily removed by stretching the silicone mesh allowing the
cake to fall off. The silicone microfiber or nanofiber mesh may be
used alone (preferred thickness of 100 microns to 1 cm).
Alternately, the silicone mesh may be constructed as part of a
layered filter using other commonly available filter materials. In
this case, the silicone may be electrospun directly onto another
material, placed on the other material during assembly or
electrospun onto a wire or other fiber mesh with large openings to
provide mechanical support.
[0091] In some embodiments, silicone meshes may be used in drug
delivery. Drugs may be incorporated into the silicone microfibers
or nanofibers for delivery to a patient. In one embodiment a
silicone mesh is formed with drug in the silicone solution and is
placed on the skin for cutaneous or transcutaneous delivery. In
another embodiment, the silicone microfibers or nanofibers are
formed into a mesh, tube or other structure and implanted to
deliver drugs internally. This could include the mouth or other
bodily orifices (e.g., delivery of fluoride, bleach or other
whitening substances to teeth).
[0092] In other embodiments, silicone meshes may be used in
barriers to modulate water penetration for controlled drug
delivery. A mesh of polysiloxane fibers such as silicone fibers
could act as a barrier to modulate drug release. For example, if a
drug delivery device has a large burst, a PDMS mesh (which is
relatively hydrophobic) can be placed around the device to prevent
or slow water contact with the device. Additionally, since silicone
is elastic, expansion of the mesh can lead to changes in its
porosity and pore size, resulting in an increase of water so as to
cause more drug release.
[0093] In some embodiments, silicone meshes may be used in
pressure-sensitive adhesive bandages. In this embodiment the
silicone microfibers or nanofibers are electrospun from a silicone
which has adhesive properties. The mesh can then be applied to skin
and will adhere well, but will provide water and air permeability
to facilitate natural skin function and health. This material can
be used in bandages, as part of a wound dressing or for drug
delivery patches.
[0094] In some embodiments, silicone meshes may be used for
oil-water separation as silicone is known to be relatively
hydrophobic. With high pore volume fraction, a silicone microfiber
or nanofiber mesh will separate oil from water. The silicone may be
surface treated, functionalized or doped with additives to make it
more oleophilic or hydrophobic. In this application, the silicone
mesh may be used as a filter or placed into oil-water mixtures to
remove oil or to separate oil from water. This may be extended to
other systems containing hydrophilic and hydrophobic materials or
phases. Because the mesh is highly elastic, the mesh can be
stretched, squeezed, or compressed to clean/remove the oil from the
pores for recovery of the oil and/or reuse of the mesh.
Additionally, silicone also absorbs organic solvent and can also be
used to separate aqueous from organic solvents. The high surface
area of microfiber meshes makes it particularly efficient and
appealing for these applications.
[0095] In some embodiments, silicone meshes may be used in
textiles. Silicone microfibers or nanofibers may also be used in
textile applications where high elasticity, durability and
permeability is desired. In other applications, the hydrophobicity
or liquid repellent nature of silicone microfiber or nanofiber
meshes (due to architecture) can be used to provide protection from
liquids while still allowing air permeability to enable the skin to
"breath".
[0096] In various embodiments, the composite fiber can be collected
into aligned fiber bundles like a yarn. These yarns will act as
strong, elastic fibers that can be used (e.g., sutures) or
processed further, including: twisting multiple yarns together into
a rope, weaving multiple yarns together into a woven sheet, tube or
other shape, braiding multiple yarns together into a stent,
scaffold or other tubular structure.
Fibers with Polymeric Core and Hydrophilic or Hydrogel Sheath
[0097] Novel materials can be produced by forming various
hydrophilic or hydrogel materials around various polymeric core
materials, which act as a reinforcing material for the hydrophilic
or hydrogel material. The encapsulated polymer material can impart
unique material properties (mechanical, chemical, thermal, etc.) to
the hydrophilic or hydrogel material that would otherwise not be
possible.
[0098] More particularly, in some embodiments, a core-forming
polymer is co-electrospun with a hydrophilic or hydrogel-forming
polymer to create novel composite fibers with a polymeric fiber
core that is at least partially surrounded by a hydrophilic or
hydrogel sheath. The electrospinning may achieved by needleless
electrospinning, coaxial electrospinning, slit-surface
electrospinning, or any other suitable technique known in the
spinning art. The result is a composite fiber that has mechanical
and hydration properties that are distinct from either material
alone. These composite fibers can be gathered, formed or processed
into various shapes (e.g., tube, mesh, yarns, etc.) for use as
medical devices or other products.
[0099] Any appropriate hydrophilic or hydrogel-forming material may
be used as the sheath polymer and, like the selection of the
polymeric core material, the hydrophilic or hydrogel-forming
material can be selected to suit the particular purpose of the
composite fiber. For example, with regard to the hydrophilic or
hydrogel polymer sheath, crosslinked PVP, PEO, PVA, and hydrophilic
polyurethanes, among other polymers, as well as xerogels, aerogels,
etc., may be used, among many other possibilities. Other hydrogel
polymers include crosslinked versions of hydrophilic polymers such
as those listed above.
[0100] Similarly, any appropriate polymer may be used for the
core-forming polymer, depending on the mechanical or chemical needs
at hand. In some embodiments, the fiber core is formed using a
relatively hydrophobic polymer. While certain embodiments employ a
covalently crosslinked silicon-based organic polymer core (e.g., a
polysiloxane such as PDMS), the core polymer does not need to be
covalently crosslinked to act as a reinforcing fiber. Thus in other
embodiments, thermoplastic polymers such as polyurethanes, PLGA,
PCL, nylon, polystyrene, acrylic polymers, polypropylene,
polyethylene and fluoropolymers, among others, can be used as the
core reinforcing fiber.
[0101] Polyurethanes represent a broad class of polymers having a
wide range of properties and, as such, can serve as core and/or
sheath materials in conjunction with the present disclosure. For
example, a thermoplastic polyurethane core may be at least
partially enclosed in a hydrophilic or hydrogel polyurethane
sheath. Many polyurethane materials exhibit physical cross-linking
and thus do not require a separate crosslinking step. Such
materials may be used, for example, in conjunction with melt-based
or solvent-based spinning processes, among others.
