Note: Descriptions are shown in the official language in which they were submitted.
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CORE-SHEATH FIBERS
AND METHODS OF MAKING AND USING SAME
RELATED APPLICATIONS
[001] This application claims the benefit of U.S. Provisional Application
No. 61/852,224,
filed March 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 August 2,
2013, 2013, entitled "Biocomponent Elastomerie-Hydrogel Fibers," each of which
is
incorporated herein by reference in its entirety.
STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH
[002] 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.
TECHNICAL FIELD
[003] 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
[004] 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.
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SUMMARY OF THE INVENTION
[005] 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.
[006] 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.
[007] Still other aspects of the present invention pertain to meshes and
other articles that
are formed using the multicomponent fibers.
[008] 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
[009] FIG. 1 shows a photomicrograph of a cross-section of the PLGA/PDMS
sheathkore
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-3B 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.
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[0014] FIGS. 6A-6B 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. 7 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 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 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 fainted in accordance
with an
embodiment of the invention (cured at 100 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 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 C
and 150 C), when dry and when wet.
[0019] FIGS. 11A and 11B shows balloon formed from a hydrated PVP-PDMS
fiber mesh
cured at 100 C, in accordance with an embodiment of the invention, at two
levels of expansion.
[0020] FIG. 12 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.
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[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).
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[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 (um)
(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.
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[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.
[00371 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 x Mesh width x Mesh thickness =
Mesh area x
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 electrospirming, 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.
[00401 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
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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
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,
[ 0¨Si
R2
- , where R1 and R2 are organic radicals, for example, linear,
branched or cyclic
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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). 121 and R2
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).
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[00481 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.
[00491 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.
[00501 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 C, among
other crosslinking
techniques.
[00511 In certain embodiments, the polymeric sheath may be formed from
hydrophilic or
hydrogel materials, which are discussed in more detail below.
[00521 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.
[00531 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 polylactie
acid) including poly-L-
lactide, poly-D-lactide and poly-D,L-lactide, poly(lactide-co-glycolide)
(PLGA),
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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.
[00551 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.
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Fibers with Hydrogel Components and Components of varying
Hydrophilieity/Hydrophobieity
100561 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 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 (x 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
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] Hydro gels 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
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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 (ND!), 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 (HuMDI),
isophorone
diisocyanate (IPDI), cyclohexane diisocyanate (CHDI), 2,2,4-trimethy1-1,6-
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
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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
õ . . .
,
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
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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)aerylates (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-
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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.
[00711 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.
[00721 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.
[00731 In some embodiments, the sheath is etched away using an acid.
[00741 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.
[00751 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.
[00761 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
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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 foimed 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 faun a hydro gel 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
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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 (urn). Target fiber diameters are
between 500 nm to
urn. 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 urn. 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
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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 thxombogenicity 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.
[00861 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
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be functionalized, using the various methods described above, to improve
tissue ingrowth or
integration.
[00871 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.
[00881 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,
antifimgals, 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
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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.
100901 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.
100911 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).
100921 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
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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".
[00961 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.
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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.
101001 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.
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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-
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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.
[01061 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.
[01071 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.
[01081 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.
101091 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.
101101 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.
[01111 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.
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101121 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 urn. 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 urn 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
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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.
In one embodiment, the hydrogel fibers are electro spun 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
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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.
[0117] 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.
[0118] 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.
[0119] 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
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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.
[0120] 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.
[01211 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.
[0122] 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
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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.).
101231 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.
[0124] 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
peimeability 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-
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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.
[0125] 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.
[0126] 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.
[0127] 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.
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[0128] 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.
[0129] 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.
[0130] 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.
[0131] 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.
[0132] 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
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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.
[0133] 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).
[01341 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.
[01351 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 C (to accelerate curing) for
3 hours and then
immersed in chlorofolin 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.
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Example 2. Further fibers with PDMS core and PLGA sheath.
101361 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 hexalluoroisopropanol (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 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 C for 7
days. As can be seen in FIG. 3B, the sheath layer was completely degraded and
removed,
leaving behind PDMS-only fibers.