[0102] The present inventors have demonstrated this concept in
conjunction with polyurethane chemistry by co-electrospinning a
hydrophilic polyurethane sheath around a more hydrophobic
polyurethane core as detailed in Example 4 below. The resulting
composite fiber has mechanical and hydration properties that differ
from either material alone.
[0103] More particularly, a composite material consisting of a
mechanically strong polyurethane core and a hydrophilic
polyurethane sheath has been created. The particular technique
employed was slit-surface, core-sheath electrospinning. As
previously noted, electrospinning creates fibers with small
diameters (micro or nanometers) which impart additional benefit and
functionality (e.g., softness, high surface area, conformability).
However, suitable fibers may also be produced using other
techniques including hot melt spinning, melt electrospinning, and
centrifugal fiber spinning, among other fiber forming
techniques.
[0104] As noted above, the pre-polymer of PDMS is difficult to
electrospin due to its low molecular weight and flowability, which
does not allow for sufficient polymer chain entanglement for fibers
to form. In addition, the silicone pre-polymer remains flowable
until it is crosslinked, so spinning fibers without some way to
preserve the fiber structure is unlikely to result in good fiber
formation. The present inventors have overcome this difficulty,
particularly for micro and nano-sized fibers, by using coaxial
electrospinning to encapsulate PDMS pre-polymer and a cross-linking
agent within a polymer sheath. In certain embodiments a hydrogel
polymer is used as a polymer sheath material. For instance, in
Example 3 below, the core polymer is crosslinked PDMS and the
polymer sheath is a crosslinked polyvinylpyrrolidone (PVP).
[0105] In some embodiments, the cross-linking of the
hydrogel-forming polymer is modified to suit the core materials, as
well as the desired properties of the composite fiber. In some
embodiments, hydrogel crosslinking is initiated by the application
of heat, along with core crosslinking. For example the core polymer
may be crosslinked PDMS and the polymer sheath may be a crosslinked
polyvinylpyrrolidone (PVP), both of which are crosslinked by the
application of heat (see, e.g., Example 3). In other embodiments,
methods to initiate cross-linking of the hydrogel polymer (and/or
core polymer) could include UV or gamma radiation, freeze/thaw
cycles, supercritical drying, and so forth. In still other
embodiments, a physically crosslinked hydrogel is selected (see,
e.g., Example 4). All of these variations of hydrogel chemistries
are within the present disclosure.
[0106] A major benefit of this aspect of the present disclosure is
that an elastic, durable, biocompatible and mechanically stable
construct may be provided for hydrogels so that the many potential
benefits of hydrogels can be utilized in applications which require
greater mechanical integrity. Another benefit is that methods of
forming core-hydrogel fibers are provided, which do not require a
separate crosslinking step, due to the physical crosslinking
attributes of the polymers selected as the core-forming polymer
and/or sheath-forming polymer.
[0107] As previously noted, small diameter fiber meshes provide,
inter alia, higher surface area, higher permeabilities and lower
pore sizes than meshes made from larger diameter fibers. This
disclosure thus provides materials which combine the benefits of
hydrogels and small-diameter fiber meshes.
[0108] As elsewhere wherein, these core-hydrogel fibers can be
gathered, formed or processed into various shapes (e.g., tube,
mesh) for use as medical devices or other products.
[0109] Other materials may also be incorporated into the core or
sheath polymer to modify or obtain new properties. For example,
water absorbing particles may be included to further improve water
retention capabilities or agents which will elute out to provide
another benefit.
[0110] Thus, in some embodiments, excipient materials are
incorporated into the fibers to increase water swelling and
retention capacities. Excipient materials include cross-linked
hydrophilic polymers such as PVP, cellulose, gelatin and starch,
among others. These materials can be incorporated as dissolved
polymers in the sheath or core during electrospinning.
Alternatively, they may be included as particulates that are not
soluble or are only partially soluble in the solvents used to
produce the fibers. In this case, the excipient materials will
present as particles embedded in or projecting from the surface of
the finished fibers.
[0111] In some embodiments, therapeutic agents such as small
molecule drugs, anesthetics, procoagulants, anticoagulants,
antimicrobials, biologics, RNAi, genetic material, genetic vectors,
vaccines, or particles such as silver nanoparticles are
incorporated into the fibers which are released upon hydration.
[0112] With regard to applications, in some embodiments, the
composite core-hydrogel fiber can be used in heart valve leaflets.
Replacement heart valves, in some cases, use synthetic materials to
recreate the native leaflets. Native leaflets are thin, highly
flexible and durable. In addition to these properties the leaflets
need to be non-thrombogenic. Encouraging endothelialization is one
of the best ways to provide non-thrombogenic implants. The hydrogel
layer sheath along with the microfiber or nanofiber architecture
will encourage endothelial cell growth. Upon hydration, the
hydrogel layer will swell and fill the pores between the core
fibers--thus preventing blood from passing through the pores of the
valve. In a preferred embodiment, microfiber meshes of
core-hydrogel fibers are electrospun to a thickness of between 100
to 1000 microns. Target fiber diameters are between 500 nm to 10
um. These meshes are then cut into appropriate shapes and attached
to a main body which will be implanted via open or
minimally-invasive surgery. Alternate embodiments include:
functionalizing the core polymer with proteins or antibodies (e.g.
CD34, VEGF) to encourage tissue ingrowth and reendothelialization
(particularly where a degradable hydrogel is selected);
electrospinning onto a frame (e.g. polymer fiber, metal wire,
contoured conductive mesh) to help shape the leaflet and/or provide
an attachment to the main body; electrospinning onto a
biocompatible fiber structure which will create a composite implant
(e.g. fibers provide additional mechanical strength or varying
stiffness across the leaflet); and coating or functionalizing the
fibers to decrease thrombogenicity (e.g. heparin).
[0113] In some embodiments, the composite core-hydrogel fiber can
be used in stent graft coverings. For example, hydrogel fibers can
be used as coverings on stents that are used in left atrial
appendage closures. These embodiments are similar to the heart
valve leaflet, but the core-hydrogel fibers are electrospun onto a
tubular collector to form a tube of microfibers or nanofibers.
Preferred mesh thickness is between 100 and 1000 microns. Target
fiber diameters are between 500 nm to 10 um. This tube can then be
attached to a stent to form the stent graft. Alternatively, the
fibers may be electrospun directly onto the stent. Alternate
embodiments described for the heart valve concept are applicable
here as well. Advantages are that composite core-hydrogel fibers
will encourage cellular ingrowth while providing an elastic,
biocompatible, durable implant.
[0114] In some embodiments, the composite core-hydrogel fiber can
be used in vascular grafts. These embodiments are similar to the
stent graft design but the tube is not attached to a stent and the
preferred mesh thickness range is larger (100 to 5000 microns).
Alternatively, these tubular meshes act as a reinforcing cuff for
vessels (e.g., vascular autografts for bypass surgeries) or other
tubular structures where the mechanical properties of the native
tissue have deteriorated, such as in abdominal aortic
aneurysms.
[0115] In some embodiments, the composite core-hydrogel fiber can
be used in bioengineered blood vessels. These embodiments are
similar to the vascular graft above, and the core-hydrogel
microfiber or nanofiber mesh can be fashioned into a tube and
seeded with cells ex vivo. These cells, typically fibroblasts,
smooth muscle cells and endothelial cells, are incubated under
various conditions (e.g. pulsatile flow, steady flow, no flow) in
nutrient-rich environments to grow tissue on the graft material.
The core-hydrogel microfibers or nanofibers may provide advantages
in encouraging cell infiltration and growth as well as provide an
elastic character typical of blood vessels. The core-hydrogel tube
may be used alone or in combination with other natural (e.g.
collagen) or synthetic (e.g. PTFE, ePTFE, polyurethane) materials.
In other embodiments, the graft is seeded with cells and implanted
without significant incubation or implanted without cell seeding.
In the latter case, cells from the host will infiltrate and
populate the graft.
[0116] In some embodiments, the hydrogel fibers are used in medical
device sealing applications. These mechanically robust, hydrogel
fibers and resulting meshes, yarns, tubes, etc. are ideally suited
for use to seal interfaces between medical devices and the body,
other medical devices or other surfaces requiring a seal. For
example, they can be used to provide a seal between an implanted
heart valve and the native valve annulus to prevent paravalvular
leakage.
[0117] In one embodiment, the hydrogel fibers are electrospun
directly onto the outer surface of the valve stent or fashioned
into a mesh, yarn or tube and applied to the valve stent as part of
the manufacturing process. Upon implantation the hydrogel absorbs
water from the blood which leads to swelling, filing of the space
between the implant and the valve annulus and thus sealing around
the valve to prevent leakage. The advantage compared to other
hydrogels is the favorable mechanical properties and durability
lead to a safer and more effective product. Other applications
include: providing hydrogel microfibers or nanofibers on the vessel
contacting side of a stent graft, vascular graft or other medical
device to seal between the graft or other medical device and the
vessel wall; providing hydrogel microfibers or nanofibers on the
outer or inner diameter of a stent graft to seal between two stent
graft components which will be assembled together (e.g., EVAR graft
main body and iliac limb extension); providing fibers on the
outside of a stent graft to be used as a chimney, snorkel, etc. as
part of another stent graft placement; providing hydrogel
microfibers or nanofibers on the outer surface of a transcutaneous
catheter, ostomy bag, or wire lead to seal between the device and
the skin and/or underlying muscle, fat or fascia; providing fibers
on the outside of a device designed for implantation into the
digestive track to prevent food contact with a segment of the
digestive system; providing fibers around an endoscopic or
laparoscopic instruments or access tubes to provide a temporary
seal with the patient's tissues to prevent bleeding, gas leakage or
fluid leakage. For those applications where the device is temporary
and will be removed the robust mechanical properties and slippery
surface of the hydrogel will aid in removal.
[0118] In some embodiments, the hydrogel fibers can be manufactured
such that they hydrate only when a strain is applied (see, e.g.,
Example 3 below). Upon hydration, the fibrous construct increases
in volume. This property can be applied to create strain-dependent
seals around stent grafts and heart valve cuffs. In some cases,
when stent grafts and heart valve cuffs are deployed, they do not
make complete conformal contact with the vessel wall or annulus,
thereby leaving open spaces between the stent graft and vessel,
which in turn may lead to leaks, device failure and poor clinical
outcomes. The hydrogel fibers can be used as a ring or stent
covering such that during delivery, the hydrogel fiber covering
does not wet, but upon stent deployment the fiber covering is
strained, resulting in wetting and swelling of the fibers that fill
empty spaces where the stent does not make conformal contact with
surrounding tissues.
[0119] In some embodiments, the hydrogel fibers are used in
non-medical sealing. For instance, the core-hydrogel fibers will be
useful in providing a seal in non-medical applications in aqueous
or non-aqueous environments. For example, in aqueous environments,
fibers positioned between two surfaces to be sealed will hydrate
upon contact with water then the swelling will seal the surfaces
and prevent flow through the microstructure. In non-aqueous
applications (e.g., oil transport), the mesh will be hydrated upon
installation creating a seal from swelling in between two surfaces
and also prevent leakage due to immiscibility with the non-aqueous
fluid.
[0120] In some embodiments, the composite core-hydrogel fiber can
be used in arteriovenous grafts or shunts. These grafts are used in
hemodialysis patients to provide better needle access for repeated
dialysis. A core-hydrogel microfiber or nanofiber mesh will provide
a robust set of mechanical properties as well as encourage cellular
ingrowth. The elasticity, durability, potential for
biocompatibility and low thrombogenicity will improve the
performance of these grafts. In one embodiment, a core-hydrogel
microfiber or nanofiber mesh is fashioned into a tube and
implanted. This tube can be pre-treated by functionalization or
coating with other materials (e.g. heparin, collagen, gelatin,
growth factors) to improve integration and cell ingrowth. In other
embodiments, a core-hydrogel mesh can be combined with other
natural or synthetic materials as sheets or meshes to form a
composite, layered structure. This layered structure may improve
the mechanical properties, the ability to contain blood immediately
after implantation or long term durability or performance.