[01371 The electrospun fibers and meshes of the present disclosure offer
different properties
than those foitned from traditional methods of constructing PDMS as a cast
film. The contact
angle of the electrospun PDMS-only mesh was measured to be 110 while a cast
film of PDMS
had a contact angle of 104 . 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.
[01381 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
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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.
[0139] Core/sheath fibers were fabricated using a sheath polymer solution
of 8wt% 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 10mL/h
while the core flow
rate was set to 2mL/h.
[0140] Meshes were collected on PTFE coated aluminum shims and then cured
at either
100 C or 150 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 C sample, swelling (by mass) was measured at 242%
48%, whereas for
the 150 C sample, swelling (by mass) was measured at 401%+ 76%.
[0141] Gel fraction data (% hydrogel) were generated. For the 100 C sample,
the gel
fraction was measured at 64% 1%, whereas for the 150 C sample, the gel
fraction was
measured at 98% 3%. These data indicate that upon water extraction of non-
cross-linked PVP,
the 100 C cured sample loses ¨40% of its mass while the 150 C sample maintains
almost 100%
of its mass. This suggests that the PVP sheath is nearly completely cross-
linked at 150 C and
may be only partially cross-linked at 100 C.
[0142] 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 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 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 6um)
and in sheath
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thickness (around lum). After hydration and subsequent drying, however, the
sheath appears to
be almost completely removed in the 100 C sample while it remains intact in
the 150 C sample.
[01431 Analysis by FTIR can indicate the presence of PVP in the sample by
the existence of
an amine peak around 1650cm-1. Additionally, PDMS does not absorb water so the
presence of a
broad peak around 3400cm-1 indicates an 0-11 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 C and 150 C, respectively.
[01441 In comparing the spectra for PVP-PDMS cured at 100 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.
[0145] Contrary to the spectra for the PVP-PDMS hydrogels cured at 100 C,
the spectra for
samples cured at 150 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.
[01461 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. Figure 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 20mm die and tested on an Instron system at a rate of 50mm/min. It
can be seen from
this data that temperature of cross-linking can affect the tensile strength
and modulus of the
material.
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[01471 This data also supports the previously stated conclusion that very
little cross-linked
PVP remains on the sample cured at 100 C given that the mechanical properties
appear to be
unaffected by the hydration state of the material. Conversely, the PVP-PDMS
cured at 150 C
behaves drastically different when it is dry than when hydrated. In the dried
state the material has
a modulus ¨200 times greater than when it is hydrated (i.e., 75MPa vs.
0.4MPa). 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.
101481 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.
[0149] As shown in FIG11A, a "balloon" was formed from a mesh of hydrated
PVP-PDMS
fibers cured at 100 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 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.
[0150] 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
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curing temperatures. The 100 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 C, on the other hand, absorbs much more water than it
does oil, because
much more crosslinked PVP is present in this sample.
Table 1.
Cure Temperature 100 C 150 C
Water Swelling (%) 242% 401%
Oil Swelling CYO 215% 150%
[0151] In addition to the swelling, ballooning and mechanical strength
differences between
samples cured at 100 C and those cured at 150 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 C quickly absorb water and become fully hydrated without any
additional handling
or manipulation. Meshes cured at 100 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.
[0152] 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.
[0153] Figure 12 shows the SEM of the fibers for each composition; fiber
diameters for all
formulations were approximately 2 microns
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[01541 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 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 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
(IVIPa)
[0155] Table 3 shows the hydration properties of the different foimulations
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.
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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 92 0.5 110.8 16.1 94 5.7
PBS absorption (%) 1750 23 1760 57 1270 201 1110 101
PBS retention (%) 52 2 58 + 1 56 2 56 1
Shrinkage (%) 57 0.5 35 2 14 4 2 6
[0156] 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.
[0157] 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.
[0158] The perfoanance 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 (ConvaTee Inc.) or Durafiber (Smith 84 Nephew), a
material has
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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).
[0159] 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 and Durafiber wound dressings. As
seen from FIG.
19, an annealed mesh material has been developed which provides PBS retention
equivalent to
that of Aquacel and Durafiber 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 and
Durafiber 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.
[0160] 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.
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Example 5. Fibers with a polyurethane core containing silver particles and a
hydrophilic
polyurethane sheath.
101611 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
electro spinning
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.
101621 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.
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