[0121] In some embodiments, the composite core-hydrogel fiber can
be used in hernia meshes. Core-hydrogel fibers (e.g., silicone or
polyurethane core with a hydrogel sheath) can be electrospun into
flat mesh configuration of thickness 500 to 5000 microns. To
improve mechanical properties, a composite may be formed with
biocompatible polymer fibers by electrospinning directly onto those
fibers in the desired configuration. These fibers may also be
provided in a configuration that improves suture-ability of the
mesh. In alternate embodiments, the mesh may be functionalized to
improve tissue ingrowth or integration.
[0122] In some embodiments, the composite core-hydrogel fibers can
be used in dural coverings. In neurosurgical procedures where the
dura are compromised, it is desirable to provide a covering to
re-seal the membrane. For example, a core-hydrogel microfiber or
nanofiber mesh, optionally combined with a polymer membrane (e.g.
silicone, PLGA, collagen) can be used for this purpose.
[0123] In some embodiments, the composite core-hydrogel fibers can
be used in wound dressing. Challenges for wound dressings include
adherence to the wound, wound exudate management and permeability
to air and water (wound exudates). For example, hydrogel
encapsulated polymer (e.g., silicone or polyurethane) may be
electrospun into a mesh between 100 and 5000 microns. Preferred
fiber diameters are between 500 nm and 10 microns. The advantage of
the reinforced hydrogel is that it provides moisture to the wound
bed while also forming a protective layer which does not adhere to
the wound. In one embodiment, a hydrogel-polymer dressing is
supplied separately and medical staff place additional gauze or
other bandages in layers on top of the core-hydrogel fiber
dressing. In another embodiment, a core-hydrogel mesh is combined
with a gauze or other backing material as part of the finished
product to aid in the absorption of fluid and protect the wound. In
still other embodiments, a core-hydrogel mesh can be fabricated
with therapeutic agents such as antibiotics, antifungals, topical
pain relievers, disinfectants (e.g. iodine) or the like. In still
other embodiments, a hydrogel-polymer mesh is fabricated for use
with negative pressure wound therapy. In this case, the mesh is
sized to be compatible with these devices and is placed on the
wound bed as negative pressure is applied. The high permeability
and porosity allow exudates removal as well as a non-adherent
dressing when it must be removed. The hydrogel sheath or core
polymer may also be useful in controlling release of therapeutic
agents to the wound (e.g., antimicrobials, antibiotics, silver
ions, growth factors, analgesics, anesthetics, debridement
compounds or enzymes, etc.).
[0124] In some embodiments, the composite core-hydrogel fiber can
be used in hemostat applications. For hemostatic applications, the
device is configured much like the wound dressings, but the
hydrogel-polymer microfibers or nanofibers are fabricated or
surface modified with a prothrombotic agent (e.g. thrombin, kaolin,
chitosan, fibrin). The fiber provides a nonadherent dressing that
can be removed easily. In addition, the high permeability and
porosity allows the blood to penetrate and contact a large amount
of the surface area with the prothrombotic agent. This open
structure also allows for coagulation factor diffusion back into
the wound promoting clot formation. This material may also be
integrated as a non-adherent layer on other dressings (e.g. Combat
Gauze); for this application the fibers may or may not be
manufactured with a prothrombotic agent.
[0125] In some embodiments, the composite core-hydrogel fiber can
be used in filtration. Composite core-hydrogel fiber meshes can be
used as filters or as part of a filter for air, gases, liquids,
slurries or particles. In particular, the low pore size and high
permeability of electrospun, microfiber or nanofiber meshes are
desirable for filters. In addition, where the fibers are
elastomeric, the elastomeric nature provides a way to clean the
filter. Simply, stretching the material biaxially,
circumferentially or otherwise will increase the pore size. Then,
backflow of gas or liquid will provide a method to clear the pores
of debris or other material. In a similar manner, cake which forms
on the intake side of a filter can be easily removed by stretching
the fiber mesh allowing the cake to fall off. The core-hydrogel
microfiber or nanofiber mesh may be characterized by high strength
and hydrophilicity, thus being useful as a filter, barrier or
separating membrane to partition oil content in water. The
core-hydrogel microfiber or nanofiber mesh can be used alone
(preferred thickness of 100 microns to 1 cm). Alternately, the
core-hydrogel mesh can be constructed as part of a layered filter
using other commonly available filter materials. In this case, the
core-hydrogel may be electrospun directly onto another material,
placed on the other material during assembly or electrospun onto a
wire or other fiber mesh with large openings to provide mechanical
support.
[0126] In some embodiments, the composite core-hydrogel fiber can
be used in drug delivery. The hydrated core-hydrogel composite
material may act as a substantially non-porous yet conformal layer.
In one embodiment the core-hydrogel material would be inserted into
the target delivery area then inflated with gas or other fluid
(e.g., a drug containing solution, etc.) to conform to the internal
structure of the target area. Direct, conformal contact of the
hydrogel with the surface leads to efficient drug delivery.
Alternatively, upon reaching a certain expansion limit on
inflation, the pores become stretched and open to allow drug
solution to be released. Once deflated, the pores seal back up thus
inhibiting drug delivery to areas not being targeted during removal
of the device. This approach is particularly applicable for
therapeutic delivery to cavities and lumens, such as the sinusoidal
space.
[0127] In various embodiments, drugs may be incorporated into the
core-hydrogel microfibers or nanofibers for delivery to a patient.
For example, a core-hydrogel fiber mesh may be formed with drug in
the core-forming solution, and placed on the skin for cutaneous or
transcutaneous delivery. Fiber meshes of the present disclosure are
beneficial in that they provide a means of targeted delivery to
difficult orifices such as sinus cavities, intestinal wall or ear
canals due to the ability to balloon open for conformal delivery. A
tubular or other shaped mesh may also be implanted to provide
sustained drug delivery. It may be implanted alone or held in place
using another medical device, such as a stent.
[0128] In some embodiments, the composite core-hydrogel fibers can
be collected into aligned fiber bundles like a yarn. These yarns
will act as strong, elastic hydrogel fibers that can be used (e.g.,
sutures) or processed further, including: twisting multiple yarns
together into a rope, weaving multiple yarns together into a woven
sheet, tube or other shape, braiding multiple yarns together into a
stent, scaffold or other tubular structure. These configurations
can be developed into novel medical devices such as hydrogel
catheters, introducer sheaths, guide wires, vascular grafts, hernia
meshes, etc.
[0129] In some embodiments, the composite core-hydrogel fiber can
be used in textiles. Core-hydrogel microfibers or nanofibers may
also be used in textile applications where high elasticity,
durability, water absorption and permeability are desired.
[0130] In some embodiments, the composite core-hydrogel fiber can
be used in tissue engineering applications. Hydrogels allow for
free diffusion of oxygen, nutrients, etc., which is desirable for
these purposes. This property is further enhanced, because
diffusion not only can occur across the hydrogel bulk, but through
the porosity created by the fibrous network. Hydrogels are used
extensively in tissue engineering applications due to their
promising biocompatibility and hydration properties. A major
benefit of the present disclosure is that fibrous hydrogels would
allow for better cell attachment and integration to form 3D
scaffolds. The hydrogel sheath would allow for cell attachment and
in-growth, which could eventually degrade away, while the core
polymer fibers would provide more permanent mechanical support. A
specific example application of this includes hyaline cartilage
repair, in which the hydrogel sheath provides a biocompatible
scaffold for stem cells to attach and differentiate into
chondrocytes while the porosity provides space for chondrocytic
secretion of collagen and ECM components.
[0131] In some embodiments, the hydrogel fibers are used as a
tissue bulking agent in cosmetic or plastic surgery. The elastic
and flexible mechanical properties and high hydration of the
hydrogel fibers can be tailored to match that of native tissue for
a more natural look and feel. The fibrous nature will integrate
with the surrounding tissue such that the bulking agent stays in
place and will not become displaced. Furthermore, the hydrogel can
be made to be nonbioresorbable and therefore maintain its bulking
capacity over time.
[0132] In some embodiments, the composite core-hydrogel fiber are
used as medical electrodes. The swelling properties of hydrogel
allow for conformable and intimate contact with tissue that can
lower electrical impedance and improve electrode performance.
Furthermore, to improve electrical conductance, the core material
can be comprised of a conductive polymer or include electrically
conductive particles or ions.
[0133] The ballooning and hydration capability of the composite
core-hydrogel fibers is a unique property that can be used for the
ablation of tissues through the use of microwaves. For example, for
ablation within a body cavity (e.g., endometrial, left atrial
appendage) or to an irregular surface (e.g., liver, esophagus,
sinuses) a mesh of composite core-hydrogel fibers can be inflated
with a gas (e.g., carbon dioxide) to make conformal contact with
the tissue. Application of microwaves from a source within the
balloon will heat the water within the hydrogel membrane, which is
in intimate and conformal contact with the cavity or tissue
surface, to thermally ablate the surrounding tissue.
[0134] This same technique may be extended to other ablation
approaches, including hydrothermal (e.g., inflate the balloon with
hot water or other hot liquid), chemical (e.g., ablative agent in
the hydrogel fibers) or cyroablation (e.g., cold source or liquid
nitrogen used to chill the balloon).
[0135] The composite core-hydrogel fibers of the present disclosure
may also be used to embolize a body lumen. The composite structure
provides a fiber or coil that can be inserted into a patient using
techniques know to those skilled in the art. The hydrogel
properties then swell the fibers to completely fill the body lumen
or aneurysm cavity. Two key advantages here are 1) combination of
fiber strength and high swelling ratio, and 2) ability to form very
small fibers or coils and/or flexible implants.
Example 1
Fibers with PDMS Core and PLGA Sheath
[0136] Core/sheath fibers are fabricated in accordance using a
high-throughput core-sheath needleless electrospinning fixture. The
sheath polymer system was a 3.5 wt % 85/15 poly(L-lactic
acid-co-glycolic acid) (PLGA) in 6:1(by vol) chloroform:methanol
solvent. The core polymer consisted of PDMS (Sylgard 184, available
from Dow Corning, a two-part liquid system consisting of part A
(pre-polymer) and part B (cross-linking agent)) mixed in a 10:1
mass ratio. The sheath solution flow rate was set to 200 ml/h while
the core flow rate was set to 20 ml/h. The fibers were deposited
onto and collected from a grounded collection plate. The fabricated
mesh was then placed in an oven at 100.degree. C. (to accelerate
curing) for 3 hours and then immersed in chloroform for 1 hour to
dissolve the PLGA sheath. The PDMS fiber mesh swelled to an extent
upon exposure to the solvent, but then shrank back to original size
after solvent evaporation. FIG. 1 shows an image of the
cross-section of the PLGA/PDMS sheath/core fibers after curing. The
different polymers in the sheath/core configuration can be
observed. FIG. 2 shows the PDMS fibers after sheath layer removal.
PDMS fibers were manufactured to be between about 1 and 5 microns
in diameter. As described elsewhere herein, however, the diameter
of the core PDMS can be tuned by modulating electrospinning
parameters.
Example 2
Further Fibers with PDMS Core and PLGA Sheath
[0137] Core-sheath fibers were electrospun with 50/50
poly(D,L-lactic acid-co-glycolic acid) (5050 PDLGA) as the sheath
over a PDMS (Sylgard 184) core, as described in Example 1. The
sheath solution was an 11 wt % 5050 PDLGA in hexafluofoisopropanol
(HFIP). The flow rate for the sheath solution was set at 10 ml/h
while the core solution flow rate was set at 1 ml/h. The fibers
were subsequently placed in a 60.degree. C. oven for 24 hours to
allow the PDMS in the fibers to cure. FIG. 3A shows a core-sheath
structure (in cross-section) that was formed. Diameters of the
fibers were measured for both top-down and cross-sectional images.
The overall fiber diameter of the fibers was approximately 7
microns (see FIGS. 3A and 3C), while the core PDMS diameter was
approximately 4.5 micron (see FIGS. 3B and 3D). The 5050 PDLGA
sheath was removed under accelerated degradation conditions by
immersing the mesh in 12 pH buffer consisting of 1.5% sodium
phosphate, 0.1% boric acid, and 0.08% citric acid at 37.degree. C.
for 7 days. As can be seen in FIG. 3B, the sheath layer was
completely degraded and removed, leaving behind PDMS-only
fibers.
[0138] The electrospun fibers and meshes of the present disclosure
offer different properties than those formed from traditional
methods of constructing PDMS as a cast film. The contact angle of
the electrospun PDMS-only mesh was measured to be 110.degree. while
a cast film of PDMS had a contact angle of 104.degree.. FIG. 4
shows the hydrophobic and oleophilic nature of the PDMS mesh formed
using the electrospinning processes of the present disclosure. A
water droplet (left) placed on the mesh remains beaded while an oil
droplet (right) wets the mesh and can move throughout the porosity
of the mesh. FIG. 5 shows the mechanical properties of the mesh
compared to a cast PDMS film. The data indicates that the PDMS
fiber mesh exhibits significantly different mechanical properties
than a cast film. The modulus of the mesh is significantly lower
(0.2 MPa vs 2.0 MPa) while its extension at max loading is
significantly higher (300% vs. 122%) relative to the cast PDMS
film.
[0139] FIGS. 6A-6C show cross-sectional photomicrographs of
electrospun fibers of the present disclosure having PDMS in the
core and 5050 PDLGA in the sheath. The electrospinning process was
carried out at sheath:core flow rates of 10:1, 10:0.25, and 20:0.25
ml/h in order to generate PDMS fibers with different fiber
diameters, as shown in the cross-sectional images of FIGS.
6A-6C.
Example 3
Fibers with PDMS Core and PVP Sheath
[0140] Core/sheath fibers were fabricated using a sheath polymer
solution of 8 wt % PVP (polyvinyl pyrrolidone) in TFE
(trifluoroethanol), while the core polymer solution consisted of
Sylgard 184, a two-part liquid system consisting of Part A
(pre-polymer) and B (cross-linking agent) mixed in a 10:1 mass
ratio. The sheath flow rate was set to 10 mL/h while the core flow
rate was set to 2 mL/h.
[0141] Meshes were collected on PTFE coated aluminum shims and then
cured at either 100.degree. C. or 150.degree. C. for 24 hours. The
meshes were then removed from the aluminum shims and submerged in
deionized water in which any non-crosslinked PVP was solubilized by
the water. The remaining PVP was crosslinked as a robust sheath
around the silicone fiber cores, which then formed a hydrogel and
swelled to >200% its initial mass, the amount of swelling is
proportional to the degree of crosslinking of the PVP (and thus the
temperature of the cure). In particular, for the 100.degree. C.
sample, swelling (by mass) was measured at 242%.+-.48%, whereas for
the 150.degree. C. sample, swelling (by mass) was measured at
401%.+-.76%.
[0142] Gel fraction data (% hydrogel) were generated. For the
100.degree. C. sample, the gel fraction was measured at 64%.+-.1%,
whereas for the 150.degree. C. sample, the gel fraction was
measured at 98%.+-.3%. These data indicate that upon water
extraction of non-cross-linked PVP, the 100.degree. C. cured sample
loses .about.40% of its mass while the 150.degree. C. sample
maintains almost 100% of its mass. This suggests that the PVP
sheath is nearly completely cross-linked at 150.degree. C. and may
be only partially cross-linked at 100.degree. C.
[0143] Similar conclusions can be drawn by cross-sectional analysis
with SEM, as illustrated in FIG. 7, which shows: (A) SEM
cross-section of core-sheath fibers where the core consists of
fully cured PDMS and the sheath is PVP cured at 100.degree. C.; (B)
SEM cross-section of the same fibers in (A) except after they have
undergone water extraction to remove non-cross linked PVP; (C) SEM
cross-section of core-sheath fibers where the core consists of
fully cured PDMS and the sheath is PVP cured at 150.degree. C.; (D)
SEM cross-section of the same fibers in (C) except after they have
undergone water extraction to remove non-cross linked PVP. Before
hydration, the two cure temperature samples look identical in core
fiber diameter (around 6 um) and in sheath thickness (around 1 um).
After hydration and subsequent drying, however, the sheath appears
to be almost completely removed in the 100.degree. C. sample while
it remains intact in the 150.degree. C. sample.
[0144] Analysis by FTIR can indicate the presence of PVP in the
sample by the existence of an amine peak around 1650 cm.sup.-1.
Additionally, PDMS does not absorb water so the presence of a broad
peak around 3400 cm.sup.-1 indicates an O--H bond and therefore the
absorption of water by the sample, which would only occur if
cross-linked PVP is present. FIGS. 8 and 9 show the spectra
obtained from the silicone fiber hydrogels in the wet and dry
states as compared to pure PDMS and pure PVP cured at temperatures
of 100.degree. C. and 150.degree. C., respectively.
[0145] In comparing the spectra for PVP-PDMS cured at 100.degree.
C. to pure PDMS and pure PVP it can be confirmed that very little
PVP remains in the sample after the initial water extraction. What
little PVP that remains can only be detected when the sample is in
the hydrated state. The dry PVP-PDMS hydrogel matches nearly
perfectly with pure PDMS and absorbs essentially no water from the
atmosphere. This supports the mass loss data and observations from
the SEM that an essentially undetectable amount of PVP remains on
the sample although it still behaves as a hydrogel.
[0146] Contrary to the spectra for the PVP-PDMS hydrogels cured at
100.degree. C., the spectra for samples cured at 150.degree. C.
show an amine peak and absorbed water even when dry. This is
further evidence to support that nearly all of the PVP is
cross-linked at this higher temperature and remains in the sample
after initial water extraction.
[0147] The mechanical properties of the fiber hydrogel are
dramatically enhanced due to the presence of a silicone microfiber
structure. Tensile properties of hydrogels are rarely reported and
difficult to find due to the poor mechanical stability of the same.
With silicone fiber reinforcement, the hydrogel has a larger
surface area for wetting while also maintaining mechanical
integrity and strength. Additionally, the presence of a
cross-linked hydrogel layer on the silicone provides another layer
of support for the silicone fibers and increases the overall
strength of the composite material. FIG. 10 shows the comparison of
different formulations of hydrogel-silicone microfiber composites
both in the wet and dry states. The samples were cut using a 20 mm
die and tested on an Instron.RTM. system at a rate of 50 mm/min. It
can be seen from this data that temperature of cross-linking can
affect the tensile strength and modulus of the material.
[0148] This data also supports the previously stated conclusion
that very little cross-linked PVP remains on the sample cured at
100.degree. C. given that the mechanical properties appear to be
unaffected by the hydration state of the material. Conversely, the
PVP-PDMS cured at 150.degree. C. behaves drastically different when
it is dry than when hydrated. In the dried state the material has a
modulus .about.200 times greater than when it is hydrated (i.e., 75
MPa vs. 0.4 MPa). It also has a much shorter elongation to break
(7% vs. 140%). When the material is in the hydrated state the PVP
swells and the mechanical properties are driven largely by the PDMS
fibers.
[0149] Another feature of silicone core/hydrogel sheath fibers is
that its mechanical features can change, depending on whether the
fibers are wet or dry. For example, a dry mesh will have high air
permeability, high porosity and will be opaque. Conversely a
hydrated mesh will have lower air permeability (because the swollen
hydrogel fills the pores), high water permeability and will be
optically clear. Upon reaching its expansion limit of the mesh the
pores open up and the gas or liquid flows through the mesh as
opposed to bursting it. This ability to expand is also affected by
the cure temperature, because as the elongation of the fibers is
dependent on cure temperature.
[0150] As shown in FIG. 11A, a "balloon" was formed from a mesh of
hydrated PVP-PDMS fibers cured at 100.degree. C. If spherical
expansion is assumed, then the volume expansion ratio is near 800%
when the balloon reaches maximum expansion (see FIG. 11B). At the
maximum expansion, the balloon doesn't burst but merely becomes air
permeable and allows air to escape through the pores. The air
permeability and porosity of the hydrated mesh can be increased
upon stretching the mesh to open up the pores. Due to the decreased
permeability of the hydrated hydrogel mesh, the material can hold
air or water and expand to very high volumes while still
maintaining mechanical integrity. In addition, the microfibers
provide a flexible balloon that can conform to irregular surfaces,
cavities or containers. Compared to a pure PDMS fiber mesh, which
expands only about 100% before bursting, the effect of the very low
amount of cross-linked PVP on the surface is quite significant. The
PVP-PDMS hydrogel cured at 150.degree. C. expands to 450% before
becoming air permeable, thus indicating that cure temperature and
therefore amount of cross-linked PVP in the sample effects this
property.
[0151] Due to the unique pairing of a hydrophilic hydrogel sheath
with an oleophilic core fiber, the PVP-PDMS fiber mesh swells in
both water and oil. Table 1 shows a comparison of the swelling
properties of the PVP-PDMS meshes in DI water and Vacuum Pump Oil
at different curing temperatures. The 100.degree. C. cured mesh
absorbs nearly the same amount of oil as it does water due to the
presence of the PDMS fibers and the very small amount of PVP on the
surface. The mesh cured at 150.degree. C., on the other hand,
absorbs much more water than it does oil, because much more
crosslinked PVP is present in this sample.
TABLE-US-00001 TABLE 1 Cure Temperature 100.degree. C. 150.degree.
C. Water Swelling (%) 242% 401% Oil Swelling (%) 215% 150%
[0152] In addition to the swelling, ballooning and mechanical
strength differences between samples cured at 100.degree. C. and
those cured at 150.degree. C., these samples also show a distinct
difference in their wettability after the initial hydration
(extraction) and drying. The PVP-PDMS fibers cured at 150.degree.
C. quickly absorb water and become fully hydrated without any
additional handling or manipulation. Meshes cured at 100.degree.
C., on the other hand, are more hydrophobic in their dry,
unstretched state. In order to hydrate the meshes, they are
manipulated (e.g., stretched). In this regard, when water is
initially applied to the dry mesh, it beads on the surface.
However, as the sample is stretched and manipulated, it eventually
becomes fully hydrated.
Example 4
Fibers with a Polyurethane Core and a Hydrophilic Polyurethane
Sheath
[0153] In this example, slit-surface, core-sheath electrospinning
was employed, in which a hydrophilic aliphatic polyether-based
thermoplastic polyurethane (HLPU) was used as the sheath material,
while a mechanically stronger more hydrophobic aliphatic
polyether-based thermoplastic polyurethane material (HBPU) was used
as the core material. The electrospinning solutions were as
follows: 4 wt % HLPU in TFE and 6 wt % HBPU in HFIP.
Electrospinning was carried out at different sheath:core flow rate
ratios. At the flow rate ratios selected, the resulting fiber was
composed of HLPU and HBPU in the following HLPU:HBPU weight ratios:
(A) 93:7, (B) 82:18, (C) 60:40, and (D) 38:62, respectively.
[0154] FIG. 12 shows the SEM of the fibers for each composition;
fiber diameters for all formulations were approximately 2
microns
[0155] Characterization of the meshes included dimensional, and
hydration measurements, which are summarized in Tables 2 and 3
below. Mechanical characterization was determined by cutting the
meshes into dog-bone shapes and performing tensile testing using an
Instron.RTM. at a pull rate of 50 mm per minute. Swelling was
characterized by immersing samples in phosphate buffered saline
(PBS) for at least 20 minutes and the PBS was allowed to drip off
before weighing. Swelling was calculated as the (wet weight-dry
weight)/dry weight. PBS retention was determined by placing the
hydrated material on filter paper and applying a weight equal to 40
mmHg for 30 seconds. The sample was then re-weighed to determine
the amount of water lost during testing. The wet tensile strength
of the different polyurethane samples are shown in Table 2 and
demonstrates an increase in mechanical properties as the amount of
HBPU in the fiber is increased. Therefore, by varying the core to
sheath material composition, one can modulate the tensile
strength.
TABLE-US-00002 TABLE 2 Formulation A Formulation B Formulation C
Formulation D HLPU:HBPU Ratio 93:7 82:18 60:40 38:62 Wet tensile
strength 0.20 .+-. 0.03 0.15 .+-. 0.02 0.26 .+-. 0.05 1.21 .+-.
0.08 (MPa)
[0156] Table 3 shows the hydration properties of the different
formulations and indicates that sample shrinkage upon hydration and
swelling were most impacted by the chemical composition of the
fibers. However, PBS retention did not appear to be significantly
impacted.
TABLE-US-00003 TABLE 3 Formulation A Formulation B Formulation C
Formulation D HPLU:HBPU Ratio 93:7 82:18 60:40 38:62 Basis weight
(GSM) 55 .+-. 4.5 .sup. 92 .+-. 0.5 110.8 .+-. 16.1 .sup. 94 .+-.
5.7 PBS absorption (%) 1750 .+-. 23 1760 .+-. 57 1270 .+-. 201 1110
.+-. 101 PBS retention (%) 52 .+-.2 .sup. 58 .+-. 1 56 .+-. 2 56
.+-. 1 Shrinkage (%) 57 .+-. 0.5 35 .+-. 2 14 .+-. 4 2 .+-. 6
[0157] A comparison of the mechanical and hydration properties as a
function of HLPU content is shown in FIGS. 13 and 14. FIG. 13 shows
that tensile strength increases as the amount of HLPU decreases
(and hence the amount of HBPU increases). However, as the tensile
strength increases, the amount of PBS absorption decreases as a
result of less hydrophilic material being present.
[0158] A comparison of the swelling (or PBS absorption) and
shrinkage data as a function of the HLPU content further reinforces
the utility of using a core-sheath fiber structure to modulate the
mechanical and hydration properties. As shown in FIG. 14, there is
an increase in swelling capacity as the HLPU content increases;
however, dimensional shrinkage (i.e., shrinkage in area) of the
mesh is also observed to increase as the HLPU content increases.
These data illustrates the formulation space for these materials
and shows a correlation between performance of the hydrogel mesh
and its chemical composition.
[0159] The performance across tensile strength, shrinkage, and
swelling has been optimized by varying the sheath to core ratio of
the polymeric materials. This is highly advantageous for numerous
applications, especially medical applications. For example,
hydrogel wound dressings are cut to fit the wound size when dry.
These dressings improve wound healing by providing a moist
environment and absorb excess wound exudate to prevent leakage.
However, excessive shrinkage may result in a dressing which
inadequately covers the wound after it starts to absorb liquid. As
shown in FIGS. 15, 16 and 17, in comparison with
commercially-available wound dressings such as Aquacel.RTM.
(ConvaTec Inc.) or Durafiber.RTM. (Smith & Nephew), a material
has been developed which provides equivalent water absorption (see
FIG. 15, Formulation A and B), much stronger mechanical properties
(see FIG. 16, all Formulations) and has minimal shrinkage (see FIG.
17, Formulations D) or shrinkage that is comparable to those
existing products (see FIG. 17, Formulations B and C).
[0160] In various embodiments, meshes in accordance with the
present disclosure are annealed at elevated temperature to improve
the properties of the same. For example, HLPU/HBPU sheath/core
fiber meshes as formed herein have been found to become less porous
upon annealing. In this regard, FIGS. 18A and 18B are
photomicrographs of a mesh formed from HLPU/HBPU sheath/core fibers
as described herein, before and after annealing, respectively.
Along with the reduction in mesh porosity, the annealing step is
accompanied by a reduction in mesh volume (and thus mesh area).
Unexpectedly, such an annealing step has been found to improve
water retention and to result in mesh expansion (rather than mesh
shrinkage). In this regard, FIG. 19 shows PBS retention values for
non-annealed (B Normal) and annealed (B Annealed) HLPU/HBPU
sheath/core fiber meshes in accordance with the present disclosure,
as well as retention values for Aquacel.RTM. and Durafiber.RTM.
wound dressings. As seen from FIG. 19, an annealed mesh material
has been developed which provides PBS retention equivalent to that
of Aquacel.RTM. and Durafiber.RTM. dressings. In this regard, FIG.
20 shows shrinkage or expansion values for non-annealed (B Normal)
and annealed (B Annealed) HLPU/HBPU sheath/core fiber meshes in
accordance with the present disclosure, as well as for Aquacel.RTM.
and Durafiber.RTM. wound dressings. Thus, as seen from the
foregoing, the present disclosure provides the ability to tailor
mesh absorption, retention and shrinkage/expansion to the
application at hand.
[0161] In addition, as noted elsewhere, the small fiber sizes
obtained also improves softness, conformability and leads to very
high surface areas. High surface area improves absorptive
capabilities, hydration kinetics and drug release capabilities,
among other properties. Moreover, the fibrous form factor allows
for formation/collection into novel form factors such as yarns,
ropes, tubes, meshes, etc.
Example 5
Fibers with a Polyurethane Core Containing Silver Particles and a
Hydrophilic Polyurethane Sheath
[0162] In this example, needle core-sheath electrospinning was
employed, in which a hydrophilic aliphatic polyether-based
thermoplastic polyurethane (HLPU) was used as the sheath material,
while a mechanically stronger more hydrophobic aliphatic
polyether-based thermoplastic polyurethane material (HBPU) was used
as the core material. The electrospinning solutions were as
follows: 4 wt % HLPU in TFE and 6 wt % HBPU in HFIP containing 30%
silver nanoparticles with respect to the polymer. The resulting
fibers exhibited a core-sheath geometry in which silver was
encapsulated and are shown in FIG. 21. Silver is well-known for its
antibacterial properties and such a mesh could be used for
sustained release of silver for wound dressing applications. In
addition to silver nanoparticles, other embodiments including
incorporation of other particles and/or excipients into the core
material to achieve different performance metrics. For example,
cross-linked celluloses or other hydrophilic polymers can be
incorporated into the core to further aid in the hydration
properties of the resulting fiber.
[0163] Although various aspects and embodiments are specifically
described herein, it will be appreciated that modifications and
variations of the present invention are covered by the above
teachings and are within the purview of the appended claims without
departing from the spirit and intended scope of the invention.
* * * * *